personalized pathway-activated systems imaging in oncology: principal and instrumentation

Personalized Pathway-Activated Systems Imaging in Oncology

Tomio InoueDavid YangGang Huang Editors

Principal and Instrumentation

123

Personalized Pathway-Activated Systems Imaging in Oncology

Tomio Inoue • David Yang • Gang HuangEditors

Personalized Pathway- Activated Systems Imaging in OncologyPrincipal and Instrumentation

ISBN 978-981-10-3348-3 ISBN 978-981-10-3349-0 (eBook)DOI 10.1007/978-981-10-3349-0

Library of Congress Control Number: 2017931586

© Springer Nature Singapore Pte Ltd. 2017This work is subject to copyright. All rights are reserved by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed.The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use.The publisher, the authors and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, express or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations.

Printed on acid-free paper

This Springer imprint is published by Springer NatureThe registered company is Springer Nature Singapore Pte Ltd.The registered company address is: 152 Beach Road, #21-01/04 Gateway East, Singapore 189721, Singapore

EditorsTomio InoueSchool of MedicineYokohama City UniversityYokohama, Japan

Gang HuangRenji Hospital / Department of Nuclear

MedicineShanghai Jiao Tong UniversityShanghai, China

David YangVyripharm BiopharmaceuticalsUniversity of Texas Health Science CenterHouston, TX, USA

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Foreword

In the past 40 years, there has been an enormous increase in the range of imaging techniques available to investigate patients with cancer. Imaging modalities includ-ing computed tomography (CT), magnetic resonance imaging (MRI), ultrasound, optical imaging, and gamma scintigraphy have been used to diagnose cancer. Although CT and MRI provide considerable anatomic information about the loca-tion and the extent of tumors, they do not adequately provide optimal staging and differentiate residual or recurrent tumors from edema, radiation necrosis, or gliosis. Ultrasound images demonstrate local regional abnormalities, but they are operator- dependent. The optical imaging showed promising results; however, its ability to detect deep tissue penetration still has rooms to improve. Radionuclide imaging modalities (positron emission tomography, PET; single photon emission computed tomography, SPECT) are diagnostic cross-sectional imaging techniques that map the location and concentration of radionuclide-labeled compounds. Beyond know-ing precisely where a tumor is and its size, shape, and viability, PET and SPECT are making it possible to target the molecular makeup of the tumor and its metabolic activity. Whereas PET and SPECT can provide a very accurate picture of metaboli-cally active areas, its ability to show anatomic features is limited. As a result, new imaging modalities have begun to combine PET and SPECT images with CT scans used for treatment planning. PET-CT or SPECT-CT scanners combine anatomic and functional images taken during a single procedure, without having to reposition the patient between scans. PET-CT or SPECT-CT becomes a tool for image-guided pathway-directed precision therapy.

18F-fluorodeoxyglucose (FDG) has been used to diagnose and stage tumors, myocardial infarction, and neurological disease by PET. Although tumor metabolic imaging using 18F-FDG has been studied in the last two decades, its clinical practice is still hampered by factors such as differentiation of infection/tumor recurrence and low-grade/high-grade tumors. To improve the diagnosis, prognosis, planning, and monitoring of the cancer treatment, the characterization of tumor tissue is exten-sively determined by the development of more tumor-specific radiopharmaceuti-cals. The novel molecular imaging agent enables the comprehensive characterization of therapeutic intervention due to their specificity in molecular targets. Molecular

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imaging agents have potential application in patient selection, pharmacokinetic, dosage-finding, and proof-of-concept studies. To improve the diagnosis and prognosis and the planning, monitoring, and predicting of the cancer treatment, novel PET or SPECT imaging agents as well as higher-resolution PET-CT or SPECT-CT scanners are needed that would allow precise measurement of molecu-lar pathways on a whole-body image upon administration of a functional molecular imaging agent. These mechanism-based agents provide image-guided therapy that may discontinue ineffective treatment of diseases in the earlier phase and benefit patients. The development of a compounded single kit for clinical imaging was reviewed in this book. The kit formulation fulfills cGMP compliance for standard-ization and optimization of the drug product; meets quality assurance of sterility, pyrogenicity, and other analyses; and could shorten manufacturing time and reduce production cost by using generator-produced isotopes. This book seeks to bridge the gap from translational science to bedside.

Radiological Sciences E. Edmund KimThe University of California at IrvineIrvine, CA, USA

Foreword

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Preface

The development of molecular biology and genetics over the past 50 years has pro-vided medical science with unprecedented chances to study the molecular basis of diseases. Imaging becomes involved in the age of molecular medicine by creating new contrast media and radiopharmaceuticals. New classes of contrast agents based on tissue-specific uptake were developed. Molecular imaging science has been focused on imaging guidance in the areas of targeting epigenetic abnormalities and tumor microenvironment in overcoming resistance in cancers. The use of image- guided technologies to select patient for personalized therapy and to monitor thera-peutic outcomes is the focus of this book.

Nuclear and optical imaging agents could play major roles in the move from imaging of structure and morphology to the visualization of the individual biologic processes underlying disease and could contribute to more accurate diagnostics and improved treatment efficacy. In the chapter, the progress in the translation of mac-roscopic optical imaging and applications—including fluorescence imaging, FISH and optical tomography, and flow cytometry—was reviewed. The trends in nuclear imaging agent development were reviewed. Regulatory compliance of an imaging agent from bench to bedside was reviewed. The emergence of novel imaging agents in conjunction with specialized reconstruction algorithms can provide not only high-quality and high-resolution MRI or CT-like images but also functional infor-mation. This improvement in imaging technology allows quantitative target assess-ment of microenvironmental hemodynamics and subsequently in the imaging session.

The topics covered in this book are advances in molecular imaging both in radio-active and nonradioactive applications in preclinical drug discovery, drug develop-ment, regulatory compliance, and instrumentation for better management of cancer patients.

Yokohama, Japan Tomio Inoue Houston, TX, USA David YangShanghai, China Gang Huang

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Acknowledgments

We are very appreciative and indebted to Hector Martinez-Valdez, M.D., Ph.D., for his tremendous efforts to make this book the best of its kind by editing chapters. We are also grateful to contributors to all chapters in a timely manner. We wish to thank Dr. Edmund Kim for his encouragement of our work. Finally, we wish to thank Ms. Hemalatha Gunasekaran and Ms. Yoko Arai at Springer Nature who supported us in the creation and editing of this book.

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Contents

Part I Chemical Biology Aspect

1 PET Radiotracers for Tumor Imaging .................................................. 3Ming-Rong Zhang

2 Theranostic Approaches for Pathway- Activated Systems in Oncology ............................................................................... 17Lan V. Pham, Jerry L. Bryant, David Yang, and Richard J. Ford

3 Optimization of Aseptic Production in PET Radiopharmaceuticals for Compliance to the Most Current GMP........................................... 43Chih-Hao K. Kao

Part II Engineering and Instrumentation Aspect

4 OpenPET Enabling PET Imaging During Radiotherapy ................... 55Taiga Yamaya and Hideaki Tashima

5 High-Resolution PET/CT Development ................................................ 85Wai-Hoi Wong and Yuxuan Zhang

Part III Social Aspect

6 Ethics, Regulations, and Clinical Development of Precision Medicine: Activating with Molecular Imaging ..................................... 105Chieko Kurihara and Tomio Inoue

7 Optical Imaging: How Far Can We Go ................................................. 127Wenbin Ma, Sara Herrera-Rodriguez, and Hector Martinez-Valdez

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Part IV Clinical Aspect

8 Applications of UIH High-Resolution PET/CT in Zhongshan Hospital ............................................................ 153Jun Zhou and Hongcheng Shi

9 Clinical Evaluation of Focused High- Resolution Breast PET............. 171Kanae Kawai Miyake and Yuji Nakamoto

10 Recent Developments with Large-Bore PET/CT ................................. 185Tomohiro Kaneta and Tomio Inoue

11 Copper-Labeled Radiopharmaceuticals in Oncology .......................... 195Hiroaki Kurihara

12 PET Radiopharmaceuticals in Oncology Beyond FDG....................... 213Shao-Li Song, Pan-Li Li, and Gang Huang

13 Molecular Imaging in Oncology: Challenges and Opportunities in the Twenty-First Century ................................................................... 223Ya-Ming Li, David Yang, and Xue-Na Li

14 Visualization of Apoptosis: Annexin V Imaging ................................... 233Skye Hsin-Hsien Yeh, Fan-Lin Kong, and Ming-Hsien Lin

Contents

Part IChemical Biology Aspect

3© Springer Nature Singapore Pte Ltd. 2017 T. Inoue et al. (eds.), Personalized Pathway-Activated Systems Imaging in Oncology, DOI 10.1007/978-981-10-3349-0_1

Chapter 1PET Radiotracers for Tumor Imaging

Ming-Rong Zhang

1.1 Introduction

Positron emission tomography (PET) is a nuclear medicine imaging technique that produces a three-dimensional functional image of the living body. This system detects pairs of gamma rays emitted indirectly by a positron-emitting radiotracer, which is introduced into the body as a biologically active tracer. Three-dimensional images of radiotracer concentration within the body are then constructed by com-puter graphic analysis. PET is both a medical and study tool used in clinical oncol-ogy (medical imaging and diagnosis of tumors and the search for metastases) and in preclinical animal studies, where it allows repeated scans of the same subject. PET is particularly valuable in cancer research because it increases the statistical quality of the data (research subjects can act as their own controls) and substantially reduces the number of animals needed for individual studies.

PET studies using radiotracers are leading to rapid advances in the personalized diagnosis and treatment of cancers. The ability to translate oncogene signatures into functional imaging data, such as those obtained using PET via specific targeted radiotracers, permits noninvasive and quantitative visualization of tumors at multi-ple time points and at the whole-body level, thereby facilitating personalized drug development, clinical trials, and patient management.

The development of labeled radiotracers using positron-emitting radionuclides is required for the PET technique. In this review, the author will introduce [18F]FDG and post-[18F]FDG radiotracers, review recent advances in the development of PET tumor imaging radiotracers, and present results to develop novel PET radiotracers for tumor imaging.

M.-R. Zhang (*) Department of Radiopharmaceutics Development, National Institute of Radiological Sciences, National Institutes for Quantum and Radiological Science and Technology, 4-9-1 Anagawa, 263-8555 Inage-ku, Chiba, Japane-mail: [email protected]

mailto:[email protected]

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1.2 [18F]FDG for Glucose Metabolism

PET scanning with the radiotracer [18F]FDG (Fig. 1.1) is widely used in tumor imaging for clinical studies or diagnosis [1–3]. This radiotracer is derived from glucose, taken up by glucose-consuming cells and phosphorylated by mitochondrial hexokinase, which is greatly elevated in rapidly growing malignant tumors. Because the hydroxyl group in glucose (replaced by 18F to generate [18F]FDG) is required for the next step in the glucose metabolism in all cells, [18F]FDG cannot participate in further reactions [4]. Furthermore, most tissues, except the liver and kidneys, cannot remove the phosphate added by the hexokinase. This means that [18F]FDG is trapped in cells that take it up until it decays, owing to their ionic properties and phosphory-lated sugars, which cannot be cleared from cells. This results in high level of radio-activity accumulating in tissues with high glucose uptake and metabolism, such as the brain, liver, and most cancers [5–10]. As a result, [18F]FDG-PET can be used for cancer diagnosis, staging, and monitoring of therapy [11]. Many individual solid tumors have been found to exhibit very high uptake of radioactivity, a fact that is useful when searching for tumor metastasis or for recurrence after the removal of a primary tumor known to be highly active. PET oncology scans using [18F]FDG make up over 90 % of all PET scans in the current medical practice.

Although PET facilities are rapidly increasing worldwide, the only PET radio-pharmaceutical currently available for diagnosis is [18F]FDG. In keeping with this notion, [18F]FDG is by far the most widely used radiotracer for clinical purposes, but its application has some shortcomings. Since [18F]FDG is a derivative of glu-cose, which is taken up by cells via glucose transporters, phosphorylated by hexo-kinase, and retained in the tissue, its high uptake is not only in tumor cells but also in normal tissues, such as the brain and heart, which have high levels of glucose metabolic activity [12, 13]. Hence, [18F]FDG-PET is not suitable for imaging tumors in these tissues. In addition, because of its high uptake in urine and fast excretion from the bladder, it is difficult to detect tumors in this organ and surrounding tissues using [18F]FDG [5]. Most importantly, because of high uptake in inflamed tissues, it is difficult to distinguish tumor from inflammation using [18F]FDG.

To increase the usefulness of PET and to overcome the disadvantages associated with [18F]FDG, it is important to develop new imaging radiotracers, which use alter-native mechanisms for tumor visualization and provide different information to that obtained using [18F]FDG. Moreover, higher tumor specificity than the one provided by [18F]FDG could be achieved using new radiotracers.

In the following sections, the author will introduce the main candidates for post-[18F]FDG tumor imaging radiotracers.

M.-R. Zhang

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1.3 Amino Acids

The amino acid analog most frequently used as a radiotracer is [11C]methionine ([11C]Met, Fig. 1.2) [14, 15]. [11C]Met is easily synthesized using [11C]methyl iodide or [11C]methyl triflate as the radiolabeling agent. Since the precursor for [11C]Met radiosynthesis is L-homocysteine, the product obtained is only the L-isomer. In the normal brain, where protein metabolism levels are low, PET radiotracers reflecting controlled protein biosynthesis/degradation rates are suitable for the detection of glioma tumors in many PET facilities [16, 17]. However, because the S-[11C]methyl group in the cell is relatively easily transferred into other positions, compared to 11C labeling in other positions, the levels of [11C]Met are insufficient to enable evalua-tion of protein synthesis ability.

Since 1960, amino acids that are stable against metabolism (unlike natural amino acids) have been developed. These artificial amino acids were labeled with 11C and used to detect tumors in preclinical studies. Among them, α-aminoisobutanoic acid, α-aminocyclobutane-1-carboxylic acid, and α-aminocyclopentane-1-carboxylic acid show high binding affinity for amino acid transporters. Moreover, these amino acid analogs do not contain chiral carbon atoms; therefore, their radiolabeled ver-sions are considered promising probes for tumor imaging.

Many 18F-labeled amino acid analogs have been developed and evaluated as can-didate of post-[18F]FDG radiotracers (Fig. 1.2). Although phenylalanine and tyro-sine analogs were found to correlate with protein synthesis, these analogs are unable to participate in protein synthesis, and their tumor uptake levels are associated with amino acid transporter activity. Then, in efforts to improve in vivo metabolic stabil-ity, 3-[18F]fluoro-α-methyl-L-tyrosine ([18F]FMT, Fig. 1.2) [18] and 4-[18F]fluoroethyl- L-tyrosine ([18F]FET) [19] were developed. More recently, [18F]FACBC has been reported as the most promising radiolabeled amino acid analog [20, 21]. This radiotracer has two stereoisomers, with the cis isomer exhibiting higher selec-tivity than the anti-isomer. In clinical glioblastoma imaging studies, which cannot use [18F]FDG-PET, [18F]FACBC can provide high-quality PET tumor images.

Fig. 1.1 [18F]FDG

(A) (B) (C) (D)

Fig. 1.2 (a) [11C]Met, (b) [18F]FMT, (c) [18F]FET, and (d) [18F]FACBC

1 PET Radiotracers for Tumor Imaging

6

Therefore, the 18F-labeled artificial amino acid analogs are viable alternatives for the detection of tumors that cannot be successfully visualized by [18F]FDG-PET.

1.4 Nucleic Acids

[3H]Methylthymidine has been synthesized and used for many years. In addition, methods for production and in vivo evaluation of the PET tracers [11C]methylthymi-dine (Fig. 1.3) and 2-[11C]thymidine have been reported since 1980 [22, 23]. However, these natural nucleic acid analogs are not stable in vivo, which prompted the modification of their chemical structures and radiolabeling them with 18F, thus creating a probe a longer half-life than 11C.

3′-Deoxy-3′-[18F]fluorothymidine ([18F]FLT) is an analog of thymidine, in which 18F is introduced in the 3′-position and shows high in vivo stability and strong resis-tance to metabolism by cellular thymidine phosphatase [24]. Blood circulating [18F]FLT is taken up by the tissues via a pyrimidine transporter, which functions in nucleic acid synthesis. As the hydroxyl group in 3′-position is replaced by fluorine, phosphorylated [18F]FLT-5′-P cannot participate in the synthesis of DNA, and it is thus retained in the cell as a monophosphate. Hence, the uptake of [18F]FLT can reflect thymidine kinase-1 (TK1) activity in tumor cells [25], which is very low in the G0 stage of the cell cycle and reaches a maximum between the G1 and S phases. Therefore, [18F]FLT has been used to evaluate cell proliferation for tumor staging and assess the therapeutic effects of anticancer drugs.

1.5 Lipid Metabolism

Cancer is characterized by the high proliferation ability of tumor cells, and during this process the synthesis of cell membrane components increases accordingly. Therefore, membrane lipid synthesis is a useful target to also evaluate the

(A) (B) (C)

Fig. 1.3 (a) [11C]Methylthymidine, (b) 2-[11C]thymidine, and (c) [18F]FLT

M.-R. Zhang

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proliferation ability of tumor cells. To this end, [11C]acetate, [11C]choline, and [18F]fluorocholine have been developed and used in clinical imaging studies (Fig. 1.4) [26].

In vitro evaluation of [11C]acetate has been performed to determine the mecha-nism of radioactivity accumulation in tumor cells [27]. Acetate is formed by the metabolism of phosphatidylcholine and neutral lipids, and, thus, the accumulation of [11C]acetate radioactivity can reflect the proliferation ability of tumor cells [28, 29].

The uptake of choline analogs can reflect the activity of choline kinase and be used to indirectly evaluate ability to synthesize membrane lipids [30–32]. In clinical studies, PET radiotracers for lipid metabolism are useful for the detection of tumors in the brain, the bladder, and the urinary tract. [11C/18F]Choline analogs are not use-ful for the detection of cancer in epigastrium tissues and other organs because they have a high radioactivity uptake in the liver. Hence, acetate- and choline-based radiotracers may be better to evaluate the therapeutic effects of radiation and antitu-mor drugs.

1.6 Hypoxia

During the process of tumor cell proliferation, insufficient supply of oxygen results in hypoxia. Hypoxic areas of tumors are relatively insensitive to chemotherapy and radiation therapy. Thus, an understanding of the hypoxic state is useful for the pre-diction of therapeutic effects and the evaluation of treatment regimens, which has led to the development of PET imaging radiotracers designed to evaluate hypoxia

(A) (B) (C) (D) (E)

Fig. 1.4 (a) [11C]Acetate, (b) [18F]fluoroacetate, (c) [11C]choline, (d) [18F]fluoromethylcholine, and (e) [18F]fluoroethylcholine

(A) (B) (C)

Fig. 1.5 (a) [18F]FMISO, (b) [18F]FAZA, and (c) 62/64Cu-ATSM


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[33–35]. [18F]FMISO (Fig. 1.5) was the first nitroimidazole analog used for imaging of hypoxia in tumors [33]. The nitro group of misonidazole analogs is reduced to form a hydrophilic amine group, and this amine product binds to cellular compo-nents and is retained in tumor cells.

In hypoxic areas, which have a low blood flow, the initial uptake of radioactivity by one-pass circulation of radiotracer is low. However, the highly lipophilic [18F]FMISO is able to diffuse into the regions with low blood flow. Unfortunately, the slow clearance of radioactivity from blood means that PET imaging with [18F]FMISO may not result in high-quality images, and thus, extended scanning dura-tions are required [36]. As an alternative, the tracer [18F]FAZA, which has low lipophilicity and high hydrophilicity, has been developed and used in clinical stud-ies [37, 38]. Compared to [18F]FMISO, [18F]FAZA shows improved solubility in water and good signal/noise contrast in PET images within a relatively short PET scanning time.

In addition to nitroimidazole analogs, 62/64Cu-diacetyl-bis(N-4- methylthiosemicarbazone (62/64Cu-ATSM) is also a useful PET imaging radiotracer for hypoxia [39, 40]. Cu-ATSM is a small lipophilic molecular complex that easily penetrates the blood-brain barrier and cellular membranes and clears rapidly from normal tissues [41]. In hypoxic area, Cu2+ binding to ATSM is reduced to Cu+ by microsomal electron transfer, and the Cu+ component is retained in the cell. Compared to [18F]FMISO, 62/64Cu-ATSM shows rapid clearance from normal tissues and blood to produce images with good contrast and signal/noise ratios within a short PET scanning time.

1.7 Receptor and Angiogenesis

By transferring signals through various receptors overexpressed in tumor cells, cer-tain genes and proteins mediate tumor phenotypes, including proliferation ability, invasiveness, metastasis, and treatment resistance. PET studies using radiolabeled tracers for imaging of these receptors are useful for understanding the tumor characteristics.

To date, a large number of PET radiotracers for imaging of receptors have been reported. Major radiotracers of this type include: epidermal growth factor receptor (EGFR) human type (HER2), which is associated with poor prognosis in breast cancer; folic acid receptor, associated with malignant proliferation; chemokine receptor, associated with metastasis; glucagon-like peptide-1 (GLP-1) receptor, associated with neuroendocrine tumors; somatostatin receptor; tumor angiogenesis integrin receptor (αvβ3); and vascular endothelial growth factor (VEGF) [42]. A number of these radiotracers have been used in clinical studies.

M.-R. Zhang

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1.7.1 Somatostatin Receptors

Somatostatin receptors are G protein-coupled transmembrane proteins that are widely distributed in normal tissues, including those of the central nervous system, pancreas, anterior pituitary, thyroid gland, spleen, gastrointestinal tract, and adrenal gland. There are five somatostatin receptors, of which somatostatin receptor-2 is overexpressed in the majority of malignant tumors, including neuroendocrine can-cers, small cell lung cancer, breast cancer, and malignant lymphoma. The endoge-nous ligand of somatostatin receptors, somatostatin, has two isoforms of 14 and 28 amino acids, both of which demonstrate high binding affinity for somatostatin receptors. Owing to its short half-life in blood (2 min), imaging using somatostatin as a radiotracer is difficult. An analog of somatostatin, octreotide, which is formed from eight amino acid residues, has a longer half-life in blood (1.7 h) and higher metabolic stability than somatostatin. Many PET radiotracers derived from octreo-tide have been developed. Their chemical structures are illustrated in Fig. 1.6.

The PET radiotracers for somatostatin receptors, TOC, TATE, NOC- conjugated1,4,7,10-tetraazacyclododecane (DOTA), 68Ga-DOTATOC, 68Ga-DOTATATE, and 68Ga-DOTANOC] have been used in clinical studies (Fig. 1.6) [43, 44].

1.7.2 Integrin Receptor Subtype αvβ3

Angiogenesis is an important process during the proliferation of solid tumors. Cilengitide is a treatment developed to target integrin receptor (subtypes αvβ3 and αvβ5), which attenuates tumor angiogenesis. Almost all PET radiotracers targeting integrins contain the amino acid sequence, arginine-glycine-asparagine (RGD) [45, 46]. The RGD sequence, which is commonly found in extracellular matrix proteins,

Fig. 1.6 PET radiotracers targeting somatostatin receptors: (a) octreotide, (b) [Tyr3]-octreotide, (c) [Tyr3,Thr8]-octreotide, (d) [1-NaI3]-octreotide, and (e) NOC-conjugated 1,4,7,10- tetraazacyclododecane


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binds to integrins and shows particularly high affinity for the integrin αv subunit. Representative integrin-targeting PET radiotracers are 18F-, 68Ga-, and 64Cu-labeled compounds, including [18F]galacto-RGD, [18F]fluciclatide (AH111585), [18F]RGD- K5, 64Cu-DOTA-RGD, and 68Ga-NOTA-RGD (Fig. 1.7).

1.7.3 Other Receptors Involved in Tumor Angiogenesis

Another promising target molecule involved in tumor angiogenesis is VEGF. VEGF and its receptor VEGFR are highly expressed in endothelial cells, and VEGFR is a promising target for PET imaging [47]. To that end the VEGFR tyrosine kinase inhibitor sunitinib and VEGF monoclonal antibody bevacizumab were developed to image this pathway. Recently, bevacizumab and ranibizumab have been labeled using 89Zr (half-life, 78 h). 89Zr-bevacizumab and 89Zr-ranibizumab have demon-strated promise as potential PET imaging radiotracers for in vivo imaging of VEGF in tumor-bearing mice [48, 49].

PET imaging studies using octreotide-based and RGD-peptide-based molecules to target somatostatin and integrin receptors, respectively, have had some success in clinical studies. The imaging data resulting from targeting these molecules may be useful in clinical diagnosis, tumor staging, therapeutic strategies, and monitoring therapeutic effects. However, the clinical importance of imaging somatostatin and integrin receptors has not yet been fully determined, and more research is required to validate their clinical usefulness.

1.8 Metabotropic Glutamate 1 Receptor

Ectopic metabotropic glutamate 1 receptor (mGluR1) shows oncogenic activity and is becoming an important target for personalized diagnosis and treatment strategies for melanomas [50]. Evidence indicates that ectopically expressed mGluR1 inde-pendently induces melanocyte carcinogenesis. We have developed an oncoprotein- based PET imaging platform in melanomas for noninvasive visualization and quantitation of mGluR1 with a novel mGluR1-specific radiotracer, 4-[18F]fluoro-N-[4-[6-(isopropylamino)pyrimidin-4-yl]-1,3-thiazol-2-yl]-N-methylbenzamide ([18F]FITM; Fig. 1.8) [50]. [18F]FITM shows excellent pharmacokinetics, namely, the dense and specific accumulation of radioactivity in mGluR1-positive melano-mas B16F1 and B16F10, compared to mGluR1-negative hepatoma and normal tis-sues. Furthermore, accumulation levels of radioactivity corresponded to the extent of the tumor and to those of mGluR1 protein expression in melanomas and mela-noma metastases in the lung (Fig. 1.8). The [18F]FITM PET imaging platform is expected to open a new avenue for defining individualized therapeutic strategies, clinical trials, and patient management, as a noninvasive personalized diagnostic

M.-R. Zhang

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(A)

(C)

(B)

(D) (E)

Fig. 1.7 (a) [18F]Galacto-RGD, (b) [18F]fluciclatide, (c) [18F]RGD-K5, (d) 64Cu-DOTA-RGD, and (e) 68Ga-NOTA-RGD


12

tool. It can also become a useful research means to understand mGluR1-triggered oncologic events in melanomas. Recently, an iodine analog of [18F]FITM has been developed, and this compound can be further developed using 124I and 131I radiola-beling for long-duration PET scanning and radiotherapeutic applications [51].

1.9 Summary

Many PET radiotracers have been developed for imaging tumors to facilitate animal studies, clinical diagnosis, tumor staging, therapeutic strategy, and monitoring ther-apeutic effects. However, PET radiotracers for imaging a large number of molecular targets remain undeveloped. Moreover, as the targeted candidate molecules include small molecule compounds, peptides, and antibodies, determining how to label them efficiently using positron-emitted radionuclides is a challenging problem for their development as novel PET radiotracers. Radiolabeling techniques, including radionuclide production, preparation of radiolabeling agents, radiochemical

(A) [18F] FITM

B16F1/[18F]FITM

9

0

%ID/g

B16F10/[18F]FITM Hepa1-6/[18F]FITM

Control /[18F]FITM

Metastasis /[18F]FITM

8

0

%ID

/g

(B) (C)

Fig. 1.8 (a) Chemical structure of [18F]FITM. (b) Representative coronal [18F]FITM PET/CT images in B16F1-, B16F10-, and Hepa1-6 tumor-bearing mice. High accumulation of radioactivity was observed in the mGluR1-positive B16F1 and B16F10 melanomas, and low uptake was observed in the mGluR1-negative Hepa1-6 tumor. Green circles indicate tumors. (c) Representative coronal and axial [18F]FITM PET/CT images in mice with pulmonary metastatic melanoma and control mice without metastasis. Intense and heterogeneous accumulation of radioactivity was observed in the lungs bearing B16F10 metastasis with very low background signals

M.-R. Zhang

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reactions, and automated production, continue to be assessed. Clearly, it will take a great deal of time and effort to identify another gold standard PET radiotracer for tumor imaging that is similar or more specific than [18F]FDG.

Acknowledgment The author thanks Dr. Masayuki Fujinaga (National Institute of Radiological Sciences) for assistance in the preparation of this manuscript.

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Chapter 2Theranostic Approaches for Pathway- Activated Systems in Oncology

Lan V. Pham, Jerry L. Bryant, David Yang, and Richard J. Ford

Abstract Theranostics is a novel concept that refers to the integration of diagnostics with therapeutics in order to generate personalized therapies and is emerging as a promising precise therapeutic paradigm. In oncology, the approach is aimed at more accurate diagnosis of cancer, optimization of patient selection to identify those most likely to benefit from a proposed specific therapy allowing the generation of effective therapeutics that enhance patient survival. Perhaps the most promising target to date for theranostics is the deregulation of cancer cell metabolism, involving the uptake of glucose and glutamate, two key nutrients that are necessary to convert into glucos-amine to stimulate protein biosynthesis for the growth and survival of cancer cells. We have recently developed a novel technology whereby the chelator ethylenedicys-teine (EC) conjugates with glucosamine to create a vehicle platform (ECG), which mimics N-acetylglucosamine (GlcNAc) that targets highly proliferative cancer cells. Moreover, ECG can be further conjugated to diagnostic/therapeutic metals (rhenium, Re, and platinum, Pt) that function as a new theranostic agent suitable for personal-ized medicine, targeting key pathways in cancer cells such as highly metabolic dif-fuse large B-cell lymphoma (DLBCL). This chapter summarizes key signaling pathways linked to dysregulated glucose metabolism in DLBCL and how deregu-lated glucose metabolism can be utilized for developing innovative new technologies with theranostic applications to eradicate cancer.

2.1 Diffuse Large B-Cell Lymphoma Pathophysiology

Malignant B-cell lymphoma represents a major health risk in the USA and world-wide [1, 2]. Non-Hodgkin lymphomas (NHLs) are a common, accounting for about 4% of all cancers, but heterogeneous group of human B lymphocytic neoplasms (NHL-B), that primarily arise (~85%) within the B-cell lineage of the immune

L.V. Pham (*) • J.L. Bryant • R.J. Ford University of Texas MD Anderson Cancer Center, Houston, TX 77030, USAe-mail: [email protected]

D. Yang Vyripharm Biopharmaceuticals, University of Texas Health Science Center, Houston, TX, USA


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system. NHL-B represents the fifth most common cancer in the USA (~72,000 new cases/20,000 deaths) as reported in 2016 [3–5]. Notably, while diffuse large B-cell lymphoma (DLBCL) is the most common (30–40%) of histologically defined NHL- B, it is also the most heterogeneous [6, 7].

As an aggressive, diverse group of lymphoid neoplasms, DLBCL is associated with constitutive activation of key NF-κB signaling pathways [8, 9], although other lymphoma histotypes like Hodgkin and MALT lymphomas may show different acti-vation patterns, signaling proteins, and pathological variants in the canonical NF-κB pathway [10–12]. DLBCL has been extensively studied in recent years by various types of microarray gene analyses (supervised or unsupervised) [13, 14] and is con-sidered to consist of at least three definable genetic “signatures” or occasionally overlapping subtypes, based on various methods of gene expression profiling [15, 16]. Although these putative DLBCL subtypes appear valid, considerable pheno-typic genotypic overlaps clearly occur [17]. The Rosenwald/Staudt group, for instance, has equated the expression of MUM1/IRF4 and CD138 immunologic markers with constitutive activation of the NF-κB1 pathway as a specific gene array “signature” that defines an activated B-cell type (ABC-like DLBCL). In contrast, the nonrandom t(14;18) bcl2 cytogenetic translocation, expressing germinal center (GC) markers bcl-6 and CD10, defines a GC B-cell subtype (GCB-like DLBCL) [18–20]. In the era that preceded the standard frontline combination chemotherapy rituximab, cyclophosphamide, hydroxydaunomycin, Oncovin, and prednisone (R-CHOP), these DLBCL subtypes were reported to have disparate clinical out-comes with significantly different 5-year survival rates [21, 22], although better differential biomarkers are still needed.

Notably, studies on DLBCL-associated oncogenes have recently revealed new molecular insights regarding the role of bcl-6 in the pathogenesis of the GCB- DLBCL. For instance, Dalla-Favera’s group has shown that bcl-6 gene expression is regulated through the CD40-NF-κB canonical signaling pathway. They have also shown that in GCB-DLBCL, CD40 appears to activate the transcriptional factor IRF4 (MUM1) gene through the p50/p65 members of canonical NF-κB pathway. Once activated IRF4 binds to the bcl-6 promoter and directly represses its transcrip-tion [23]. This suggests that resistance to CD40-NF-κB-IRF4 signaling is an essen-tial mechanism of bcl-6 deregulation in GCB-DLBCL. The data also suggest that while CD40-NF-κB pathway is active in GCB- and ABC-type DLBCL, activation occurs by different mechanisms [24, 25].

Moreover, several studies have examined the potential role(s) of TNFR and related signaling pathways on DLBCL growth and survival (G/S) [26, 27], whose mechanisms are also shared by normal B lymphocytes. However, the difference between normal and malignant cells is that these G/S signaling pathways are aber-rantly dysregulated in the latter [28, 29], the hallmark of aggressive NHL-B patho-genesis. These studies initially identified the CD40 signalosome, consisting of the TNFR, CD40, and its cognate ligand (CD40L, CD154) as constitutively expressed in DLBCL cell lines and primary lymphoma cells from patients, along with the signaling components (TRAFs 2,6: IKK complex, c-rel/p65) of the canonical NFkB1 pathway [30]. When the CD40 signalosome was disrupted by antibodies to

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CD40 or CD154, the signalosome structure and NF-κB1 signaling pathway were interdicted, and cell death was induced through apoptosis [31–33]. Further studies on aberrant survival mechanisms in DLBCL led to the discovery of the role of another TNF superfamily member, the B-cell survival factor BLyS/BAFF, with con-stitutive expression of their receptor BR3 in aggressive NHL-B (DLBCL and MCL). These studies further revealed that constitutively activated BLyS/BR3 receptor- ligand interactions result in noncanonical (alternative) NF-κB2 pathway signaling, which in analogy to the CD40/CD154 cascade, the BLyS/BR3 signalosome pro-vides a constitutive positive ligand-/receptor-mediated feedback to the NF-κB2 pathway [31].

While earlier studies [33, 34] had also revealed that in addition to constitutive expression of the canonical NF-κB1 pathway, there is evidence that at least some members of the alternative NF-κB2 pathway (e.g., p52 and RelB) are also constitu-tively activated and expressed in DLBCLs and other aggressive NHL-B (e.g., MCL) [34, 35]. A key study examined a series of validated tissue microarrays (TMA) in DLBCL (ABC and GCB types) cell lines and primary patient samples and found that both NF-κB1 and NF-κB2 signaling pathways were constitutively activated in both ABC and GCB subsets, but that the pattern of activation and NF-κB dimer utilization was characteristically different [36]. These studies suggest that multiple interactive cell signaling pathways, including both canonical and alternative variant “hybrid” NF-κB pathways [37–39], contribute to G/S regulatory mechanisms in DLBCLs. While the canonical/classical NF-κB1 pathway has been studied in many cell types [40, 41], the alternative NF-κB2 pathway has only recently begun to receive attention [31, 42, 43]. Most of these studies, however, have been mainly performed in genetically engineered mice, which basically provided a general out-line of some of the regulatory interactions [44].

Examples of genetically altered murine models demonstrate that the alternative NF-κB pathway activation is controlled through a negative feedback mechanism involving increased protein levels of negative regulators of the adaptor proteins TRAF2/3, which inhibit the key upstream NF-κB2 kinase, NIK. Overexpression of wild-type NIK leads to B-cell hyperplasia caused by the amplification of BLyS- induced alternative NF-κB signals. Interruption of the interaction between TRAF3 and NIK induces constitutive BLyS-independent activation of the alternative path-way and leads to a large accumulation of mature B cells in lymphoid organs and disruption of structural integrity. Other studies have proposed a model where inter-actions between TRAF2 and TRAF3 constitutively block B-cell survival via the inhibition of activation of the alternative NF-κB2 pathway, suggesting a mechanism by which NIK accumulation is prevented, since NIK bears a TRAF3 interaction site that can result in NIK degradation by TRAF3 [45–48]. Interestingly, when BLyS occupies BR3, it sequesters TRAF3 and prevents interactions with TRAF2. This blocks TRAF3/2 interactions that would lead to greater NIK accumulation, subse-quent NF-κB2 processing, and hence increased B-cell survival [49]. The results from these findings imply that deregulated NIK expression may contribute to B-cell malignancies, particularly aggressive lymphomagenesis, even if NIK protein remains undetectable at the protein level [50].

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Moreover, there is compelling transgenic mouse model that obtained evidence of an important role of CD40 in B-cell lymphomagenesis, resulting from constitutively active CD40 receptor expression, which leads to B-cell-specific enforced activation of the noncanonical NF-κB pathway. Consistent with these findings, LMP1-/CD40- expressing mice developed a high incidence of B-cell lymphomas, indicating that interactions of the signaling pathways induced by constitutive CD40 signaling are sufficient to initiate a neoplastic B-cell process, likely leading to the development of B-cell lymphomas.

On the other hand, the noncanonical NF-κB2 pathway has been occasionally implicated in B lymphoid malignancies associated to some cases of chromosomal abnormalities leading to the production of truncated p100 proteins with diminished NF-κB inhibitory ability [45–47]. For instance, two studies in primary multiple myeloma (MM) samples and cell lines exhibited genetic aberrations that affect mediators of NF-κB activation, mostly involving the alternative NF-κB2 pathway. The aberrations led to the absence of negative regulators of NF-κB, such as TRAF3, TRAF2, and c-IAP1/2, or to overexpression of NIK [51, 52]. These studies imply that deregulation of the TRAF3-NIK axis could also play an important role in B-cell lymphomagenesis [50]. Although the data further indicate that elevated canonical and noncanonical NF-κB activity by deregulation of NIK directly contributes to disease progression in primary MM, they also suggest that TRAF3 can function as an important suppressor of lymphoid neoplasia through the negative regulation of both the canonical and noncanonical NF-κB pathways [53, 54].

It must be noted that despite its B-cell lineage, MM cells represent a neoplasm mimicking plasma cells, which is quite different from DLBCL, although some plasmablastoid lymphomas can be quite plasmacytoid. In keeping with this notion, there is little information on the intrinsic nature of NF-κB2 pathway in DLBCL. However, a report by Kim et al. [55] demonstrated the activation of NF-κB2 components by BLyS in DLBCL subsets, while others have described the involvement of several oncogenes in DLBCL and identified NFAT expression as a candidate oncogene in the ABC subtype. Similarly, a variety of genetic abnormali-ties were identified in DLBCL that were associated with either the ABC or GCB- DLBCL subsets [56, 57].

Although deregulation of NF-κB signaling can be a key mediator of transcription factor (TF) heteromer formation, which targets NF-κB-regulated G/S genes in DLBCLs, other regulators like NFAT (distantly related to NF-κBs [58]) are known to be central for chromatin structural remodeling, which in turn actively modulates gene transcription. The NFAT family of proteins are also Ca2+-inducible transcrip-tion factors that prominently stimulate the expression of a wide range of immune response genes in activated T cells [59]. However, we have shown that NFAT plays key regulatory roles in B lymphocytes, particularly in aggressive NHL-B [34, 60]. NFAT-dependent promoters and enhancers rapidly undergo extensive chromatin remodeling to form deoxyribonuclease I (DNAse 1)-hypersensitive sites (HSSs)

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[61, 62]. NFAT is likely to be at least a driving force involved in chromatin remodeling, which has recently been described as a major NFAT function [63], since NFAT sites are necessary and sufficient to activate DNA-1-driven chromatin HSSs. Chromatin remodeling may well be a primary function of NFAT elements, since even high-affinity NFAT binding to promoter targets results in relatively weak transcriptional activation without bonding between NFAT and transcription factor (TF) associates [64]. The activating protein 1 (AP-1) is the most common TF partner that is directly recruited by NFAT, at the HSS domains, and dimeric NFAT-AP-1 and DNA response elements are very efficient in removing nucleosomes [62]. NFAT-AP-1 complexes recruit both histone acetyltransferases (HATs) and the ATP- dependent SWI-SNF family of chromatin remodelers [65, 66], which together provide the necessary functional proteins required to modify and rearrange nucleosomes.

NFAT may also help to organize chromatin domains and enable enhancer- promoter communication [67]. In activated T cells, inducible intrachromosomal looping occurs between the tumor necrosis factor (TNFα) gene promoter [68] and two NFAT-dependent enhancers located within 9 kb upstream (-) and 3 kb down-stream (+) of the promoter region [69]. This topology places the TNF gene and the adjacent lymphotoxin (LT) genes in separate loops, thereby allowing independent regulation of the TNFα gene within multigene loci. This new data supports earlier studies proposing that NFAT functions through the disruption of nucleosomes within enhancers, mobilizing nucleosomes across extensive chromatin domains and linking enhancers and promoters. These studies identify NFAT as a factor that cre-ates a chromatin environment, which is permissive for both the recruitment and aggregation of factors to coordinately control transcriptional processes at promoter and enhancer regions [70]. The specific role that NFAT plays in the multifaceted process of locus activation is still unclear, but its role as an effective orchestrator of essential steps in creating an accessible chromatin environment is compelling [63].

NFAT functions have recently also been linked to tumor immunity [71, 72] that includes the development of T-cell dysfunctions such as CD4+ T-cell anergy [73] and CD8+ T-cell exhaustion, which occur in a variety of cancers [74]. It is conceiv-able that NFAT regulates PD-1 expression in anergic T cells [75], and thus, the inhibition of NFAT or the kinase-regulating NFAT pathway would enhance T-cell function [76, 77]. These findings support the premise that targeting NFAT pathway in cancer patients could lead to tumor cell killing, either directly by abrogating NFAT-dependent cell survival or indirectly by blocking T-cell activation. Consistent with this hypothesis, a recent study by Ron Levy’s group [78] demonstrated that a combination treatment with anti-PD-L1 antibodies and ibrutinib, a Bruton’s tyro-sine kinase (BTK) inhibitor, leads impressive antitumor effects in animal lymphoma models, as well as in breast and colon cancer models. Their findings lead to the speculation that ibrutinib concomitantly targets BTK and interleukin-2-inducible kinase (ITK) to inhibit NFAT activation and thus could play a key role in T-cell- mediated therapies [79].


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2.2 Linking Deregulated Signaling Pathways to Cancer Cell Metabolism in DLBCL

Thioredoxin (TRX) is an integral antioxidant system, which maintains the intracel-lular redox state and hence a strong candidate to coordinately target a family of proteins to restore sensitivity to chemotherapy [80]. One family member is thiore-doxin- 1 (TRX-1), a low-molecular-weight (10–12,000) cellular redox protein, which is present in the nucleus and cytoplasm to regulate the activity of various enzymes, including those that counteract oxidative stress within the cell [81]. Intracellular TRX-1 exerts most of its antioxidant properties through scavenging of reactive oxygen species (ROS). Moreover, it plays an important role in the regula-tion of redox-sensitive transcription factors [82] and acts as a proto-oncogene that stimulates tumor growth and inhibits both programmed and drug-induced cell death [83]. Its increased expression is associated with enhanced HIF-1α (hypoxia-induced factor 11α) levels and transactivation in cancer cells [84], which result in high pro-duction of vascular endothelial growth factor (VEGF) and enhanced tumor angio-genesis [85]. Additionally, its overexpression has been correlated with aggressive tumor growth, poorer prognosis, and shortened patient survival [86].

TRX-1 appears to have an important role in maintaining the transformed pheno-type of some human cancers as well as their resistance to chemotherapeutic drugs. These functions make it a rational target for cancer drug development, and recent experiments support such potential. To that end, the Leukemia/Lymphoma Molecular Profiling Project (LLMPP) used a microarray technology to define a molecular profile for each of 240 patients with DLBCL and developed a molecular outcome predictor score that accurately determines patient survival. The study found that DLBCL patients with the worst prognosis, according to the outcome predictor score, had decreased expression of TXNIP, a protein that naturally inhibits TRX-1 activity [87]. However, these studies had little or no follow-up on the patho-physiologic impact of TRX-1- and TXNIP-controlled reduction-oxidation (redox) state in B-cell lymphomas.

TXNIP is known to regulate the cellular redox state by binding to and inhibiting thioredoxin in a redox-dependent fashion [88]. Recent studies, however, demon-strated that TXNIP is also a potent negative regulator of glucose uptake [89, 90]. In response to glucose uptake, cells activate a key TF complex that includes the Mondo member A and the Max-like protein (MondoA:Mix), which then enters the nucleus, binds the TXNIP promoter, and upregulates TXNIP gene transcription. How TXNIP blocks glucose uptake is not completely clear, but maintenance of energy homeosta-sis is clearly regulated through TXNIP. Furthermore, conventional TXNIP-deficient mice revealed that mitochondria were functionally and structurally altered, leading to reduced oxygen consumption and enhanced anaerobic glycolysis [91, 92]. On the other hand, tissue-specific knockout mice showed that TXNIP is essential for main-taining hematopoietic stem cell (HSC) quiescence and homeostatic interactions between HSCs and the bone marrow niche. In addition, targeted deletion of TXNIP causes cardiac dysfunction in response to pressure overload, primarily due to the

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dysregulation of mitochondria, which switches from oxidative phosphorylation to anaerobic glycolysis [92].

Notably, the energy homeostasis is defective in cancer since the TXNIP gene is repressed in many tumors by posttranscriptional and translational mechanisms [91]. These collective data indicate that active glucose metabolism together with acti-vated TRX-1 plays a key role in the pathophysiology of DLBCL. Since TXNIP controls both redox and glucose levels in cancer, its expression could be a promi-nent target for the therapy of DLBCL. In agreement with this hypothesis, it has been previously shown that epigenetic histone deacetylase (HDAC) inhibitors like vori-nostat and histone methyltransferase inhibitors like 3-deazaneplanocin A (DZNep) can reactivate TXNIP gene expression and inhibit TRX-1 in cancer cells [93, 94]. Of note, one of the main functions of DZNep is to disrupt the polycomb-repressive complex 2 (PRC2) by inhibiting the enhancer of the zeste homolog 2 (EZH2) pro-tein [95]. EZH2 is the catalytic subunit of PRC2, which is a highly conserved his-tone methyltransferase that targets lysine-27 of histone H3 [96]. This methylated H3-K27 chromatin site is commonly associated with the silencing of differentiation genes in organisms ranging from plants to humans. Studies in human tumors have shown that EZH2 is frequently overexpressed in a wide variety of malignant tissues, including lymphomas [97]. Although the mechanistic contribution of EZH2 to can-cer progression is not yet determined, functional links between EZH2-mediated histone methylation and histone acetylation suggest a partnership with the gene- silencing machinery implicated in tumor suppressor loss. Of particular pertinent are key studies using whole-genome sequencing in primary DLBCL, which identified frequently recurrent somatic heterozygous mutations in the EZH2 locus [98, 99]. The EZH2 mutations usually result in a gene gain-of-function that acts dominantly to increase histone methylation, particularly histone H3 Lys-27-trimethylation (H3K27me3) [100]. Taken together, these data suggest that epigenetic repression of TXNIP by the PCR2-EZH2 complex is involved in the TRX-1 gene-dependent hyperactivation of glucose metabolism in DLBCL and that targeting this pathway by small-molecule inhibitors has significant potential to reverse the resistance of DLBCL to chemotherapy [101].

On a different front, the c-myc proto-oncogene has also been shown to be involved in controlling key metabolic pathways in cancer [102, 103]. For instance, MYC overexpression has been recognized in aggressive B-cell lymphomas, primar-ily due to chromosomal translocations, which inevitably bear an adverse prognosis [104]. The MYC transcriptional network has been also shown to include noncoding microRNA (miRNA) regulators, such as miR-101 and miR-26a, which are linked to the epigenetic control of EZH2 metabolic pathways [105–107]. Previous studies have shown that MYC is transcriptionally regulated by chromatin enhancer func-tions, which involve the transcription factors NFATc1, NF-κB, and STAT3 [108], underscoring the impact of genetic and/or epigenetic dysregulation of these meta-bolic pathways in DLBCL (Diagram 2.1).

The incidence of DLBCL has been rising in recent decades, a situation that underscores the need to improve therapy with greater efficacy and fewer adverse effects. The fundamental problem is that while standard frontline combination


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chemotherapy of DLBCL with rituximab, cyclophosphamide, hydroxydaunomycin, Oncovin, and prednisone (R-CHOP) achieves lasting therapeutic remissions, it does not usually lead to complete cure. Furthermore, the adverse effects are too toxic for many older patients and pose long-term risks for younger patients. The development of new, affordable, effective, and low-toxicity frontline regimens against DLBCL, which target specific pathways, is feasible but will take many years to achieve and may still be suboptimal if pursued by conventional means.

Diagram 2.1 Epigenetic dysregulation of the metabolic signaling pathways that control cell growth, survival, and chemoresistance mechanisms in DLBCL. Our previous studies have shown that deregulated NF-kB, NFAT, and STAT3 signaling pathways alter the expression of MYC, a key oncogene in DLBCL that is frequently amplified as a result of chromosomal translocations. MYC has recently been shown to negatively regulate miR-101 and miR-26a, which are known to sup-press EZH2 expression. Our model hypothesizes that deregulation of EZH2 leads to the epigenetic silencing of the thioredoxin-interacting protein (TXNIP), a key negative regulator of thioredoxin, glucose metabolism, and bcl-6. The result is the hyperactivation of thioredoxin, glucose metabo-lism, as well as bcl-6, which are highly activated in some DLBCL, which causes uncontrolled tumor cell growth survival, and chemoresistance, a hallmark of lymphomagenesis. Hence, these pathways are rational targets for the design and application of innovative therapies, including ther-anostic approaches, to specifically reverse the resistance of DLBCL to chemotherapy

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2.3 Deregulated Glycolytic Pathway in DLBCL by Monitoring Through Nuclear Imaging

The energy consumed by the cells in the form of adenosine triphosphate (ATP) is generated from two main sources, glycolysis and the tricarboxylic acid (TCA) or Krebs cycle, which are required for normal and malignant cell proliferation and survival. The Warburg effect describes a mechanism by which most cancer cells consume glucose to be converted into ATP via aerobic glycolysis. DLBCL is known to be an aggressive disease, which exhibits high cell proliferation and glucose metabolism rates and influences the response to therapy. As a result of avid glucose consumption, DLBCL cells show higher uptake of fluorine-18F-deoxyglucose (18FDG) by positron-emission tomography (PET) than any other B-cell NHL. Moreover, recent studies link the increased glucose transporter type 1, 2, and 3 (Glut1, 2, and 3) expression and hexokinase II (HKII) activity to the pathogenesis of many hematological malignancies. In support of those studies, we have also found that aggressive B-cell lymphomas express high Glut1, Glut3, and HKII in DLBCL cell lines and primary tumor cells from patients (Fig. 2.1). Several onco-genes and signaling pathways have been implicated in the regulation of cancer cell glycolysis, particularly in DLBCL (Diagram 2.1).

2.4 Cancer Metabolism and Theranostic Approaches in DLBCL

Theranostics is a novel concept that refers to the integration of imaging diagnostics and therapy, which is emerging as a promising therapeutic paradigm [109]. It is an evolving field related to but different from traditional imaging and therapeutics. It embraces multiple techniques to arrive at in vivo molecular imaging, comprehen-sive diagnostics, and a personalized treatment regimen. Over the past decade, tre-mendous effort has been put forth to design and develop methods to produce highly efficient delivery vehicles for theranostic approaches. Liposomes, polymeric nanoparticles (including gold and other metals), dendrimers, carbon nanotubes, and quantum dots are examples of nano-formulations that can be used as multifunc-tional platforms for cancer theranostics [110]. However, these platforms have their limitations, and they have not been thoroughly developed for effective clinical utili-zation (Fig. 2.2).

In keeping with the unique features of the previously discussed pathways, the most promising target for personalized theranostics is targeting glucose metabolism of cancer cells because unlike normal cells, they metabolize glucose by aerobic glycolysis. Briefly, aerobic glycolysis, also known as the Warburg effect, is charac-terized by increased glycolysis and lactate production [111], which is often accom-panied by increased cellular glucose uptake. Notably, glucose uptake can be imaged in patient tumors by 18FDG-PET [112, 113]. 18FDG-PET is used clinically as a


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Fig. 2.1 Overexpression of glycolytic pathway proteins in DLBCL. (a) Purified whole cell extracts from normal B lymphocytes obtained from five healthy donors and representative DLBCL cell lines were subjected to Western blot for Glut1, Glut2, HK2, and actin (loading control). (b) HK2 protein expression level in DLBCL cells was compared to normal B lymphocytes. Quantitatively, HK2 protein expression in DLBCL is significantly higher, approximately fourfold higher, in nor-mal B lymphocytes. (c) Tissue microarray (TMA) analysis of HK2 protein expression in 93 cases of primary DLBCL tissue. (d) Table showing the summary of the HK2 expression, low (11%), intermediate (43%), and high (46%) in the TMA

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staging tool for diverse types of cancers, including DLBCL [114, 115], and experi-mental PET tracer probes can distinguish cancer cells from normal cells on the basis of increased glucose metabolism. In addition, positive images of residual 18FDG after therapy are predictive of a poor prognosis and survival of patients with refrac-tory aggressive lymphomas [116, 117]. Initial reports suggested that 18FDG-PET/computed tomography (CT) scans, performed early during treatment (interim PET) after 2–4 courses of CHOP chemotherapy in aggressive B-cell lymphoma (DLBCL in particular), could identify patients who were likely to relapse (Fig. 2.1) [117–119]. However, there are conflicting data on the PET/CT scans performed before treatment (initial PET) in lymphoma patients [120]. On the other hand and irrespec-tive of the interim PET results, the studies indicated that such imaging modalities could swiftly identify lymphoma patients who were likely to respond poorly to induction therapy or frontline treatment, which would prompt an indication for a shift to either intensified regimens or a theranostic approach. Importantly, cancer cells not only consume glucose but also large amounts of glutamine, a key amino acid involved in tumor protein synthesis [121, 122]. Among its various roles, gluta-

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mine is a precursor amino acid, which in combination with high glucose levels initi-ates the hexosamine biosynthetic pathway to synthesize glucosamine [123]. Briefly, fructose-6-phosphate from the glycolytic pathway combines with glutamine in the presence of the initiating enzyme glutamine:fructose-6-phosphate transferase (GFAT) to synthesize glucosamine-6-phosphate. A series of subsequent enzymatic steps lead to the production of uridine diphosphate-N-acetylglucosamine (UDP- GlcNAc), a substrate for O-linked glycosylation that is regulated by the terminating enzyme O-linked GlcNAc transferase (OGT). OGT is the enzyme responsible for the addition of a single N-acetylglucosamine (GlcNAc) residue to the hydroxyl groups of serine and/or threonine residues of target proteins. The hexosamine sig-naling pathway terminating in O-linked GlcNAc (O-GlcNAc) cycling has been implicated in cellular signaling cascades and regulation of transcription factors involved in cancer biology [124–127]. The biological relevance of the hexosamine biosynthetic signaling pathway has not been completely elucidated, and hence, assessing the impact of altered O-GlcNAc metabolism in tumors such as DLBCL would be useful to determine whether the pathway is a relevant target for the design of personalized theranostics. Analogous to glucose, GlcNAc can be taken up by the cellular glucose transporters [128, 129] and can replace glucose in glucose-depleted cancer cells [130], which supports its relevance as a theranostic probe.

2.5 Development of Targeted Molecule ECG that Mimics GlcNAc for Theranostic Approaches

We have devised a metabolic agent that mimics GlcNAc by conjugating the chelator ethylenedicysteine (EC) to two molecules of D-glucosamine [131]. The end result is ECG, a metabolic agent containing two molecules of GlcNAc. At the core of ECG is the chelator, which can bind to diagnostic/therapeutic metals, which can then trace and kill cancer cells (Diagram 2.2).

In terms of diagnostic imaging, the technetium-99m-based 99mTc-ECG radio-pharmaceutical has been shown to be an effective imaging agent for various types of cancers in both rodents and humans [131, 132]. The biopharmaceutical company Cell>Point is currently sponsoring Phase III clinical trials for its first 99mTc-ECG product for diagnostic imaging in oncology. The multicenter clinical trial is compar-ing 99mTc-ECG/single-photon emission CT (SPECT) imaging with FDG-PET imag-ing to assess and stage patients with non-small cell lung cancer. Remarkably, the Phase I/II results indicate that 99mTc-ECG/SPECT has a higher specificity than 18FDG-PET imaging for detecting tumor metastasis and differentiating between inflammation/infection and tumor recurrence [132]. Following this trial, Cell>Point (Centennial, CO) plans to sponsor Phase IV clinical trials to evaluate 99mTc-ECG in non-Hodgkin lymphoma (Diagram 2.2) and other types of cancer, for the diagnosis and staging of the disease process. Unlike FDG, ECG is not taken up into the brain or inflammatory/infection tissues and therefore has a lower false-positive rate in

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cancer diagnosis. More importantly, ECG has little or no toxicity to normal tissues in the human body, suggesting that ECG is an excellent vehicle to deliver therapeutic metals to cancer cells. The current interest is to integrate ECG imaging function to its therapeutic potential in a theranostic approach to treat refractory DLBCL and highly metabolic cancers.

To date, platinum-based drugs like cisplatin remain one of the most effective classes of chemotherapeutic agents in clinical use. However, the clinical use of cis-platin is quite limited by dose-dependent adverse effects. More effort should be directed to combat the severe systemic toxicity of traditional platinum anticancer agents by designing therapy systems that exclusively deliver platinum or other metallic complexes to tumor cells. To that end, we have chosen two cold metallic agents, rhenium 187 (Re) and a cis-platinum derivative (Pt), for conjugation with our metabolic agent ECG (see Diagram 2.1). We have tested the feasibility of the metal-ECG conjugation technology and the specific targeting of glucose and gluta-mine metabolism as a novel theranostic approach in refractory DLBCL (Fig. 2.3) (Diagram 2.3) [133].

Diagram 2.2 Synthesis of metalic ECG. D-glucosamine hydrochloride salt was added to ethyl-enedicysteine, giving rise to ethylenedicysteine with two sugar moiety similar to N-Acetyl-Glucosamine on both sides. The core of ECG is the chelator that binds to diagnostic (99mTc, 68Ga, and Gd) or Therapeutic (187Re or 188Re and Cis-Pt) metals


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Fig. 2.3 Effect of metallic Re-ECG in DLBCL. (a) In vitro viability assays were assessed in 14 representative DLBCL cell lines treated with increasing concentration of Re-ECG. (b) Representative DLBCL cell lines sensitive to Re-ECG or less sensitive to Re-ECG. (c) DLBCL cell lines that are sensitive to Re-ECG are highly proliferative in comparison to DLBCL cell lines that are less sensitive to Re-ECG. (d) Cellular uptake of Re-ECG is more significant in DLBCL

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Control Re-ECG Pt-ECG Control Re-ECG Pt-ECG

SUDHL-4 OCI-LY10

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Fig. 2.3 (continued) cells that are sensitive to Re-ECG (LY-10) in comparison to less sensitive cells (MS). (e) Western blot analysis showing Re-ECG inducing the DNA damage marker pH2AX in two representative DLBCL cell lines. (f) Confocal microscopy analysis showing the induction of the DNA damage marker pH2AX in OCI-LY10 cells treated with Re-ECG. (g) Cellular damage by measuring DNA activity with increasing platinum-ECG concentrations. (h) Western blot analy-sis showing platinum-ECG inducing the DNA damage marker pH2AX in two representative DLBCL cell lines


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2.6 Theranostic Potential of Metallic ECG in DLBCL

Our model proposes that key growth/survival transcription factors (NFATc1 and p65) in DLBCL, which are activated by upstream signaling pathways, are also regulated by glucose metabolism through the O-linked GlcNAc/hexosamine biosynthetic pathway (Diagram 2.4). Fluxes through the hexosamine biosynthetic pathway involve interaction between the substrate of the glycolytic pathway (fructose- 6- phosphate) and glutamine, which in the presence of the initiating enzyme GFAT synthesizes glucosamine-6-phosphate. A series of subsequent enzymatic steps leads to production of UDP-GlcNAc, a substrate for O-linked glycosylation that is regulated by the terminating enzyme OGT. OGT is the enzyme responsible for the addition of a single GlcNAc to the hydroxyl groups of serine and/or threonine residues of target proteins. The metabolic agent ECG mimics GlcNAc and can be taken up easily by cancer cells through glucose transporters and hexosamine path-way; it enters the nucleus via “piggybacking” with OGT-modified nuclear proteins. We propose that increased fluxes through the hexosamine biosynthetic pathway accordingly yield elevations in O-GlcNAcylation status in DLBCL, constituting a new pathophysiologic process in the regulation and activation of key transcription factors that control growth/survival mechanisms. ECG, when conjugated to metallic agents, could be a promising theranostic agent, for treating as well as imaging patients’ tumors.

Diagram 2.3 99mTc-ECG vs. FDG monitoring during the course of therapy in a lymphoma patient. In FDG-PET images, the lesion appears enlarged and fuzzy because of inflammation caused by chemotherapy, while the lesion appears its actual size and has clearer outlines on 99mTc-ECG- SPECT/CT

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2.7 Concluding Remarks

Molecular imaging in oncology has focused on the identification of tumor-specific markers and the application of these markers to evaluate patient response to radiation therapy, chemotherapy, or chemoradiotherapy. The main application of molecular imaging that dominates until now has been the modality known as 18PET/CT, which is intended to help in the evaluation and management of drug dosage for safety and effectiveness. However, 18PET/CT has fallen short to its premise mostly because of the limitations of the tracer drugs to monitor treatment over the course of treatment. More importantly, the radiotracer should have the ability to assess, noninvasively, disease treatment endpoints, which up to now, almost exclusively, rely on the histo-pathological diagnosis of biopsies because of the inflammatory process after

Diagram 2.4 Glucose Metabolism and the Hexosamine Biosynthetic Pathway Link to Key Growth/Survival Signaling Pathways in DLBCL. This diagram depicts the connection between key growth/survial signaling pathways to the metabolic pathway in DLBCL. The hexosamine bio-synthetic pathway give rises to UDP-GlcNAC, which is equivelent to ECG, that modifies key transcription factors, such as NFATc1 and NF-kBp65, allowing these transcription factors to migrate to the nucleus and bind to DNA


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treatments. In order to develop personalized therapies to achieve optimal diagnosis and early cure, the objective is to design metal-based molecular imaging radiotracers that can image the entire repertoire of metabolically active carbohydrate/sugar sub-strates (glycome). Chelator-based imaging technology is the cornerstone for ther-anostic applications, which aim to enable the assessment of target therapies and patient selection. For example, L,L-ethylenedicysteine (EC) is a family of bis-amino-ethanethiol (BAT) tetradentate ligands that are known to form stable 99mTc(V)O com-plexes in which an oxotechnetium core is bound to the thiol-sulfur and the amine-nitrogen atoms. One metal which is being used for detecting cancer which is relatively inexpensive has a long half-life, is easily accessible, and has strong 99mTc which has a complexing property of such N2S2-tetraligand systems that can form label protein linkage or peptide linkage. It has been found that EC is a unique chela-tor because EC has the potential to be involved in signature pathway events. It’s been observed that EC-homing conjugates are able to mimic pathways and monitor changes in the target expression from pre- to posttreatment. Moreover, target-specific biomarkers that are designed as a universal imaging tracer probe, such as ECG, can assess GP, HBP, and broad glycome status from broad to a specific transitional appli-cation in cancer and other metabolic diseases; it is conceivable that the knowledge gained will be helpful to optimize therapies against these disorders.

The imaging agent 99mTc-ECG is already in clinical trials for various cancers and has shown great potential to become the next-generation theranostic imaging agents. The premise of therapeutic and diagnostic capabilities of 99mTc-ECG imaging for refractory DLBCL and other types of metabolically active cancers is already in prime time. Such an important approach should have great potential for clinically translatable advances that can have a positive impact on the overall diagnostic and therapeutic process, which will also enhance the quality of life for cancer patients and other diseases of patients.

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131. Yang DJ, Kim CG, Schechter NR, Azhdarinia A, Yu DF, Oh CS, et al. Imaging with 99mTc ECDG targeted at the multifunctional glucose transport system: feasibility study with rodents. Radiology. 2003;226(2):465–473. doi: 10.1148/radiol.2262011811. PubMed PMID: 12563141.

132. Yang D, Yukihiro M, Yu DF, Ito M, Oh CS, Kohanim S, et al. Assessment of therapeutic tumor response using 99mtc-ethylenedicysteine-glucosamine. Cancer Biother Radiopharm. 2004;19(4):443–456. Epub 2004/09/30. doi: 10.1089/cbr.2004.19.443. PubMed PMID: 15453959.

133. Yang DJ, Kong F-L, Oka T, Bryant JL. Molecular imaging kits for hexosamine biosynthetic pathway in oncology. Curr Med Chem. 2012;19(20):3310–4.

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http://dx.doi.org/10.1089/cbr.2004.19.443


Chapter 3Optimization of Aseptic Production in PET Radiopharmaceuticals for Compliance to the Most Current GMP

Chih-Hao K. Kao

3.1 Introduction

Molecular imaging plays a major role in drug discovery and development because of its ability to quantify drug properties in vivo. Gradually, its potential in revealing molecular abnormalities in disease has prepared itself for an ever important applica-tion in clinical disease management. A crucial part in molecular imaging develop-ment is the scientific design and characterization of a molecular imaging agent. Among the various categories of molecular imaging agents, positron emission tomography (PET) radiopharmaceuticals are the most sensitive, which can also provide target-specific information for biochemical pathways and molecular mech-anisms. For clinical consideration, PET radiopharmaceuticals do not produce detectable pharmacologic effects but provide important information concerning the etiology of various diseases. Therefore, PET radiopharmaceuticals are able to assist in the determination of optimal therapeutic dosing, delineate differential diagnosis between patients, and conceivably predict treatment responses.

These promising outlooks of PET radiopharmaceuticals result in a need for their production in industrial environments, which otherwise would not be possible to achieve in clinical settings. In keeping with this notion, manufacture of PET radio-pharmaceuticals has taken steps toward mainstream pharmaceutical industry. Hence, small-scale radiolabeling performed by experienced chemists in radiochem-istry laboratories or small batch production for a few patients in clinical trials is no longer a common practice. Traditional radiochemists and other PET radiopharma-ceutical developers, who once found good manufacturing practice (GMP) an unfa-miliar subject, were ready to equip themselves with basic indispensable concepts

C.-H.K. Kao (*) Buddhist Tzu Chi Foundation, Hualien Tzu Chi Hospital, Hualien, Taiwan

PET Pharm Biotech Co Ltd, Taipei, Taiwane-mail: [email protected]


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for a potential expansion of their profession into industrial PET radiopharmaceuti-cal production. While terminologies such as quality assurance programs, qualifica-tion and validation, and environmental monitoring are no longer foreign, there are still significant challenges ahead before a fully integrated “PET drug production” is able to claim itself readily compliant to the most current GMP standards. Some of the challenges were too easily deemed “impossible to overcome” or “not applica-ble” for PET radiopharmaceutical before today’s production. However, they were in fact norms to all other types of pharmaceutical production and thus “nonnegotiable” in the eyes of GMP inspectors. Can we still use “small batch volume” and “short half-life” as excuses to circumvent these issues? More often than not, a solution was just not thought out hard enough, and ingenious techniques were eventually created to resolve the problem only upon GMP inspector’s insistence. Nevertheless, PET drug production is progressing and getting its acceptance into pharmaceutical industry. This chapter describes a part of this endeavor that is the most unfamiliar to traditional PET radiopharmaceutical production and provides a few examples on how these could be optimized.

3.2 Trends in PET Radiopharmaceutical Development

For decades, development of PET radiopharmaceuticals centered on molecules labeled with fluorine-18 (18F, half-life, 110 min), carbon-11 (11C, half-life, 20 min), and a few other halogen isotopes with much longer half-lives such as bromine-76 and iodine-124. These radionuclides are cyclotron produced and are incorporated into biochemically active molecules via covalent chemistry. Fluorine-18 radiophar-maceuticals in particular have prevailed due to, among other factors, its more appro-priate half-life. For instance, the technology for producing 18F-fluorodeoxyglucose (FDG) is now so mature that a large quantity of FDG can be obtained by touches of a few buttons, and its quality as a medicinal product meets the highest standard for a sterile injectable solution. FDG is currently the single most widely used PET radiopharmaceutical, which has been so successful in imaging disease with high glycolytic rate and, thus, has become one indispensable tool in the management of many cancers [1]. However, FDG has several limitations, which can give rise to false-positive/false-negative diagnoses and poor predictive value of tumor chemora-diation therapy responses. In addition, certain tumors such as the neuroendocrine type have poor uptake of FDG [2]. Therefore, there is a high demand to develop new radiopharmaceuticals beyond FDG in the oncology field, and indeed many are already in various stages of development. One important feature that sets them apart from FDG and many other PET radiopharmaceuticals of the earlier days is not sci-entific but methodological in terms of how they were developed. Most of these new and promising PET radiopharmaceuticals were being moved forward in industrial

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settings, and their synthesis methods were chosen based on automation capabilities and yields. A pharmacologically favorable PET radiopharmaceutical might never make its way into clinical medicine if a commercially viable production method could not be found.

A similar argument can be made for PET radiopharmaceuticals in the neurologi-cal field. In fact, the first new PET radiopharmaceuticals in decades to be granted marketing authorizations in major countries and regions such North America and Europe are those designed for the detection of abnormal protein plaque deposition in the brain [3].

A bright prospect in new PET radiopharmaceuticals is the development of gal-lium- 68 (68Ga, half-life, 68 min) compounds, which represent an entirely different category, in terms of their chemistry, their production, and supply routine. The generator- produced metallic radionuclide 68Ga is conjugated to a small peptide via coordinating ligands and serves as a convenient alternative to 18F for routine clinical practice. A number of 68Ga-labeled peptides have proven their efficacies in imaging neuroendocrine tumors and prostate cancers [4].

3.3 Regulations and Standards for PET Radiopharmaceutical Manufacture

Regardless of 18F or 68Ga compounds, these short-lived radiopharmaceuticals are characteristically different from nonradioactive drugs and even from other longer- lived radiopharmaceuticals such as iodine-131 capsules and technetium-99 m com-pounded solutions. Nevertheless, they are medicinal products for human use, and, hence, their manufacture must be subjected to proper regulations in order to con-trol their quality. Moreover, because PET radiopharmaceuticals are virtually always solutions for intravenous injection, extremely good practice manufacture has to be guaranteed, to eliminate the risk of product contamination. Since they are usually not able to wait through a longer process of terminal sterilization to com-pletely kill potential contaminating pathogens, aseptic handling procedures are critical during production. Collectively, all these necessary steps make the manu-facture practice of short-lived PET radiopharmaceuticals one of the toughest GMP processes to follow and comply.

An official GMP guidance is usually issued and enforced by individual national governments, and it generally covers the same manufacturing process and facility, quality guidelines, and personnel training. However, some standards and thus the quality of products may vary to an unacceptable degree. To solve this problem, the Pharmaceutical Inspection Convention (PIC) and the Pharmaceutical Inspection Co-operation Scheme (PICS) have created a universally concurred GMP guide for medicinal products. PIC and PICS (PIC/S) are two international instruments, which

3 Optimization of Aseptic Production in PET Radiopharmaceuticals…

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jointly provide an active and constructive cooperation in the field of GMP, between countries and pharmaceutical inspection authorities. Their mission is to lead the international development, implementation, and maintenance of harmonized GMP standards and quality systems of inspectorates in the field of medicinal products (http://picscheme.org/).

With obvious distinct characteristics, radiopharmaceutical production was given a dedicated section, Annex 3, in PIC/S GMP guide. Annex 3 covers all radiophar-maceuticals while offering some exemptions for short-lived PET radiopharmaceuti-cals. The exclusion of the cyclotron portion in the production section of this guide means that “GMP” starts from the point in which the radioactive starting material (like a solution containing 18F) enters into the production process. Additionally, a very rare process of “conditional release” of final product is allowed so that a short- lived product can be released for patient use, before completion of the sterility test. However, being an aseptically processed product, short-lived PET radiopharmaceu-ticals are subjected to one of the most difficult sections of all: Annex 1, for sterile medicinal product, which makes this PIC/S GMP regulation very difficult to comply.

In the United States, a version of GMP is specifically designed for PET radio-pharmaceuticals in Section CFR 21 part 212, which is a cGMP rule for PET drug production. This US version is generally considered easier to comply because it acknowledges that a short-lived PET drug is in fact “short-lived” which implies that the potential harm a contaminated product can cause is considerably limited. The risk it carries and thus the rule for its manufacture should not be comparable to those of just any other sterile medicinal products.

3.4 Example of a Cleanroom Designed Specifically for Aseptic Production of PET Radiopharmaceuticals

The PIC/S GMP guide specifies cleanroom grading with Grade A, as being the cleanest, versus Grade D, designated as the least clean. Table 3.1 lists the limits of particle number permitted in each clean grade. This guide also specifies grading requirements for various types of activities to be conducted. For example, an automatic synthesis module needs only be placed in a Grade C room, but the final filling of a filtered solution needs to be done in a Grade A space, with a Grade B background. The design of a cleanliness-controlled complex for a PET radiophar-maceutical production laboratory (lab) should follow those guides.

A conventional PET radiopharmaceutical production room in a university lab or research hospital usually has multiple hotcells enclosing various synthesis modules. Each module is connected by a dedicated fluid transfer tube that transports a formu-

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lated product to a distant hotcell for dispensing and packaging [5, 6]. This dispens-ing space is responsible for handling multiple incoming formulated products from different synthesizers. Aseptic filtration of formulated products is usually performed in this hotcell as well. This type of space is not too difficult to be transformed into a cleanliness-controlled complex according to the above guidance, as long as the dis-pensing hotcell can be kept in a Grade A room, which is located at Grade B environ-ment, although this is only from the viewpoint of cleanroom grading (Fig. 3.1). There are many other designs that might better suit various purposes in PET radio-pharmaceutical production. As an example, this section describes a new concept in an industrial PET radiopharmaceutical manufacture facility. This site is a central-ized PET radiopharmaceutical provider in Taiwan, and it has been certified accord-ing to PIC/S GMP guide by the Taiwanese authority (Fig. 3.2).

Table 3.1 Maximum permitted number of particles/m3 equal to or greater than the tabulated size

GradeAt rest In operation0.5 pe 5.5 m 0.5 pe 5.5

A 3520 20 3520 20B 3520 29 352,000 2900C 352,000 2900 3,520,000 29,000D 3,520,000 29,000 Not defined Not defined

Fig. 3.1 This is a conceptual drawing of a conventional PET radiopharmaceutical production complex that follows basic cleanroom requirement of PIC/S GMP guidance. A row of three hot-cells for synthesizers is in a Grade C room, while a Grade A dispensing hotcell is in a separate Grade B room. Dotted lines indicate transport tubes from each synthesizer. This drawing does not depict necessary functions such as pass boxes for material in and product out


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Fig. 3.2 This is the actual cleanroom complex in PET Pharm Biotech Co Ltd. (a) Floor plan. Operators enter through left bottom corner (green arrow) into a Grade D corridor. The two produc-tion rooms are assessed via series of small buffer spaces. Only the upper production room is depicted with a hotcell in this drawing. (b) 3D rendition of the floor viewing from the left side of the plan. Crimson color floor is Grade D while gray C and yellow B. The final gowning and two production rooms are in Grade B. Two rooms with pass boxes in Grade C are for passing of materi-als and products. (c) Viewing from the opposite angle

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This alternative design argued that multiple small Grade B rooms, each setup with a hotcell encasing one Grade B synthesis compartment and one Grade A asep-tic filling compartment, would be a more efficient way for complying the toughest GMP, especially in a site that more than one kind of PET radiopharmaceuticals are regularly produced and dispensed.

Entering via a series of small rooms that increase in the cleanliness grades, the Grade B final gowning room opens to two Grade B production rooms. In each production room (12 m2) stands one hotcell with two main compartments. The syn-thesis compartment is not controlled for its air particles to a specific standard, but because it opens to the production room, and thus it is in the same level of cleanli-ness as the production room itself. The other compartment of the hotcell is for asep-

Fig. 3.2 (continued)


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tic filling; it is therefore designed to be an isolator of Grade A. Inside this isolator, particle counts are monitored continuously and air is sampled for bacteria.

The concerned cleanroom complex is only 38 m2 including buffer, final gowning, raw materials and products in and out, and two production rooms. There is a system established to monitor temperature, humidity, radiation dose rate, and air pressure in each room. The ventilation control system has chambers to collect waste gas as well as to vent through hotcell. With each production room only 12 m2, one hotcell, and maximum two operators, Grade B is easily maintained. The regular and fre-quent disinfection practice required for the Grade B room is also achieved without difficulty. Within hotcell the two compartments are adjacent, and thus the product solution transport line is short enough to allow the use of the prepackaged dispos-able sterilized tubing. More importantly, two production rooms can be in operation simultaneously with diminished risk of cross contamination. The government audi-tors have deemed this design in compliance with PIC/S GMP requirements.

Cross contamination among different drug products is one of the major issues that PIC/S GMP requirements designed to avoid. Yet, in an effort to comply with these rules, most PET radiopharmaceutical manufacture facilities still set up their production rooms with multiple hotcells, so that various PET radiopharmaceuticals could be produced on the same day even simultaneously. To make the matters worse, these different products often share the same aseptic filling space. Though the syn-thesizer room is only required to maintain a Grade C cleanliness, the Grade A asep-tic filling still needs to be stationed in a Grade B environment, which requires a preceding Grade B gowning room as depicted in Fig. 3.1. The overall use of space is not as economical as one might think, and it would be hard to convince GMP auditors that cross contamination would be unlikely.

This “one room, one product” concept simplifies the work flow and allows two rooms in operation simultaneously. No different batches of PET radiopharmaceuti-cals will be processed in the same hotcell space, which greatly lowers the risk of cross contamination. Though now the synthesis module is set up in Grade B envi-ronment, this only enhances the quality of production. In conclusion, this clean-room design is effective and in compliance with the most current and toughest GMP requirements.

3.5 Challenges in Aseptic Process in Short-Lived PET Radiopharmaceutical Manufacture

Short-lived PET radiopharmaceuticals such as FDG are usually produced as a batch of solution in small volume. The volume of the original solution coming out of a dedicated FDG synthesizer can be as small as 12 mL, but it is usually around 20 mL. The product solution is filtered to achieve sterility by flowing from the synthe-sizer through a filter-connected tube into a batch vial. After its final closure, the vial is properly packaged and sent out as a final product to a radiopharmacy, where it is

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subsequently dispensed into a unit-dose syringe for each individual patient. This final product package is called a multidose vial because it usually contains more than enough FDG for a single-patient dose. Alternatively, one can dispense into multiple unit-dose syringes filtered product solution directly. A filled syringe is then the product’s final package. Irrespective of what final product packaging is chosen for one manufacture site, this process has to be performed in Grade A compartment and handled aseptically, which means following the Annex 1 of the PIC/S guide in all details.

In this Grade A hotcell compartment that is validated for HEPA functions, air-flow speed and pattern, air pressure differential, temperature, and humidity are merely basic requirements. Aseptic processing for manufacturing a sterile product has many other obligatory functions that demand real challenges.

Design Qualification (DQ) for a Dispensing Hotcell DQ is a process to be done prior to the better known 3Q (IQ, OQ, and PQ). It is actually a requirement in many parts of the GMP but quite often neglected. One, however, would find oneself feel-ing stranded if DQ was neglected for a dispensing hotcell. Here are non-exhaustive examples explaining this situation. A hotcell needs to have certain built-in features in order to accommodate certain obligatory functions, which in turn show how important the DQ is for this hotcell:

1. Continuous monitoring of particle counts during aseptic processing is nonnego-tiable and thus needs to be done. A built-in particle counter in the hotcell cer-tainly fulfills the demand, but an off-the-self hotcell is not likely to have this feature. Alternatively, a commercial particle counter with a long particle- collecting tube might also serve the purpose, if only the tubing could penetrate through the hotcell wall.

2. Continuous air sampling is another similar situation because air exhaust from the sampler needs to go outside the hotcell.

3. Fumigation by disinfectant such as hydrogen peroxide before each production, which usually means at least daily for PET radiopharmaceutical, is another requirement. A normal hotcell interior material and fixtures usually do not expect an environment as harsh as this, and things break down much easier than usual.

Sampling for Bioburden Analysis “Bioburden” is a term that describes microbio-logical burden brought by an unsterilized product. To analyze this is to understand sources of microbiological contamination for a specific production process. In order to do this, one must take a sample from the product before it is being subjected to sterilization and microbiological analysis, which is subsequently tested. This proce-dure must be performed for each and every product batch produced.

PET radiopharmaceuticals such as FDG are usually produced by an automatic module, and the process includes radiochemical reactions and product purification and formulation, which are performed in a close tube system all the way to sterile filtration and until the product solution is dispensed into final the container. Now in order to take a sample for bioburden test, this “nonstop and all-the-way flow through


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in a close tube” system could suddenly become impossible. More difficult still is the requirement of minimum volume sampling that usually is not less than 10 mL, whereas the batch of product might only be 20 mL at best. Achieving such task is not trivial and it will really put one’s ingenuity to the test. It is not practical to pro-vide any answer here as manufacture sites vary in their synthesis module, filtration apparatus, and many other configurations. Yet readers need to consider these issues in the design of their own production flow.

3.6 Conclusion

The design of most current GMP facilities for PET radiopharmaceutical manufac-ture around the world has gradually gained momentum during the past decade. While basic requirements in premises, equipment, personnel, documentation, and operations are no longer headaches to some, details abound with difficult issues for most involved. Perhaps, regulatory auditors may finally yield on a few things based on low-risk factors. This chapter merely initiates the process by pointing out challenges by presenting some examples. It is the management and operators every-where who would eventually provide specific solutions to optimize the respective facilities.

References

1. Delbeke D, Coleman RE, Guiberteauetal MJ. Procedure guideline for tumor imaging with 18F- FDG PET/CT 1.0. J Nucl Med. 2006;47(5):885–95.

2. Eriksson B, Bergström M, Orlefors H, Sundin A, Oberg K, Långström B. Use of PET in neu-roendocrine tumors. In vivo applications and in vitro studies. Q J Nucl Med. 2000;44(1):68–76.

3. Sabri O, Sabbagh MN, Seibyl J, Barthel H, Akatsu H, Ouchi Y, et al. Florbetaben PET imaging to detect amyloid beta plaques in Alzheimer’s disease: phase 3 study. Alzheimers Dement. 2015;11(8):964–74.

4. Sharma P, Singh H, Bal C, Kumar R. PET/CT imaging of neuroendocrine tumors with 68Gallium-labeled somatostatin analogues: an overview and single institutional experience from India. Indian J Nucl Med. 2014;29(1):2–12.

5. Kao CHK, Hsu WL, Xie HL, Lin MC, Lan WC, Chao HY. GMP production of [18F]FDOPA and issues concerning its quality analyses as in USP “Fluorodopa F 18 Injection”. Ann Nucl Med. 2011;25(5):309–16.

6. Kao CHK, Hsu WL, Kao PF, Lan WC, Xie HL, Lin MC, et al. An efficient and aseptic preparation of “sodium fluoride (18F) injection” in a GMP compliant facility. Ann Nucl Med. 2010;24(3):149–55.

C.-H.K. Kao

Part IIEngineering and Instrumentation Aspect


Chapter 4OpenPET Enabling PET Imaging During Radiotherapy

Taiga Yamaya and Hideaki Tashima

Abstract The OpenPET is the world’s first open-type 3D PET scanner for PET image-guided particle therapy such as in situ dose verification and direct tumor tracking. The key technology, which enabled the OpenPET realization, is a depth- of- interaction detector. Even with a full-ring geometry, the OpenPET has an open gap between its two detector rings through which the treatment beam passes. Following the initial proposal of the dual-ring OpenPET (DROP), a small prototype was developed to show a proof of concept. The single-ring OpenPET (SROP) was also proposed as a more efficient geometry than DROP in terms of manufacturing, cost, and sensitivity. A small SROP prototype was developed, and feasibility of visualizing a 3D distribution of beam stopping positions inside a phantom was shown with the help of radioisotope particle beams, used as primary beams. Following these results, a full-size whole-body DROP prototype was developed.

Keywords PET • Particle therapy • Depth of interaction (DOI) • Image-guided radiotherapy • IGRT

4.1 Introduction

Cancer is a major cause of death in developed nations, and the disease death rate continues to increase. Therefore, many efforts have been made toward better diag-nosis and better treatment. Among diagnostic tools, positron emission tomography (PET), which can visualize functions such as metabolism by injecting radioisotope tracers labeled with positron emitters, is expected to enable earlier and more precise cancer diagnosis. However, big improvements of PET could still be made, including spatial resolution, sensitivity, and manufacturing costs. For example, the sensitivity of present PET scanners does not exceed 10%, and thus, more than 90% of the gamma rays emitted from a patient are not utilized for imaging.

T. Yamaya (*) • H. Tashima National Institute of Radiological Sciences (NIRS), National Institutes for Quantum and Radiological Science and Technology (QST), Chiba, Japane-mail: [email protected]; [email protected]



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A depth-of-interaction (DOI) detector, for which various positioning methods have been studied [1–7], will be a key device to get any significant improvement in sensitivity while maintaining high spatial resolution (Fig. 4.1). In order to maintain enough detection efficiency, the scintillation crystals should be 2–3 cm long. In con-ventional detectors, the crystal thickness causes uncertainty in position identifica-tion, which results in degraded spatial resolution at the peripheral area of a field of view (FOV). On the other hand, the DOI detector can reduce the parallax error while maintaining the efficiency.

Figure 4.1b shows a four-layered DOI detector based on a light sharing method [8, 9], one of the successful proofs of concept made for the jPET-D4, which is a brain prototype PET system with the DOI detectors. It has a nearly uniform spatial resolution of around 2 mm all over the FOV, which was obtained using iterative image reconstruction with the geometrically defined system matrix [10]. This detec-tor has been upgraded to get better spatial resolution at lower production costs. Moreover, the successful identification by the 32 × 32 × 4 array of LYSO crystals (1.45 × 1.45 × 4.5 mm3 in size) with a 64channel flat panel PMT (H8500, Hamamatsu Photonics K.K., Japan) [11] is the base for Shimadzu’s new positron emission mam-mography products [12].

For treatment, on the other hand, radiotherapy is essential for effective cancer treatment with minimized side effects. Specifically, particle therapy such as proton and carbon ion therapy are expected to be the ultimate radiotherapy because they can concentrate the dose even in a deep tumor. Thus, there has been remarkable progress in PET and radiotherapy, but no one has looked into the great potential to be obtained by the combination of both.

Fig. 4.1 Shows the comparison between a conventional PET detector (a) and the depth-of- interaction (DOI) detector (b). The DOI detector eliminates the parallax error, caused by the thick-ness of the crystals in conventional detectors

T. Yamaya and H. Tashima

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An open-type PET geometry (OpenPET) is a new idea to visualize a physically opened space between two detector rings [13], which will move researchers toward a future joint PET imaging and radiotherapy system. OpenPET is expected to achieve in-beam PET, which is a method to monitor in situ charged particle therapy. Without injecting any PET tracer, positron emitters are produced through fragmen-tation reactions between the projectiles and the atomic nuclei of the tissue during patient irradiation [14]. Compared with conventional radiation therapy, charged par-ticle therapy can highly concentrate the dose in a tumor. This means if there is any difference between the actual irradiation and the treatment plan, the tumor treatment will be compromised, and the normal tissue around the tumor will be damaged. Therefore, quality of treatment must be assured for in-beam PET.

Dual-head PET cameras have been developed at the Gesellschaft für Schwerionenforschung (GSI), Darmstadt, Germany [15], and at the National Institute of Radiological Sciences (NIRS), Chiba, Japan [16]. They have also been applied to proton therapy [17, 18]. However, they are basically limited to 2D imag-ing, in which the resolution along the axis vertical to the detector faces is insuffi-cient [19]. Therefore, the OpenPET is expected to be the first practical 3D geometry that does not interfere with the beam paths (Fig. 4.2).

4.2 Dual-Ring OpenPET (DROP)

4.2.1 Theory

4.2.1.1 Geometry

The long and continuous FOV including a 360-degree opened gap between two detector rings can be visualized by fully 3D image reconstruction from all the pos-sible lines of response (LORs). Based on the redundant characteristics of 3D PET data, oblique LORs between two separated detector rings are expected to compen-sate for the missing LORs in the gap, and hence, there is a limited width range in the gap.

Fig. 4.2 Provides examples of a PET geometry: (a) a typical cylindrical PET geometry, (b) the OpenPET geometry, and (c) a dual-head PET geometry which has been proposed for in-beam PET

4 OpenPET Enabling PET Imaging During Radiotherapy

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Figure 4.3 illustrates a dual-ring OpenPET (DROP) geometry of two detector rings of diameter D and axial length W, separated by a gap G. When G>W, the FOV becomes axially discontinuous. In addition, the minimum diameter of the FOV is represented as

D

DW

G W¢ =

+( ).

(4.1)

This means that G must satisfy

D

DW

G W0 £ +( ),

(4.2)

when an FOV of more than D0 in diameter is required. Therefore, in order to obtain the axially continuous FOV of D0 in diameter, the range of G is limited to

G W

DW

DW£ -æ

èç

ö

ø÷min .,

0 (4.3)

Fig. 4.3 Shows the geometrical analysis of the possible range of the gap G for the DROP geom-etry of two detector rings of diameter D and axial length W. (a) is the case when G is larger than W, and (b) is the case G = W or shorter. D’ is the minimum diameter of the FOV


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4.2.1.2 Image Reconstruction

DROP becomes practical with iterative image reconstruction methods since image reconstruction of DROP is an analytically incomplete problem. Therefore, a 3D ver-sion of the ordered subset expectation maximization (OS-EM) [20] with a precom-puted system matrix was employed. The basic software was originally developed for the jPET-D4 [10]. The system matrix is designed taking into consideration the geometrical arrangement and the radiation penetration of the crystals. Notably, it is possible to compress the system matrix by eliminating zero elements and applying rotational symmetry and an axial shift property of the crystal arrangement pre- calculation of the system matrix. For the DROP geometries with a variable open space, however, it is not practical to precompute and store a set of system matrixes for each size of open space. Therefore, we proposed a “masked” OS-EM for the DROP geometries with the following conditions:

• The crystals are arranged axially at a regular interval of Δz.• The gap width, G, is divisible by Δz.• The detector response functions are only dependent on the ring differences and

are axially shift-invariants.

Next, a long gapless scanner having N = 2NW + NGmax rings was modeled, which completely covers all possible axial FOVs. Here Nw and NGmax are the numbers of crystal rings, which satisfy W = Nw Δz and Gmax = NGmax Δz, where Gmax is the maxi-mum gap width to be supported. Then we implemented the “masked” OS-EM by applying a mask to the original gapless system, as

ff

w aw a

g

w a fjn m j

n m

l l lj i S mi ij

i

k i ik k

,,

( )

,+

Î

=å å å

1

(4.4)

where fjn,m is the j-th image element in the m-th sub-iteration of the n-th iteration. gi

and aij are the data element and the system matrix element for the N-ring scanner, respectively. S(m) represents a subset of LORs which are used in the m-th sub- iteration. The mask {wi} was defined as

w

ii =

-ìíî

0

1

( )

( ).

when the thLORismissingbythegap

else (4.5)

4.2.2 Simulation

4.2.2.1 Scanner

In order to evaluate the imaging performance of the DROP geometries, we simu-lated a DROP of the two detector rings (D = 827.0 mm, W = 153.6 mm) separated by a variable gap (G ranging from 0 to 2 W) as shown in Fig. 4.4a. The maximum


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gap size to have an axially continuous FOV is W. However, in order to investigate the image quality when the gap exceeded the limitation, we varied the gap size up to 2 W. Each one of the dual scanners consisted of 4 rings of 72 block detectors. Each block detector consisted of 8 × 8 array of BGO crystals with dimensions of 4.1 mm (transaxial) × 4.4 mm (axial) × 30.0 mm (depth). The axial crystal interval was Δz = 4.8 mm (center to center), and the number of crystal rings for each one of the dual scanners was NW = 32.

Figure 4.5 shows the minimum diameter of the FOV of the simulated DROP as a function of the gap according to Eq. (4.1). The possible range for the gap to have the axially continuous FOV is also shown according to Eq. (4.3). At the maximum gap for the axially continuous FOV, i.e., G = W, the minimum diameter of FOV is decreased to 413.5 mm.

To cover the maximum gap of 307.2 mm, i.e., NGmax = 64, the long gapless scan-ner of N = 128 crystal rings was simulated. The voxel size was 2.4 mm3 and the “mash” factor of two was applied, i.e., two adjacent views were averaged to reduce angular samplings. Then, with the restricted FOV of 268.8 mm in diameter, the dimensions of the projection data became 120 bins × 144 views × 1282 ring pairs. Projection data were generated by the forward projection using the system matrix. Attenuation and scattering were not included in the simulation.

4.2.2.2 Phantom

In order to demonstrate the imaging performance of the DROP geometry, we simu-lated a “warm” cylinder of 230.0 mm in diameter and 614.4 mm length, which included 63 “hot” spheres of 4.0 mm diameter (Fig. 4.4b). The warm cylinder was

Fig. 4.4 Depicts the simulated DROP scanner (a) and the simulated cylinder phantom (b). Each detector ring of the DROP has similar dimensions to the HR+ (32 crystal rings, 153.6 mm axial length)


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filled with background activity and the hot spheres contained higher activity. The contrast between background and the hot spheres was 1:5. The mask for the gaps of G = 0.0 mm, 76.8 mm, 153.6 mm, 230.4 mm, and 307.2 mm was applied to the noise-free projection data. The masked 3D OS-EM with 8 subsets and 20 iterations was applied.

4.2.2.3 Results

Reconstructed images are shown in Fig. 4.6. Since the coronal and the sagittal views are identical, only the sagittal views are displayed. In addition, two transaxial slices at the center and at 114.9 mm off-center are shown. The central slices (slice A) of G ranging from 0.0 to 307.2 mm were very similar. The off-center slices (slice B) of G ranging from 0.0 to 153.6 mm were also similar. However, when G>153.6 mm (i.e., the maximum limitation for axially continuous FOV), the blank areas where there is no LOR appeared on both sides of the open space. With the blank areas, distorted point spread functions and low-frequent artifacts were also observed.

The simulation results showed that the DROP visualized a long FOV including the open space. However, it is necessary to minimize the open space according to applications because there is a trade-off relationship between image quality and the size of the open space. In the simulation, artifacts including distorted point spread functions started to appear on both sides of the open space when the gap exceeded 153.6 mm, i.e., the maximum limit for the axially continuous FOV.

4.2.3 Development of a Small Prototype

4.2.3.1 Prototype

We developed a small prototype to show a proof of concept of the OpenPET imag-ing [21]. The prototype was designed as a compact system (Fig. 4.7) so that it can be easily transported between PET areas and therapy areas such as the Heavy Ion

Fig. 4.5 Reveals that the relation between the gap G and the minimum diameter of FOV D’ for the simulated DROP. NG is the number of truncated rings, which satisfies G = NG Δz


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Medical Accelerator in Chiba (HIMAC) [22]. Two detector rings of 110 mm diam-eter composed of 8 block detectors were placed with a variable gap. The height of the center of the detector ring was 125 cm, which equals the height of the HIMAC beam line. Power supply devices and the data acquisition (DAQ) system including positioning and coincidence circuits were placed in the base rack. Details of the DAQ system are available in [23]. In order to reduce radiation damage to electronic circuits caused by secondary particles such as neutrons, front-end circuits such as preamplifiers were separated from photomultiplier tubes (PMTs) and connected by 120 cm long coaxial cables.

Each block detector, which has four-layer DOI capability [11], was composed of a 14 × 14 × 4-layer array of 2.9 × 2.9 × 5.0 mm3 Lu1.8Gd0.2SiO5:Ce (LGSO) crystals (Hitachi Chemical Co., Japan) and a H8500 PMT. Axial length of the crystal blocks W was 42 mm. The gap G (i.e., axial length between crystal blocks) was also adjusted to 42 mm, which was the maximum value possible to obtain the axially continuous FOV. The real gap between the two gantry ends of two detector rings

Fig. 4.6 Shows reconstructed images of the simulated warm cylinder phantom with hot spheres for the DROP (G ranging from 0.0 to 307.2 mm). In addition to sagittal views, the central slice (slice A) and the off-center slice (slice B) are shown. White dotted lines represent the boundary of the area where detectors are located


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was limited to 27 mm by the gantry structure, but the gap will be over 20 cm, which is enough for radiation therapy, if the gantry is extended to the scale of a human body. The energy window was 400–600 keV. At this development stage, we have not completed the timing correction, and the coincidence time window was 60 ns.

As the image reconstruction algorithm, the 3D OS-EM algorithm [20] was applied. Acquired list-mode data were transformed to histogram data before image reconstruction. Random correction was applied by subtracting delayed coincidence, but attenuation correction and scatter correction were not applied. System matrix elements were calculated based on the geometrically defined detector response functions [24]. Iteration parameters were 8 subsets and 10 iterations. The FOV defined in the image reconstruction was 114 mm in diameter and 126 mm in axial length. Two choices of voxel size, (1.5 mm)3 and (0.75 mm)3, were applied. The major specifications of the small DROP prototype are listed in Table 4.1.

4.2.3.2 Imaging Performance

Absolute Sensitivity We obtained the absolute sensitivity at the center of the FOV by measuring a 0.047 MBq 22Na point source for 10 min. In order to evaluate loss of sensitivity due to the gap, the absolute sensitivity at the center of the FOV was measured by reducing the gap to the minimum as G = 42 mm, 33 mm, 24 mm, and 15 mm. Then the absolute sensitivity for G = 0 mm was estimated using an extrapo-lation (Fig. 4.8). The sensitivity for G = 0 mm is estimated at 8.7% by linearly extrapolating measurement data, while the sensitivity for G = 42 mm is 6.6%.

Fig. 4.7 Depicts the small DROP prototype: (a) a photograph of prototype and (b) sketches show-ing the positions of the detectors, front-end circuits, data acquisition (DAQ) system, and power supply


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The results comparing the gap case (G = 42 mm) and the non-gap case (G = 0 mm) clearly show what we have sacrificed for the open geometry: about 24% of the sen-sitivity is lost due to the solid angle effect.

Phantom Imaging For demonstration, we measured a small rod phantom, which had a similar structure to a micro-Derenzo phantom (Fig. 4.9). The phantom con-sisted of an outer hollow cylinder and an inner solid cylinder of 36.1 mm diameter and 17.8 mm length. The inner solid cylinder contained rods of various diameters (1.0 mm, 1.6 mm, 2.2 mm, 3.0 mm, 4.0 mm, and 4.8 mm). First, the phantom was placed vertically with the rods parallel to the axial direction (“vertical” position), and then, it was placed horizontally (“horizontal” position). In addition, the phan-tom was placed in two positions: in-ring and in the gap. The phantom (filled with 1.3 MBq 18F solution) was measured in the order of in-ring vertical, gap vertical, in-ring horizontal, and gap horizontal. Scan times, which were adjusted to correct for radioactive decay, were 20 min, 23 min, 28 min, and 34 min, respectively. With the random rate varying from 15 to 6%, the measured prompt coincidence counts were 14.1 M, 29.7 M, 21.2 M, and 28.3 M, respectively. The 0.75 mm voxel was used in image reconstruction. The comparison of transaxial slices obtained in-ring

Table 4.1 Major specifications of the developed small DROP prototype

Scanner Crystal material LGSO Crystal size 2.9 × 2.9 × 7.5 mm3

Crystal array 14 × 14 × 4-layer PMT 64ch flat panel PMT (Hamamatsu H8500) Detector arrangement 8 blocks × 2 rings Ring diameter 110 mm Gap between crystal blocks (G) 42 mm (minimum 15 mm)Image reconstruction Algorithm 3D OS-EM Detector response function Geometrically defined Voxel size (1.5 mm)3 or (0.75 mm)3

Fig. 4.8 Reveals the absolute sensitivity at the center of the FOV with varying the gap. The sensi-tivity for G = 0 mm was estimated with the linear extrapolation


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(Fig. 4.9a) and in the gap (Fig. 4.9b) supports the feasibility of the OpenPET imag-ing since almost the same image is obtained even in the gap. The image in the gap seems to be rather superior to that in the ring, because sensitivity in the gap is about twice as high as that in the ring. Rods of 2.2 mm diameter are completely separated. On the other hand, the effect of the loss of axial spatial resolution is seen in the comparison of coronal slices obtained in-ring (Fig. 4.9c) and in the gap (Fig. 4.9d), although the rods are still separated in the gap.

4.2.3.3 Small Animal Imaging

Colon-26 cancer cells (1.0 × 106) were inoculated subcutaneously into both flanks of a female BALB/c-nu/nu mouse (16.5 g weight, Japan SLC Inc., Japan). After 10 days, the mouse was fasted overnight before PET scanning. Then 1.6 MBq 18F-labeled fludeoxyglucose (FDG) was injected intravenously via a tail vein. After 70 min, the animal was sacrificed and placed so that the tumors were located in the gap. Then 10.3 M prompt coincidence counts and 0.6 M delayed coincidence counts were measured for a 10 min PET scan. The 0.75 mm voxel was used in image recon-struction. A reconstructed image is shown in Fig. 4.10. The tumors located in the

Fig. 4.9 Shows the reconstructed images (0.75 mm voxel) of the small rod phantom at four differ-ent positions (a–d)


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gap are clearly visualized. This result also shows the effect of the extension of the axial FOV. An axial FOV of 126 mm is obtained with the detectors that originally cover only an 84 mm axial FOV.

4.2.4 In-Beam Imaging of Carbon Ion Beam

For a proof of concept of PET imaging during radiation therapy, our group carried out in-beam experiments in the HIMAC to visualize in situ particle distribution in a phantom. In a usual carbon (12C) beam irradiation, it has been reported that the activity of positron emitters, produced by fragmentation reactions, is generally low in addition to the theoretical difference between the dose distribution and positron emitter distribution [14]. Therefore, instead of the 12C beam, we used a 11C radioac-tive (RI) particle beam as an incident beam, directly [25] in the secondary beam port. This beam port is where various RI beams can be generated as secondary beams from the 12C beam by irradiating a Be target. Because the projectiles them-selves are positron emitters, we expected to obtain PET images directly correspond-ing to the distribution of primary particles.

Fig. 4.10 Shows the results of small animal imaging. A reconstructed image (0.75 mm voxel) of the tumor inoculated mouse using 18F-FDG. The figure reveals coronal and transaxial slices cross-ing two tumors (indicated by the arrows)


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4.2.4.1 Experiment Setup

Figure 4.11 shows the experimental setup. The small DROP prototype was posi-tioned so that the beam passed through the gap. A polymethyl methacrylate (PMMA) cylindrical phantom (40 mm diameter, 100 mm length) was placed in the center of the FOV and irradiated along the radial direction by a 11C pencil beam. The beam intensity was about 5 × 106 particles per second (pps), which was the maximum possible in the secondary beam port, and the beam energy was 332.6 MeV/u. Therefore, a PMMA range shifter of appropriate thickness (150 mm) was used to stop the beams near the center of the phantom. After passing beams through the brass collimator (5 mm diameter, 150 mm length), the beam intensity was limited to 1.7 × 105 pps. It has been reported that PET data, which are measured at the moment of irradiation, include increased random coincidences by prompt gamma rays [26, 27]. On the other hand, the irradiation in the HIMAC has a 3.3 s cycle, in which the particles are delivered as a “bunch” of about 1.8 s duration followed by about 1.5 s pause. Hence, for the in-beam PET measurements, we extracted a useful list-mode data for each irradiation and synchronized it with this bunch cycle.

As a demonstration of the range resolution performance, an additional PMMA board of 5 mm thickness was used to cover half of the collimated beam, as shown in Fig. 4.11c. The phantom was irradiated by the 11C beam for 20 min and in-beam PET data were measured during irradiation. The voxel size in image reconstruction was also 1.5 mm.

Fig. 4.11 Shows the experimental setup for in-beam imaging tests of the developed small DROP prototype, installed in the HIMAC. (a) is a schematic top view illustration showing details of the photograph in (b). Insertion of an additional 5 mm thick PMMA board (c) is indicated by the arrow, which covers bottom half the collimated beam


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4.2.4.2 Results

Reconstructed images from in-beam PET data during the 20 min 11C irradiation are shown in Fig. 4.12. The 5 mm difference in the range was clearly visualized both in the coronal slice and the transaxial slice, which supports our conclusion that at least a 5 mm resolution in range could be detected three-dimensionally, with in-beam measurement by the DROP with 11C irradiation.

4.2.5 Real-Time Imaging for Tumor Tracking

An image reconstruction calculation usually requires a time period in the order of minutes, but there are big demands in real-time imaging for OpenPET. For example, real-time PET-guided medical procedures such as radiation therapies and biopsies are feasible by visualizing and tracking a target tumor, which is labeled with radio-active tracer in addition to the in-beam imaging. It should be noted that tumor tracking is a much more challenging application in terms of the real-time require-ment. Real-time tracking is desirable, especially for mobile organs such as the lungs, which move with respiration. In contrast to conventional tumor tracking methods, which use implantation of surrogates, the real-time OpenPET imaging enables the direct visualization of the tumor. Real-time imaging of PET is a chal-lenging task because of its computational burden in image reconstruction. Conventionally, PET images are reconstructed after all data are acquired. The time to display an image takes at least several minutes. Thus, in order to perform

Fig. 4.12 Presents a schema of the resolution performance for a 5 mm difference in range. (a) represents reconstructed images (1.5 mm voxel) from in-beam PET data during 20 min 11C irradia-tion with the PMMA board of 5 mm thickness covering the right half, while (b) is the bottom half of the collimated beam


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real-time imaging, a cycle of data acquisition, data transfer, image reconstruction, and display needs to be completed within less than 1 second. For the real-time image reconstruction, we implemented the 3D one-pass list-mode dynamic row-action maximum likelihood algorithm (DRAMA) [28, 29] on a graphics processing unit (GPU). In addition, a data transfer control system was implemented, so that the reconstruction speed is stabilized by limiting the number of list-mode events pro-cessed at one time [30].

4.2.5.1 Experimental Setup

We demonstrated the system performance in terms of the real-time tracking ability by a point source tracking experiment. Figure 4.13 shows the experimental setup and an image frame captured during the demonstration. In the demonstration, a 22Na point source (1 MBq) was put on a moving stage with a long bar so that the source was placed between the gaps of the small DROP prototype. The number of counts to be used in the GPU-implemented 3D list-mode DRAMA was limited below 5000, and voxel size was set to 3.0 × 3.0 × 3.0 mm3 for better frame rate perfor-mance. The reconstructed image size was 38 × 38 × 42 voxels.

Fig. 4.13 Depicts the tracking ability demonstration setup: (a) is the schematic illustration of the top view; (b) reveals the side view; (c) shows a photograph of the demonstration setup; and (d) provides example images captured during the tracking demonstration. 22Na point source (1 MBq) was set between the ring gap and moved up and down with a 30 s cycle


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4.2.5.2 Results

In the demonstration, transaxial, coronal, and sagittal slices of the reconstructed OpenPET images and movies of a moving point source captured by optical cameras were simultaneously displayed on the PC screen (Fig. 4.13d). The point source moved up and down with a 30 s cycle sine curve, and the real-time viewer showed the reconstructed images with a frame rate of 2 frames per second (fps). Figure 4.14 shows the displayed point source position on the screen. From the sine curve fitting, an average delay time of 2.1 s was obtained and the mean error was 2.0 mm. The time-delay correction method using a supporting sensor was also developed in order to compensate for the delay [31].

4.3 Single-Ring OpenPET (SROP)

For the purposes of in-beam PET and tumor tracking, the FOV that we need to image appears only in the gap area between the two detector rings. It is not neces-sary to image the FOV in the detector ring area for these applications. Therefore, our group proposed a second-generation OpenPET geometry, single-ring OpenPET (SROP), as a more efficient geometry [32]. The SROP can provide an accessible and observable open space with higher sensitivity and a reduced number of detectors compared to the previous generation geometries. Figure 4.15a shows a conceptual illustration of the SROP comparing with the first-generation OpenPET; 4.15b depicts DROP, and 4.15c displays an in-beam geometry, in which the conventional PET is positioned at a slant angle relative to the bed to form an accessible space

Fig. 4.14 Highlights displayed positions of the point source in the images obtained by the optical camera and by the OpenPET during the tracking demonstration. From the sine curve fitting, the delay was determined as 2.1 s


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(hereinafter referred to as “slant PET”). Compared with the slant PET, the single- ring OpenPET is expected to provide higher sensitivity with a smaller number of detectors. Thus, in the case of in-beam PET, it provides a closer positioning of the beam port, which minimizes beam broadening. Other possible applications of the SROP are shown in Fig. 4.16.

4.3.1 Theory

The sensitivity of PET is highly correlated with solid angle coverage where the coincidence detection is possible; the solid angle coverage is hereinafter referred to as geometrical sensitivity. On the other hand, the detector surface area is correlated with the number of detectors used for constructing the actual system. Figure 4.17 illustrates geometries of SROP, DROP, and slant PET. The open space with the size of C is achieved with the patient bed width of B and the detector ring width of W.

Fig. 4.15 Shows a conceptual illustrations of PET geometries with an accessible open space to the patient: (a) reveals that the SROP has the shape of a cylinder cut at a slant angle; (b) that the beam port can be placed near the patient bed, the same as in DROP; and (c) that conventional cylinder PET can also provide an open space by placing it at a slant angle relative to the patient bed

Fig. 4.16 Presents the applications of the SROP: (a) as a simultaneous PET-CT multimodality system and (b) as a as a PET-guided biopsy and medical treatment


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The geometrical sensitivity R and the detector surface area S for each geometry are calculated as follows:

SROP

RW C W

W C W B

W C W

W CSROP =

+ +

+ + +-

+ -

+ò

1

4 0

2

2 2pf

f

fp( )sin

{( )sin }

( )sin

{( ))sin },

ff

- +

ìíï

îï

üýï

þïW Bd

2 2

(4.6)

S WBSROP = p . (4.7)

DROP

RW C

W C B

C

C BDROP =

+

+ +-

+

2

2 2 2 2 2( ),

(4.8)

S WBDROP = 2p . (4.9)

Slant PET (SP)

R

W

d WSP =

-

æ

èç

ö

ø÷sin arctan ,

2 2

(4.10)

S W d WSP = -p 2 2 , (4.11)

Fig. 4.17 Depicts the geometrical analysis of the SROP (a), DROP (b), and slant PET (c). The widths of the open space, detector ring, and patient bed are notated as C, W, and B, respectively


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where d is determined by solving the following system of equations:

d C b B

d C B W b

= + +

= + + -( )

ìíï

îï

( ).

2

2

2 2

2 2 22

(4.12)

We note that an additional margin is required to place the beam port in the slant PET case. The margin M is calculated using the following equation:

M W b= -2 2 , (4.13)

where b is determined by solving Eq. (4.12).Figure 4.18 shows the geometrical sensitivity at the center plotted against the

detector surface area. As fixed parameters, the bed width of B = 600 mm and open gap of C = 300 mm were selected to provide a sufficient open space during ion beam therapy. The number of block detectors is defined as the surface area of each geom-etry divided by the area of one block detector, which is also shown above the graph to give an intuitive insight. Here, we assumed that the area of the block detector was 2500 mm2 (i.e., 50 mm × 50 mm). When the detector surface area was 500,000 mm2, in which the equivalent number of block detectors was 200, the geometrical sensi-tivities of the SROP, DROP, and slant PET were 27.8%, 23.9%, and 22.1%, respectively.

Fig. 4.18 Shows the geometrical sensitivity at the center against the number of block detectors for the open gap length C = 300 mm and bed size B = 600 mm. When the detector surface area was 500,000 mm2, in which the equivalent number of block detectors was 200, the geometrical sensitivities of the SROP, DROP, and slant PET were 27.8%, 23.9%, and 22.1%, respectively


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4.3.2 Simulation

In contrast to the DROP, the SROP has a FOV focused at the open space. Therefore, the SROP can provide a wider open space and higher sensitivity at the center of the FOV with the same number of detectors. However, realization of the SROP with rectangular block detectors remains a challenging problem because the cut ends of the detector ring are elliptic. So far, we have considered two candidates for the detector arrangement method to realize the SROP. The first method is a slant- ellipsoid type of SROP (SE-SROP), in which the ellipse-shaped detector rings are stacked at a slanted angle (Fig. 4.19a) [33]. The second method is an axial-shift type SROP (AS-SROP), in which the block detectors originally forming a conventional cylindrical PET scanner are axially shifted little by little so that the center of each detector surface, positioned on the parallel planes, is slanted against axial direction (Fig. 4.19b) [34]. We conducted a numerical simulation study to compare these two methods for constructing the small SROP with respect to their imaging performance and the effect of utilizing DOI detectors. For this simulation, each block detector was composed of a four-layer 16 × 16 array of 3.0 × 3.0 × 7.5 mm3 Gd2SiO5 (GSO) scintillators. In the non-DOI cases, each block detector was composed of a 16 × 16 array of 3.0 × 3.0 × 30.0 mm3 GSO scintillators. The small SROP was designed with two detector rings of 16 block detectors. For both types, an open space with the width of 96.0 mm was obtained. A cylinder phantom including point-like sources was simulated. The diameter of the point-like sources was 4.0 mm, and the contrast ratio of the radioactivity in the point-like sources and the background cylinder was 10:1. Projection data were generated by forward projection using a geometrically defined system matrix based on the sub-LOR model [25]. Noise, attenuation, and scattering were not included in the simulation. Images were reconstructed by the 3D OS-EM method with 8 subsets and 10 iterations. The voxel size was 1.5 × 1.5 × 1.5 mm3, which was half of the scintillator size.

Fig. 4.19 Presents simulation geometries of the small SE-SROP (a) and AS-SROP (b), with illus-trations of the simulated cylinder phantom including the point-like sources


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Figure 4.20 shows the reconstructed images for the simulation of the small SE-SROP and small AS-SROP with and without DOI measurements. In the case of no DOI simulation, artifacts appeared in the coronal slice of the reconstructed image for the SE-SROP and in the transaxial slice of that for the AS-SROP. In the case of DOI simulation, no visible artifacts appeared in both types. The full width at half maximum (FWHM) was calculated by averaging for transaxial and axial FWHMs of all point sources to evaluate spatial resolution. As a result, we found the DOI detector improved the spatial resolution in both cases. Therefore, we concluded that both types had almost the same imaging performance of an acceptable quality using DOI detectors.

4.3.3 Development of a Small Prototype

Figure 4.21 shows the design illustrations of the small AS-SROP prototype [34] that had two modes (open and closed), which are shown in Fig. 4.22. The open mode formed the SROP geometry and the closed mode formed the cylindrical PET geom-etry. Sixteen detector units each of which consisted of two DOI detectors were arranged to form a perfect circle with a diameter of 250 mm. Detector units had an axial shift mechanism so that they could be transformed into the SROP having an open space of 139 mm. In the open mode, the center of each detector surface,

Fig. 4.20 Shows the reconstructed images in the simulation of the small SE-SROP (a) and AS-SROP (b) types. The top row is without DOI measurement and the bottom row is with the DOI measurement. For each pair of images, the left is the transaxial slice and the right is the coronal slice. The average FWHMs are shown


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positioned on the parallel planes, was slanted 45° against the axial direction. For shifting purposes, each detector unit was connected to the neighboring detector units by linear guides. Transformation between the cylindrical PET (Fig. 4.22a) and the SROP (Fig. 4.22b), which was controlled by one rotation handle, was completed within 10 s.

Each DOI detector consisted of a H8500 PMT and the four-layer 16 × 16 array of Zr-doped GSO (GSOZ) scintillators (Hitachi Chemical Co., Japan) with a size of 2.8 × 2.8 × 7.5 mm3. We should note that while the lutetium (Lu)-based scintillators are widely used for PET, Lu has intrinsic radioactivity. Therefore, we used Lu-free scintillators to reduce the background activity as low as possible for low activity measurements. The output signals of two DOI detectors stacked axially were pro-jected onto one 2D position histogram by an Anger calculation. Each DOI detector was in coincidence except for the neighboring DOI detectors. Coincidence data were corrected as list-mode data. The random correction was done by delayed coin-cidence data subtraction. The energy window was 400–600 keV and the timing window was 20 ns.

Fig. 4.21 Shows design illustrations of the small AS-SROP prototype: (a) depicts the front, (b) the side, and (c) the top views. Dark boxes indicate scintillator blocks


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The compact and all-in-one design of the prototype allowed us to use it for in- beam experiments in the HIMAC. The gantry stage had wheels, all the electronics, data acquisition (DAQ) system, and power supply stored in the gantry stage (Fig. 4.21a). Therefore, the entire system could be easily transported between PET areas and the irradiation room in the HIMAC.

We evaluated the basic performance of the prototype. The spatial resolution and sensitivity were 2.6 mm and 5.1% for the open mode and 2.1 mm and 7.3% for the closed mode. The AS-SROP enables in-beam PET imaging at a slight cost of imag-ing performance. The decrease of the performance can be minimized, for example, by transforming into the close mode immediately after the irradiation while main-taining the open space only for the in-beam PET measurement.

For an imaging demonstration, a small rod phantom, which had a similar struc-ture to the micro-Derenzo phantom, was measured. The rod phantom consisted of an outer hollow cylinder and an inner solid cylinder with a diameter of 36.1 mm and a length of 17.8 mm. The phantom was filled with 18F solution with total activity of 2.5 MBq, and it was measured at the center of the FOV with both modes. We measured two positioning angles of the phantom, standup and laydown, so that the inner rod holes were directed in the z direction and y direction, respectively. Measurement time was controlled so that the total number of decays became almost the same for each measurement. Because the number of coincidence events was large, the list- mode data were converted to the LOR histogram. Images were recon-structed by the 3D OS-EM algorithm with 8 subsets and 10 iterations. The random correction was applied by subtracting delayed coincidence but attenuation correc-tion and scatter correction were not applied. The voxel size was 1.5 × 1.5 × 1.5 mm3.

Fig. 4.22 Provides photographs of the small AS-SROP prototype taken from the side in the closed conventional PET (a) and OpenPET (b) modes. The transformation from the closed mode to the open mode was completed within 10 s


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Figure 4.23 shows reconstructed images of the small rod phantom for the open mode and the closed mode using the 3D OS-EM with 8 subsets and 10 iterations. Center slices perpendicular to the rod directions are shown. The measured numbers of counts were 189 M (standup) and 181 M (laydown) for the closed mode and 132 M (standup) and 129 M (laydown) for the open mode. Rods of 2.2 mm diameter were clearly separated for both modes and both positioning angles of the phantom.

4.3.4 In-Beam Imaging of Carbon Ion Beam

The small prototype was brought into the HIMAC irradiation room where we con-ducted the in-beam imaging tests. A PMMA phantom was irradiated with an RI beam of 11C, which has a half-life of about 20 min. The RI beam was generated as a secondary beam from the 12C irradiation onto a Be target. Figure 4.24 shows the experimental setup. The dimensions of the rectangular cuboid PMMA phantom were 40 × 40 × 100 mm3, and an additional PMMA board with a thickness of 9 mm was used to cover the upper half of the beam. The irradiation time was about 10 s and a dose of about 2.5 Gy was given. The beam intensity was about 5 × 106 parti-cles per second and the beam energy was about 340 MeV/u. Therefore, the Bragg peak position in PMMA was at about 168 mm. A PMMA range shifter of 115 mm thickness was used so that the Bragg peak position was moved to 53 mm from the entrance of the phantom. PET data were measured during the irradiation and for 20 min after the irradiation ended. Because the irradiation beam in the HIMAC had a cycle of 3.3 s with beam-on time (spill on) of about 1.8 s and a beam- off time (spill off) of about 1.5 s, PET data for the spill-off time were extracted. Because the num-ber of counts acquired for the in-beam measurement was small, the 3D list-mode maximum likelihood expectation maximization (LM-MLEM) was employed for the PET image reconstruction. The number of iterations was 50 and the voxel size was

Fig. 4.23 Reveals reconstructed images of the small rod phantom


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1.5 × 1.5 × 1.5 mm3. Random correction and attenuation correction were applied, but scatter correction was not applied because the phantom size was small enough so that the scatter effects can be ignored.

Figure 4.25 shows the in-beam measurement results for the 11C beam irradiation. The count rates for coincidence events were recorded. The phantom was irradiated with spills (3.3 s irradiation cycle) to give the total dose of about 2.5 Gy. The count rate graph shows that there were huge numbers of counts during the spill-on time, due to the prompt γ-rays. However, they were single γ-rays and could not be used for PET imaging. Even if there were counts from positron emitters, it was hard to distinguish them from background single γ-rays. On the other hand, coincidence events in the spill-off time, when the prompt γ-rays were not detected, could be used for imaging. Figure 4.25b, c shows images reconstructed from counts measured between spills and for 20 min after the irradiation. The number of counts was 626 k counts. Figure 4.25d shows the profiles of single voxel lines along the beam direc-tion for the upper part with the 9 mm PMMA board and for the lower part without it. The peak position of the lower part was calculated by parabola fitting as 55.0 mm (planned Bragg peak at 53 mm). The difference in peak positions was clearly observed and measured as 8.8 mm.

4.4 Development of a Full-Size DROP System

4.4.1 System Development

The full-size DROP prototype consists of two detector rings, and each detector ring has two subrings of 40 detectors (Fig. 4.26). Each detector consists of 16 × 16 × 4 array of GSOZ (2.8 × 2.8 × 7.5 mm3). The portable gantry has a compact design;

Fig. 4.24 Depicts the setup for PMMA phantom irradiation by the 11C beam and in-beam PET measurements. A photograph of the setup (a) and illustrations for front (b) and top (c) views are shown


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each detector ring has a 940 mm outer diameter and 171 mm axial length for the detector inner bore of 640 mm diameter and 113 mm axial length.

The key technology, which enabled OpenPET, is the four-layered DOI detector. In order to measure radiations from the limited activity produced though fragmenta-tion reactions, GSOZ was chosen for the scintillators instead of Lu-based scintillators because GSOZ contains less natural radioactivity, although timing performance was compromised. On the other hand and in order to compensate for the limited light yield, we used 64channel flat-panel PMTs with a super-bi-alkali photocathode (Hamamatsu R10552–100-M64), which had a 30% higher quantum efficiency [35]. Then, to enable stable in-beam PET measurement even under high background radiations, voltage divider circuits were designed so as to have five times higher linearity.

Lastly to avoid severe radiation damage to ASICs, gain control circuits were not implemented in the front-end circuits, while position analyzer circuits were placed with a 15-m cable extension. A data acquisition system was developed based on the single events collection.

Fig. 4.25 Shows the results of the in-beam experiment, which demonstrates the results for imag-ing the 9 mm difference given by the Bragg peak positions: (a) reveals count rates in the first 70 s of the in-beam measurement of the 11C beam, (b) is the center slice of the image reconstructed from counts measured between spills and for 20 min after the irradiation, and (c) is the 3D visual-ization of the image with the shape of the PMMA phantom, indicated by dotted lines and gray rectangular parallelepipeds. Profiles along the beam direction for upper and lower lines indicated in (b) are plotted in (d), in which the center of the FOV is x = 0


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4.4.2 Imaging Test

The system was tested with a carbon ion beam of clinical intensity at HIMAC as shown in Fig. 4.27. Phantom images were obtained by applying a GPGPU-based list-mode iterative reconstruction algorithm with geometrical detector response modeling. A PMMA phantom (10 × 10 × 30 cm3) was irradiated by a 12C ion beam (mono energy pencil beam). About 2.5 Gy was given during about 10 s irradiation and PET data were measured for 20 min immediately after the irradiation. Figure 4.28 shows a reconstructed image of the distribution of the positron emitters, which were produced through fragmentation reactions. The peak position in the PET image is highly correlated to the Bragg peak position.

Fig. 4.26 Developed full-size DROP


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4.5 Conclusion

This chapter describes basic concepts and component technologies of our OpenPET systems, which led to the world’s first open-type 3D PET scanner. Even with a full- ring geometry, the DROP has an open gap between its two detector rings. The SROP is a more efficient OpenPET geometry in terms of the gap size and the sensitivity. Demonstrations with prototypes showed a proof of concept of PET image-guided particle therapy such as in situ dose verification and direct tumor tracking.

Fig. 4.27 Depicts the in-beam measurement test of 12C beam irradiation at HIMAC

Fig. 4.28 Shows an example of reconstructed DROP images of a PMMA phantom (12C beam irradiation)


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Chapter 5High-Resolution PET/CT Development

Wai-Hoi Wong and Yuxuan Zhang

Abstract The main clinical application of positron emission tomography (PET) is in oncology, where it is used to diagnose malignant tumors, cancer staging, treat-ment response, follow- ups, and recurrence of diseases. The PET camera can detect therapeutic changes earlier than anatomical imaging modalities, because the struc-ture being studied must significantly change in size and shape before it is detectable by the latter devices. Such important features of PET in oncology are, however, reduced by the image resolution and quality (noise) of clinical PET/CT systems, thereby limiting its effectiveness to diagnose lesions under a centimeter in size. Improving the PET image resolution and quality would enhance the oncologic effi-cacy of PET/CT by detecting smaller lesions with more accuracy of tracer uptake. It would also lead to earlier cancer detection, more accurate cancer staging, and more sensitive monitoring of treatment responses.

5.1 Introduction

The main clinical application of positron emission tomography (PET) is in oncol-ogy, where it is used to diagnose malignant tumors, cancer staging, treatment response, follow-ups, and recurrence of diseases. The PET camera can detect thera-peutic changes earlier than anatomical imaging modalities, because the structure being studied must significantly change in size and shape before it is detectable by the latter devices. Such important features of PET in oncology are, however, reduced by the image resolution and quality (noise) of clinical PET/CT systems, thereby limiting its effectiveness to diagnose lesions under a centimeter in size. Improving the PET image resolution and quality would enhance the oncologic efficacy of PET/CT by detecting smaller lesions with more accuracy of tracer uptake. It would also lead to earlier cancer detection, more accurate cancer staging, and more sensitive monitoring of treatment responses.

W.-H. Wong (*) • Y. Zhang Department of Cancer Systems Imaging, The University of Texas MD, Anderson Cancer Center, 77030 Houston, TX, USAe-mail: [email protected]


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The intrinsic spatial resolution of clinical PET or PET/CT systems has been lim-ited to 4.5–6.5 mm in the last three decades. But recently, a 3 mm or better PET intrinsic spatial resolution has been achieved, with the advent of:

(a) L(Y)SO scintillation crystal that has higher light output and faster scintillation (b) Small silicon photomultipliers (SiPM) (c) Higher-resolution detector design, such as the PMT-quadrant-sharing (PQS)

design (d) Time-of-flight data acquisition and reconstruction (e) Point-spread-function (PSF) iterative image reconstruction (f) Larger axial field-of-view PET camera (g) Localization of the depth of interaction of the gamma ray in the PET detectors

The new generation of clinical PET/CT incorporates some, but not all, of above advancements, which provides substantive PET/CT resolution to improve the application of PET for personalized oncology with more accurate tracer uptake quantification and diagnosis of smaller lesions. The ultimate incorporation of all the available advancements in future PET/CT systems can drive the molecular imaging capability of PET/CT to significantly higher levels for a better grasp and manage-ment of cancer. The technological tools to achieve higher levels of performance are discussed below.

5.2 Scintillation Crystals for High-Resolution PET/CT

The detector materials and how each material is used (detector design) fundamen-tally determine the intrinsic imaging resolution of PET. PET detectors require high effective atomic number (Zeff) and high density (ρ) to effectively capture the ener-getic 511 KeV gamma rays for PET imaging. While the ability to stop the 511 KeV gamma rays in the detector is necessary for high-resolution PET imaging with nar-row detector pixels, it is an insufficient condition should a gamma ray be absorbed in the detectors by multiple Compton scatterings. In such scenario, the energy signal would be distributed randomly over multiple neighboring crystals, thereby misre-porting the first-entry location of the gamma ray. Thus for high-resolution PET imaging, the first detector interaction must have a high probability of photoelectric interactions (photofraction), which absorbs all the gamma energy at once. The pho-toelectric interaction probability is proportional to (Zeff)4. The characteristics of some usable detector materials for the 511 KeV gamma detection [1–3] are listed in Table 5.1.

High Zeff and photofraction (P) also provide high coincidence-detection sensitiv-ity, as a function of P2 because only photopeak events are accepted for the annihila-tion gamma rays to minimize the high scattered event noise in clinical PET images. Furthermore, high Zeff and high stopping power materials have the lowest-resolution degradation caused by the uncertain depth of interaction (DOI), as thick detectors are needed to stop the 511 KeV gamma rays. From the above fundamental considerations

W.-H. Wong and Y. Zhang

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of photofraction, sensitivity and DOI resolution degradation, for high resolution and high sensitivity, BGO is a good detector material, especially considering its low cost. Among the currently available detector materials, BGO can conceptually yield the highest resolution PET/CT if the scintillation light from each BGO pixel is detected by its own dedicated photosensor (1:1 coupling).

However, the current photosensor technology is limited to either the large size (18–50 mm) photomultiplier tubes (PMT) or the small (1–6 mm) but relatively costly silicon photomultipliers (SiPM). Notably, both hinder a 1:1 coupling between crystal pixels and photosensors. Coupling many crystals to a single photosensor would need the Anger positioning method for which the imaging resolution also depends on the light output of the scintillation crystal. The L(Y)SO crystal has four times the light output of the BGO and would provide higher spatial resolution when the Anger positioning is used for the detector design, despite its worse DOI blurring effect outside the field-of-view (FOV) center. In addition, the fast timing and high light output of LYSO can provide much more accurate timing of arrival for the gamma ray to 300–600 ps (Δx = 4.5–9 cm), which provides enough time-of-flight (TOF) information to improve the image signal-to-noise ratio by more than two times, especially for large patients. Hence, L(Y)SO is a good choice for high- resolution clinical PET/CT systems, despite its lower Zeff and four times higher cost than BGO. However, for small FOV PET, such as brain, breast, and small-patient systems, where TOF is less beneficial and ultrahigh resolution and low cost are highly desirable, BGO crystals with 1:1 photosensor SiPM coupling would provide the highest-resolution imaging.

The fast scintillating LaBr crystal potentially has two times better TOF resolu-tion than L(Y)SO and much better energy resolution for reducing scattered event noise [4, 5]. However, these two potential gains may not make up for its low coinci-dence photofraction-detection sensitivity that is 1/6 of BGO and 1/3 of that of L(Y)SO, for the same detector system geometry and axial FOV. CZT has the same coin-cidence photofraction-detection sensitivity disadvantage as LaBr and compounded by its poor timing that is inadequate to provide TOF information to enhance the image quality. Certainly, the LaBr and CZT system detection sensitivities can be increased by having a much larger axial FOV than BGO/LSO systems, but would increase the production cost. Both LaBr and CZT are expensive materials to grow, which limits the axial FOV of PET systems. The low 18% photofraction of the LaBr and CZT (low Zeff) means that 82% of the gamma rays detected would be scattered

Table 5.1 Characteristics of PET detector materials for 511 KeV gamma rays

BGO L(Y)SO NaI(Tl) LaBr CZT

Zeff 75 66 51 50 50Photofraction (P) of 1st interaction 43% 32% 19% 18% 18%Light output (relative to NaI) 18 75 100 175 NAEnergy resolution 11% 12% 8% 3% 2%Decay time (ns) 300 40 230 20 NADepth of Interaction blurring Lowest Moderate High High High

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multiple times in the detectors, which would degrade the imaging resolution. In addition, the low stopping power also worsens the DOI image blurring effect. Hence, the LaBr and CZT are not ideal detector materials to make ultrahigh- resolution PET systems.

5.3 Solid-State Silicon Photomultiplier (SiPM) Photosensors

Traditionally, the scintillation light produced by the crystals is detected by photo-multiplier tubes (PMT) and converted into electronic signals for the front-end elec-tronics to process. The arrival of the silicon photomultipliers (SiPM) has provided a potentially exciting alternative to PMT in the design of high-resolution PET/CT systems because SiPM are small (1–3 mm square), thin, lightweight, resistant to high magnetic field, and much lower operating voltage. The disadvantages are higher noise, more temperature sensitive, and more costly to cover the same crystal footprint, at least for the time being. It needs 81–144 SiPM (3–4 mm in size) to replace a single 38-mm PMT. Due to lower gain and higher noise, each SiPM chan-nel needs its own amplifier and supporting electronics. Replacing a PMT with a 100 SiPM increases the processing electronics by a 100-fold, which further increases the production cost and substantially increases the heat generation that requires much better cooling.

One advantage of SiPM is the small size that enables an ultrahigh-resolution PET system of 1–3 mm to be achieved without using analogy decoding schemes (Anger decoding) used by PMT systems. On the other hand, the best analog- decoding scheme using PMT can also achieve 1–3 mm resolution in animal and clinical systems [6, 7]. The second advantage of SiPM is its immunity to strong magnetic fields, thus making the PET/MR hybrid system a reality [8–10]. However, for PET/CT systems, PMT remains a viable low-cost choice while able to achieve very high resolution, approaching the fundamental PET resolution limit with the latest detector design presented in the next section.

5.4 The Fundamental PET Resolution Limits

Assuming the ideal detector system, the fundamental limit of PET image resolution is imposed by (a) the non-colinearity (0.5–0.54°) of the pair of 511 KeV gammas emitted, which is about 2 mm for the large clinical systems with 85 cm detector ring diameter and (b) the positron range of the tracer isotopes in water and the lung, as shown in Table 5.2 [11–12]. The non-colinearity effect originates from the Doppler effect of center-of-mass energy (Ec) of the short-lived positronium atom before the positron-electron annihilations; the FWHM of Ec is about 2.4 KeV [13]. The posi-tron range is inversely proportional to the density of the surrounding tissues, which for the normal lung tissue is four times larger than soft tissues.


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The fundamental intrinsic PET resolution limits can be derived from Eq. 5.1. The fundamental resolution limit (FWHM) does not include the detector resolution:

Fundamental resolution limit FWHM detectormean( ) ( ( .= +R 2 0 0022 ring diameter) )‰2

(5.1)

The fundamental intrinsic PET resolution limits in water-equivalent tissue and the lung is shown in Table 5.3 for different tracer isotopes.

For a large whole-body PET with F-18 tracers, such as FDG, FDOPA, F-Choline, FMISO, FLT, FAZA, and fluoride, the fundamental PET resolution is 2 mm in water-equivalent tissues and 3 mm in the lung. With the generator produced isotopes 68Ga, 62Cu, and 82Rb, the fundamental PET resolution limits are in the range of 3.5–6 mm in water-equivalent tissues and 11–23 mm in the lung.

The fundamental PET resolution for a dedicated brain PET with a smaller 40-cm detection ring, assuming the ideal detector system, is shown in Table 5.4.

With F-18, the fundamental PET resolution is about 1 mm for a small dedicated brain PET system. With the generator produced Ga-68, the resolution limit is 3.0 mm. Hence, the small brain PET geometry can potentially improve PET resolu-tion from 2 mm to 1 mm for F-18 tracers, but would not improve the resolution limit for the generator produced isotopes.

Current clinical PET/CT systems are whole-body systems with a detector ring diameter of 80–87 cm and use analog-decoding block detector designs. The typical image resolutions of these typical clinical PET systems are in the range of 4.5–6 mm, which is substantially worse than the fundamental PET resolution limits (2–3 mm), shown above for the cyclotron-generated isotopes, especially for the F-18 tracers.

Table 5.2 Mean energy, maximum energy, and mean range of positron emitted

Isotopes Emean (KeV) Emax (KeV) Rmean (mm) water Rmean (mm) lung18F 252 635 0.6 2.411C 390 970 1.1 4.213N 488 1190 1.5 5.815O 730 1720 2.5 9.668Ga 844 1899 2.9 1162Cu 1280 2926 6.1 2382Rb 1551 3378 5.9 23

Table 5.3 Fundamental resolution limits in water and lung for a whole-body PET18F 11C 13N 15O 68Ga 62Cu 82Rb

PET resolution limit in water (mm) 2.1 2.3 2.5 3.2 3.5 6.4 6.2PET resolution limit in lung (mm) 3.1 4.7 6.1 9.8 11.2 23.1 23.1

Table 5.4 Fundamental PET resolution limits of a dedicated brain PET (40-cm)18F 11C 13N 15O 68Ga 62Cu 82Rb

PET resolution limit in water (mm) 1.1 1.4 1.8 2.7 3.0 6.2 6.0


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Thus substantially higher-resolution PET imaging, with the resolution size reduced by two times or more, can be realized if (a) cost-effective higher-resolution PET detector designs are achievable and (b) the count collected is proportionally increased by 23 times or more. The increase in collected counts is necessary because, unlike CT, PET imaging is a count-deficient imaging modality with the image quality fundamentally determined by the statistical noise (photon statistical varia-tions) in the signals obtained by the detector system. To maintain image quality in a higher-resolution image, the number of counts collected in each voxel must be maintained. With the number of resolution pixels increased by two times in all three dimensions or by an eight-time increment in the number of voxels, the counts recorded by the detector system should be increasing by eight times. This can be accomplished by increasing the axial field of view (AFOV) by about 81/2 times or more, assuming the accidental and scatter fractions are constant. Hence, to effec-tively achieve a two times improvement in PET resolution to approach the funda-mental limit of PET imaging without increasing image noise, it is necessary to have higher-resolution detector designs, with half the detector crystal size and a 2.8 times increase in AFOV, from the current 20 to 56 cm.

5.5 Higher Spatial Resolution Detector Designs

(a) PMT-based PET detector design. For high-resolution PET/CT, PMT is a cost- effective photodetector to assess the light from the scintillation crystals because of its large surface area in comparison to the SiPM. PMT is still being utilized in the current clinical whole-body PET/CT systems. As shown in Fig. 5.1, the common detector design for clinical PET systems is the block detector design, which consists of an array or a block of small scintillation crystal elements coupling to four PMT units. When a detector block senses a gamma ray, the location of the scintillating crystal element where the interaction occurs is determined, by the Anger analog-decoding scheme [14]. This block detector design can decode detector crystal size of 6 × 6 mm2 with the low light output

Fig. 5.1 Traditional block detector design in clinical PET/CT systems


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BGO crystals and 4 × 4 mm2 with L(Y)SO crystals which has four times more scintillation light output than BGO. These crystal sizes give rise to the 4–6 mm image resolution in current clinical PET and PET/CT systems. As discussed in the last section, the fundamental PET resolution limit for soft tissue is 2–3 mm for the large whole-body PET and 1–2 mm for a small brain PET system. However, it is possible to further improve the clinical PET imaging resolution from the current 4 to 6 mm.

Figure 5.2 shows that a new ultrahigh-resolution whole-body PET/CT has been developed, using the photomultiplier-quadrant-sharing (PQS) detector design [15–17]. This PQS detector design successfully decoded detector crystal blocks with 16 × 16 L(Y)SO crystals of 2.35 × 2.35 mm2 (15.2 mm deep), using a very large PMT with 39 mm diameter [7]. This PQS detector-based whole-body PET/CT has achieved an image resolution of 2.8 mm, using the 3D filtered back-projection reconstruction and 1.4 mm resolution with the OSEM-PSF reconstruction [7]. The time-of-flight (TOF) resolution was 473 ps [18]. With a small detector diameter ring to reduce the non-colinearity effect, image resolution of 1 mm has also been achieved with the PQS detector design [17, 19], which would be useful for building dedicated brain PET systems. Hence, the PQS detector-based PET/CT systems can provide PET image resolution approaching the fundamental limit of PET imaging. The clinical images of the ultrahigh-resolution PET/CT system will be presented later in this chapter.

(b) Silicon photomultiplier (SiPM)-based PET detector designs. Unlike the 20–40 mm diameter PMT, the small SiPM photosensors allow each scintillation crystal element as small as 1 mm to be coupled to an individual SiPM, without using analog-position decoding as in the PMT-based PET detector designs. As a result, ultrahigh-resolution PET detectors approaching the fundamental PET resolution can be conceptually achieved with SiPM as photosensors. As discussed earlier, one of the disadvantages of SiPM is the higher cost of the numerous SiPMs and their individual processing electronics to cover the same

Fig. 5.2 Upper: traditional PET block detector design using four PMT to decode the crystal location. Lower: the PMT-quadrant-sharing detector design using four one-quadrant of four PMT to decode the crystal location while reducing the crystal and block size by half to improve imaging spatial resolution by two times in each dimension of the imaging space


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crystal footprint of one PMT (Fig. 5.3). Namely, 81–144 SiPM (3–4 mm in size) are needed to replace a single 38-mm PMT used in clinical PET systems. Due to lower gain and higher noise, each SiPM channel requires its own amplifier and supporting electronics. Replacing a PMT with a 100 SiPM increases the processing electronics by a 100- fold, which further increases the production cost and substantially enhances the heat generation, thus requiring much better cooling. For building affordable and practical whole-body PET systems, the number of SiPM used needs to be reduced, which leads to the need of some form of analog decoding. Currently, SiPM-based PET/CT technology is still evolving to make it more affordable and practical, cost competitive, and perfor-mance competitive with PMT-based PET/CT systems. For PET/MR systems, SiPM with its magnetic field immunity is the only option, which ultimately contributes to the high cost of PET/MR systems.

In comparison to the PMT-based clinical PET/CT systems, with an image resolu-tion of 2.8–6 mm [7, 20–22], commercial clinical PET/MR systems have PET reso-lution in the range of 3–4 mm [8–10]. The resolution of the PET in PET/MR systems using small SiPM is still short of the fundamental PET resolution limit of (1–2 mm). With future development of more cost-effective SiPM-based PET detector designs, ultrahigh-resolution PET/MR, and PET/CT approaching the fundamental limit can be achieved.

5.6 Improving PET Resolution with Depth-of-Interaction Positioning

Even with the high stopping power, scintillation crystals such as BGO and L(Y)SO, PET detectors need to be deep or thick (15–30 mm) to efficiently detect the ener-getic 511 KeV gamma rays. The deep detectors coupled with the coincidence detec-tion of PET and the PET’s detector ring geometry radially broaden the point spread function (resolution) (Fig. 5.4). The radial resolution worsens with deeper detectors

Fig. 5.3 One PMT and 16 SiPM units in a 4 × 4 array


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and a smaller detection ring, despite the fact that PET detection sensitivity can be improved with deeper detectors and smaller detection rings (Fig. 5.5). As depicted in Fig. 5.6, such depth-of-interaction (DOI) resolution degradation also affects the axial resolution with a large axial field of view (AFOV), despite the fact that a large AFOV significantly increases PET’s true count sensitivity to (AFOV)2. If the DOI of a detected gamma ray can be measured, both the transaxial and the axial resolu-tion of PET can be improved. Furthermore, in ultrahigh-resolution PET systems that have very narrow detectors of 1–2 mm2, the DOI resolution degradation effect of deep crystals of 15–30 mm becomes more significant. Hence, it is essential to measure the DOI in ultrahigh-resolution and high sensitivity PET systems.

The classic DOI PET detector design is the dual-end DOI that has one solid-state photosensor such as SiPM, coupled to each end of the deep scintillation detector

Fig. 5.4 The coincidence line-of-response point spread function widens with deep detectors in a ring geometry

Fig. 5.5 Monte Carlo simulation of a small 65-cm PET detector ring. The 4-mm detector (20 mm deep) has resolution that degrades from 3.5 mm at center to 7 mm at position of 200 mm. The 2.4-mm detector (20 mm deep) has its resolution that degrades from 2 to 6.5 mm


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pixel [23, 24], and the depth of interaction is derived from the difference in scintil-lation signals received by the two photosensors. This design can achieve DOI reso-lution of 2 mm, but it is costly to be implemented into a large clinical PET system, as it doubles the already large number of photosensors and supporting electronics.

Notably, there is no clinical whole-body PET system with DOI sensing at the present time. However, DOI sensing is important for providing the highest possible resolution in PET imaging. There are ongoing developments to engineer more prac-tical DOI PET detectors, using fewer SiPMs by means of decoding schemes [25, 26] and a monolithic crystal block design [27]. Another direction of resolving the DOI degradation is using iterative image reconstruction that incorporates the system DOI degradation model and the detection point spread function to deconvolute the DOI effect in the estimated image [28, 29], which is also known as the high- definition (HD) image reconstruction. However, the HD iterative reconstructions may produce image edge artifacts [30].

Because of the importance of DOI sensing to improve PET resolution and facili-tate large axial FOV PET that significantly improve PET sensitivity, DOI sensing PET will be a reality in the future.

5.7 Time-of-Flight Data Acquisition

By measuring the difference in the time of arrival of the two coincidence gamma rays, the location of the positron annihilation site can be estimated, which would improve the image quality (noise), which is equivalent to an increase in the effective detection sensitivity of PET. The timing uncertainty of this measurement is called the time-of-flight (TOF) resolution. In current clinical PET/CT that uses L(Y)SO crystals, the TOF resolution is about 390–525 ps in clinical PET/CT and PET/MR systems [8–10], depending on the speed of the photosensors, the efficiency of capturing scintillation light in the detector, and the accuracy of the electronic time- digital converter. The TOF positron localization uncertainty is 7.5 cm with a TOF resolution of 500 ps, which has been estimated to improve the effective PET detection sensitivity by 1.5–3 times [7, 21, 22]. Such sensitivity improvement depends on the patient size, the body cross section, and the metric of measuring this

Fig. 5.6 The coincidence PSF worsens with large AFOV that can increase sensitivity significantly


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effective gain. Using LaBr scintillation crystals, the TOF resolution can be improved to below 400 ns translating to a positron localization uncertainty < 6 cm [31]. However due to the low photofraction of LaBr, the three times reduction in coincidence- detection sensitivity relative to L(Y)SO crystals (see Sect. 5.2), the loss in detection sensitivity of LaBr may exceed the higher TOF resolution gain relative to L(Y)SO, as a result of the low atomic number (Z). Relative to the low-cost BGO, LaBr suffers a 5.7 times reduction in first interaction photoelectric coincidence-detection sensitivity.

The advantage of TOF PET systems has been well demonstrated and quantified in phantom studies [32, 33]. For general clinical use, the TOF advantage is real but has more variability [34]; Fig. 5.7 shows that the average TOF gain in signal-to- noise ratio (SNR) was 1.1–1.3 times for the head, which corresponds to 1.2–1.7 times in sensitivity (Fig. 5.7). For the lung, the average SNR gain was 1.4 times, corresponding to a sensitivity gain of two times. The largest TOF gain was found in the abdomen, with SNR gain of 1.2–1.6 times depending on the body mass index (BMI), which is equivalent to 1.4–2.6 times increase in detection sensitivity. For the

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Fig. 5.7 Measured SNR gain factor vs. BMI, for small lesions (<2 cm) in different parts of the body for a group of 100 patients: (a) 144 lesions located in the abdomen, (b) 67 lesions located in the lung, and (c) 30 lesions located in the head and neck (This research was originally published in JNM by C. Lois et al. [34])


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predominant clinical PET of whole-body cancer staging application, the overall average effective TOF sensitivity gain that equals (SNR gain)2 is approximately two times, which corresponds to a reduction in scan time by 50%.

5.8 Large Axial Field-of-View PET Camera

Increasing the AFOV is an effective means to significantly improve the count starva-tion characteristics of PET imaging because PET detection sensitivity is proportional to (AFOV)2. The same effective sensitivity gain of two times from the current TOF PET/CT may be accomplished by increasing the axial field of view (AFOV) by 1.4 times. The current mainstream clinical PET/CT systems have AFOV of 20 cm. If this AFOV is increased from the current 20–25 cm to 1–1.2 m, it would cover the head and torso of more than 95% of the US male population [35], and detection sensitivity would be increased by 25 times from current commercial clinical PET/CT systems [36, 37]. The order of magnitude increase in sensitivity can be deployed in the fol-lowing ways: (a) it can reduce the whole-body imaging time to 1–2 minutes with the current whole-body cancer staging clinical protocol of imaging one time at 30–45 min postinjection of FDG. This would significantly increase the clinical patient through-put to lower PET/CT imaging cost. Furthermore, a short one-minute imaging would significantly reduce patient movement comparing to a 30 min scan time, thereby substantially reducing artifacts from patient movements. (b) It would significantly reduce the image noise to provide much more accurate quantitation of molecular activities and significantly lower the molecular concentration levels that can be detected, which is important for imaging receptors and small lesions; and (c) current 20-cm AFOV PET/CT systems prevent dynamic imaging of the whole body, as the system has to step through 5–7 body sections at different times to image the whole body. The long AFOV PET/CT will enable whole-body dynamic tracer uptake imag-ing, assessing the entire body at exactly the same time. Studies have shown that a long AFOV PET will increase the effective detection sensitivity (EDS)* by 20–40 times [36, 37] depending on the AFOV size and crystal type, thus enabling the cap-ture of changing time activity of tracers for the entire body.

Dynamic imaging is the only way to obtain quantitative physiology parametric (QPP) images for the entire body [38–44]. QPP images are more consistent among different clinical sites, which may reduce the barriers of multicenter studies and cross-center references. Hence, an affordable long AFOV PET may unleash the full potential of PET to study human physiology and quantitative molecular functions. It is particularly useful for studying the systemic whole-body therapy effects of drugs and their whole-body toxicity for normal tissues. Other advantages of a long AFOV PET that cover the whole torso and head are:

(1) The whole-body PET cancer staging imaging takes 20 min (3 min for 6–7 bed positions) with a 10 mCi of FDG injection, in which time and dosage can be respectively reduced to 1–2 min and 5 mCi. Since accidental coincidence noise


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is proportional to (radioactivity)2, reducing dose by 1/2 would reduce accidental noise by four times [45]. Hence, the system would improve image quality, lower the patient exposure to radiation by half, and shorten imaging time by ten times.

(2) High patient throughput: Assuming that the 1-m PET takes 2-min imaging plus 8 min for patient preparation (10 min total), compared to the current protocol of 20 min + 8 min + bed movement time for seven positions (30 min), patient throughput can be increased by more than three times.

(3) With 10 min/patient throughput and a 5 mCi dose per patient, the current “unit dose” of 10 mCi used for one patient can be used for two patients, since there is little radio decay for FDG after 10 min. Coupling this two times reduction in FDG cost to the two times higher patient throughput, the cost of cancer staging by PET would be substantially reduced.

(4) A 20 times increase in detection sensitivity will proportionally increase the counts per image pixel and facilitate the imaging of low-level receptor studies and other reaction studies, which are not currently achievable. For FDG dynamic studies, a 20 times increase in counts/pixel would produce whole-body meta-bolic rate images. Such images will have much higher accuracy when the counts in each pixel can be processed by the Patlak method or the 3-compartment model curve fitting process to yield whole-body images of the metabolism rate constants [41, 43, 44], without relying on the tedious and less revealing ROI drawings in each region to get enough count statistics.

5.9 Current State-Of-The-Art High-Resolution Clinical PET/CT Imaging

The recently developed high-resolution clinical PET/CT can provide image resolu-tion of 2.9 mm with 3D filtered back-projection reconstruction (NEMA standard) and 1.4 mm when PSF iterative reconstruction is used [7]. The high-resolution brain images from 8-min to 20-min patient studies are shown in Figs. 5.8 and 5.9, respec-tively. Figure 5.8 shows brain images in transaxial, coronal, and sagittal directions, whereas Fig. 5.9 depicts nine contiguous slices of the transaxial image set with a fine axial sampling of 1.2 mm and 2.4 mm slice thickness from the high-resolution PQS PET detector system, which reveals small structural changes from slice to slice. Figure 5.10 compares the high-resolution PET image with the MR T1-weighted brain images of the same patient. The high-resolution PET images from this clinical PET/CT closely resemble that of the T1-weighted MR images from a 1.5 T MRI system, demonstrating that the recently developed high-resolution clinical PET system is capable of imaging the molecular features of small anatomic structures in the body.

In conclusion, while clinical PET technology has advanced a long way in the last 30 years, the full potential of PET imaging remains to be completely grasped. Current PET detector technology can achieve PET resolution approaching 1 mm or


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Fig. 5.8 Transverse, coronal, and sagittal views of a human brain imaged for 8 minutes with an ultrahigh-resolution PET/CT system with 2.8 mm resolution

Fig. 5.9 Nine contiguous transverse images of a human brain with 20-min imaging. The slice- sampling spacing was 1.2 mm. Subtle changes in the brain structures between slices can be observed with the high-resolution axial sampling


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even smaller, using either the traditional PMT or the new SiPM photosensors. The bottleneck hindering the realization of the full potential of high-resolution PET imaging is the deficiency in detection sensitivities that the current time-of-flight tech-nology alone cannot adequately overcome. This would require the axial field of view of PET to be economically extended multiple times to cover the preferable whole torso or more.

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http://msis.jsc.nasa.gov/sections/section03.htm

Part IIISocial Aspect


Chapter 6Ethics, Regulations, and Clinical Development of Precision Medicine: Activating with Molecular Imaging

Chieko Kurihara and Tomio Inoue

Abstract “Precision medicine” is becoming a keyword toward new and more effective healthcare in the twenty-first century, a concept evolved from “personal-ized medicine.” Therefore, it is a prerequisite for the community of molecular imag-ing to clarify elements of ethics, regulations, and clinical development strategies to achieve the goal of precision medicine, activated with imaging technologies. Through literature review and continuous discussion with people of related com-munities, we identified key elements from view of regulations and clinical develop-ment strategies as follows: (1) quality assurance and standardization of methodologies and procedures of imaging technologies and (2) formulation of larger-scale global clinical trial network and imaging archives, both of which would accelerate regula-tory approval of new therapeutic drugs and diagnostic technologies. Additionally, key elements of ethics are as follows: (1) view of individual ethics to protect human rights and human dignity, i.e., (i) privacy protection, (ii) right to know and right not to know, as well as (iii) presymptomatic diagnosis consultation, and (2) view of col-lective ethics to assure social value such as (i) clinical trial registration and data sharing, (ii) justifiable commercialization, and (iii) preventing exploitation and stig-matization. In conclusion, precision medicine can be activated with molecular imaging, through more global collaborative initiatives, which recognize and have a profound understanding of the characteristics of science, regulations, and ethics in the era of precision medicine.

C. Kurihara (*) National Institute for Quantum and Radiological Science and Technology, Chiba, Japane-mail: [email protected]

T. Inoue School of Medicine, Yokohama City University, Yokohama, Japan


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6.1 Introduction

6.1.1 Precision Medicine Initiative

Voices for “precision medicine” have been skyrocketed since the plan to launch the “Precision Medicine Initiative” was announced by the United States (US) President Barack Obama on January 2015 [1]. The definition of precision medicine by the White House is described as “an innovative approach that takes into account individual differ-ences in people’s genes, environments, and lifestyles.” They expect that this strategy will overcome traditional “one-size-fits-all” approach, which has been designed for the “average patient.” They unveiled “$215 million investment in the President’s 2016 Budget” for research [2], which allocated $130 million to National Institutes of Health (NIH) to develop a cohort of a million or more volunteers, who agree to conditional sharing of their data with responsible researchers; $70 million to the National Cancer Institute (NCI) to accelerate genome-based cancer research; $10 million to Food and Drug Administration (FDA) to acquire additional expertise, including a new approach to evaluate next-generation sequencing technologies of DNA or even entire genomes; and $5 million to Office of the National Coordinator for Health Information Technology (ONC) to develop standards and requirements for privacy protection and security of data exchange across systems. They state that the mission is for the “development of individualized treatments” along with their “Guiding Principles for Protecting Privacy and Building Trust” [1]. These perspectives suggest their intention to strengthen ethi-cal, regulatory foundation of research and development in the era of genome-wide association study (GWAS) and “big data” analysis.

The term “precision medicine” has emerged during 2008–2011 to give a more “complete picture” to “personalized medicine” [3]. The World Molecular Imaging Society welcomed the term, which spotlights their technology with the announce-ment on October 2015 of “Precision Medicine… Visualized” [4]. In this context, our discussion concerning “precision medicine” aims to clarify key elements of ethics and regulations associated with the clinical development of personalized medicine, which can be activated utilizing technologies of molecular imaging.

6.2 From Personalized Toward Precision Medicine: A Change of the Trend?

Strategic reformation of drug development beyond “one-size-fits-all” had already been called for just around the year of 2000, when the US President Bill Clinton announced the completion of the first draft of human genome that resulted from the Human Genome Project, launched in 1990. The genesis of the idea for this Project dates back to the 1984 Alta meeting, which sought new DNA analytical methods to detect mutations among Hiroshima and Nagasaki survivors of the atomic bombing [5]. The term “personalized medicine” could only be insinuated in a simple mean-ing of patient-centered, good medical care [6]. Then and particularly in the context

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of genome science, this term came to be used to facilitate genome-targeted drug development. This covers drug seed hunting and prevention of genetic polymorphism- related adverse reactions. Several numbers of such drugs or diagnostics have been approved by regulatory agencies, and the approach to develop this kind of drugs has been called “pharmacogenomics” or “pharmacogenetics.” People imagined a medi-cal system in the near future, which aims to get whole-genome screening tests before or after the time of birth. Then, by means of an individual identifier code, people could access the best-fit medical care anytime and anywhere utilizing security- protected web system. These medical records, including individual’s genome information, are followed and utilized for large-size cohort analysis with-out identification of personal information. Ethical, social, and legal issues are to be well discussed, and rigorous human right protection measures need to be estab-lished. Such health system has been imagined and often described with the term “personalized medicine.”

So what is changing in “precision medicine?” Some argue that the term “person-alized medicine” creates a misleading image of a medical practice that is tailored to each individual, while others argue that there is no difference between them [3]. A major technological breakthrough toward “a new era” is the development of the so- called next-generation sequencing which enables to decipher the whole genome of one individual in a short period of time, at a reasonable cost, and thus enhances the affordability of the “whole genome-wide association study (GWAS).” At the same time, data-sharing strategy such as the US President “big data” initiative [7] led NIH to release 1000 genome data into the Amazon Cloud [8]. This kind of increased availability of genome data caused legitimate concerns of whether such data can be actually anonymous by means of traditional de-identification procedures. This question was posed by both, the US Department of Health and Human Services [9] and the Presidential Commission for the Study of Bioethical Issue [10]. The latter stressed the value of public benefit and responsibility of investigator as well as dem-ocratic deliberation. We should explore fundamental conflicts between “individual ethics” which is to ensure individual human rights and well-being and “collective ethics” which is to achieve public benefits. This conflict between individual and col-lective ethics was formerly articulated by the established clinical trial statistician Pocock [11]. Precision medicine must be achieved by means of seeking higher stan-dards of both collective and individual ethics. In keeping with this notion, International Cancer Genome Consortium (ICGC) promulgated standardization guidelines for the whole-genome sequencing analysis after benchmarking exercise among their study sites [12]. Here, we will be able to characterize “precision medi-cine” as follows [13]: Completion of human genome project and development of genome sequence technology have enabled or will enable large-size cohort studies and therapeutic options, which more precisely fit each subgroup of patients, beyond one-size-fits-all but not individually personalized. It will also enable more effective disease prevention strategies, through strengthened ethical foundation and scientific integrity, as well as regulatory reformation.

Then, is it necessary to define the role of molecular imaging in the era of preci-sion medicine? If so, this article is an overview of the regulatory reformation in the

6 Ethics, Regulations, and Clinical Development of Precision Medicine: Activating…

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USA from the beginning of the enterprise to establish regulatory frameworks for positron emission tomography (PET) imaging, as a key technology of molecular imaging. We called this process “PET Drug American Dream World History” [14–16] because it suggests a “gold standard” for regulatory framework of PET imaging and it has been more or less followed or referred to by other countries in the world [17]. Secondly, in the later part of this article, we discuss ethical issues to be deliber-ated more in depth in the era of “precision medicine,” from both perspectives of individual and collective ethics.

6.3 Regulatory Reformation Toward Personalized Medicine

6.3.1 Critical Path to Achieve Product Approval

6.3.1.1 Radioactive Drug Research Committee (RDRC) and Investigational New Drug (IND) Application/New Drug Approval (NDA)

It was in 1975 that the US government authorized the usefulness of “basic” research to administrate radiopharmaceutical drugs of limited dose to subjects with the aim to explore the human pathophysiology and drug metabolism. The code of federal regulations of this year established the RDRC [18], which states that basic research without the intention of diagnostic, therapeutic safety or efficacy evaluation of drugs can be exempted from IND regulations. IND requires investigators to apply for FDA authorization to initiate clinical trials. Alternatively, the investigator submits the protocol to an FDA-approved RDRC, along with ordinary process to submit to an institutional review board (IRB). While more than 70 approved RDRCs have to submit their annual reports to FDA, the agency does not review each protocol for authorization to initiate clinical trials [19]. This exemption from IND regulations is limited to the cases in which administered radioactive doses are limited in the range described below, which is known from previous experience in the literature. This means that first-in-human study without such previous experience is excluded from this RDRC framework to be followed, along with other limitations (e.g., limited to capable adults, number of study subjects, etc.).

– In whole body, active blood-forming organs, eye lens, and gonads: 3 rem (30 mSv) for single dose and 5 rem (50 mSv) for a cumulative annual dose

– In other organs: 5 rem (50 mSv) for single dose and 15 rem (150 mSv) for a cumulative annual dose

As for PET drug manufacturing, FDA Modernization Act (FDAMA) in 1997 [20], which covered whole regulations under the agency’s jurisdiction, caused great change of regulations for clinical development and clinical practice of PET drugs. According to the FDAMA and related regulations, after June 2012 (half a year extended deadline responding to the voices of related community), any of


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commercial, hospital, or academic institute which manufactures PET drugs for clin-ical use (excluding research use under IND or RDRC regulations) has come to manufacture PET drugs in compliance with PET drug-specific GMP (good manu-facturing practice/“PET-GMP,” hereafter) [21]. This means that they must obtain NDA, which means marketing authorization or abbreviated NDA from FDA, pass-ing FDA’s inspection. There had been monographs of 12 well-known generic PET drugs described in the US Pharmacopeia (USP) [22], among which 8 had not been approved but exceptionally included. These 12 PET drugs were exempted from the newly established PET-GMP regulations, as far as they were manufactured accord-ing to USP. However, as originally defined in FDAMA as well as in voices of PET drug specialists [23], this exemption expired, and these eight PET drugs were removed from USP at the end of 2014. Now, an increasing number of research/academic institutions or hospitals have obtained approvals of generic PET drugs such as FDG [14, 24]. In the same context, companies got approval of innovative PET drugs including amyloid imaging agents. Many of other promising PET drugs have been studied in pipelines of private companies and academic institutions.

6.3.1.2 Critical Path Initiative

Along with the above regulatory establishment, FDA’s Critical Path Initiative report in 2004 [25] highlighted the importance of biomarker assessment, one of which is PET imaging, among other various drug development tools. Demand of the citizens for more “personalized,” safe, and effective drugs has caused inflation of cost and stagnation of success in new and innovative drug development. FDA issued in 2005 a draft guidance entitled “Drug-Diagnostic Co-Development Concept Paper” [26] to introduce key critical steps during drug development that translated basic research into clinical applications, as a means to get regulatory approval, through several steps of validation process from biomarkers to diagnostics. It explained prospective “enrichment” stratification strategy to define subgroups of subjects in study proto-col according to defined diagnosis, as well as more flexible retrospective subgroup analysis after completion of study. Description in package insert (product specifica-tion) would be different according to each strategy. This guidance is mainly for genetic diagnosis but can be directly applied for imaging diagnosis. Imaging agents can be developed through a molecular targeting probe optimization process as well as pharmacological assessment and its validation, toward some diagnostic drugs, biomarker or companion diagnostics, or radioisotope therapeutic drugs.

Simultaneously in 2004, FDA issued a set of three parts of guidance to provide instructions for development of medical imaging agents for (1) nonclinical and clin-ical safety data assessment [27], (2) clinical indications [28], and (3) design, analy-sis, and interpretation of clinical studies [29] (Fig. 6.1). This set of guidance clearly showed “critical path” of imaging drug development in IND framework and how to get through and reach to the final goal of NDA. The priority of PET drug is that you can utilize both the RDRC and IND frameworks for efficient imaging drug develop-ment as well as biomarker assessment tool validation: When you have some


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information of the radiation dose within the above limitation, based on first-in-human study results under IND in the USA or under other regulatory frameworks in anywhere else, you can explore this imaging tool more in depth in RDRC frame-work. Then, if this agent is found to be promising, you can start clinical develop-ment toward regulatory approval again in the IND framework [30, 31]. Through the efforts of related communities more recently, the concept of “drug-diagnostic co- development” was switched to concept of “in vitro companion diagnostics [32]” and can be applied in molecular imaging [33]. This reflects demands for achieve-ments of more validated diagnostic products being approved by regulatory authority.

6.3.1.3 Clinical Trial Network and Standardization

Accordingly, more and more quality assurance and validation of imaging technolo-gies have been promoted in order to utilize them for the development of therapeutic drugs or otherwise developed for diagnostics approval in clinical practice. A num-ber of activities of clinical trial networking and standardization of PET imaging and PET drug manufacturing have taken place. The National Cancer Institute (NCI) carried forward their “shared IND” strategy [34] to share their IND information with those who are starting clinical trial submitting INDs to FDA (Fig. 6.2). This

INDFDA authorizationIRB authorization

RDRCRDRC authorizationIRB authorization

FDA guidance imaging/PET drugPart 1: Safety assessment

Non-clinical and clinicalPart 2: Clinical indications

A: Structure delineationB: Disease/pathology

detectioon/assessmentC: Functional, Physiological, or

Biochemical AssessmentD: Patient Management

Part 3: Clinical studies design

FDA guidance on RDRC- Basic clinical research to determine the metabolism of drug or human physiology, the pathology, and the biochemistry, without immediate therapeutic or diagnostic intention or intention of efficacy/safety assessment of a therapeutic/diagnostic drug - Mass Balance study - Molecular imaging studies to investigate biochemical processes

21CFR212FDA guidance on CGMP for PET products

USP823

Drug development Basic, exploratory clinical research

GMP

Collaboration between industry and academia

Fig. 6.1 Constructions of guidance documents by FDA concerning medical imaging drug devel-opment and radioactive drug clinical research


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means that NCI will share with others, under mutual agreement, IND information including toxicity studies and chemistry, manufacturing, and control (CMC) assess-ments, which were authorized by FDA as supporting information for the conduct of clinical trials. This is an ethical and efficient strategy to avoid duplication of unnec-essary animal experimentation to acquire toxicity data and duplication of massive paperwork for IND submission [15]. Also, the Society of Nuclear Medicine and Molecular Imaging (SNMMI)-Clinical Trial Network (CTN) has been promoting “central IND” strategy to share their IND with therapeutic drug companies under mutual agreements [15] (Fig. 6.2). SNMMI-CTN is interested in IND of PET bio-marker imaging agents, while therapeutic drug companies are interested in IND of their therapeutic drugs for which imaging agent is just a tool for their true objec-tives. Because the quality of this network has to be assured enough for the use of sponsor companies, SNMMI-CTN facilitates registration of manufacturing PET drugs and standardization and validation of PET imaging sites to be utilized by col-laborating companies.

There are other excellent clinical trial networking activities led by academic societies and universities. One prominent example is the American College of Radiology Imaging Network (ACRIN) [35], which started earlier than SNMMI- CTN, and involves a larger number of radiological physicians and scientists. Also, the Radiological Society of North America (RSNA) organizes a standardization activity group named Quantitative Imaging Biomarker Alliance (QIBA) [36]. This initiative

FDA (Food and Drug Administra�on)

SNMMIbiomarker F

NIH-NCIImaging drug F

Company BTherapeu�c drug “b”

Company CTherapeu�c drug “c”

IND to FDA

IND to FDA

CMCLOA LOA

LOA

LOA

LOA

CMC

IND INDIND

DMF

This figure doesn’t show the numbers and scales of projects but only show each relationship and usage of LOANCI is the original holder of IND of FLT and leads more various organizations of shared INDs with other manufacturers, and IND holders. SNMMI uses IND info from NCI (tox, pharmacology) based on LOA, and submits IND of FLT to FDA.SNMMI submits in their IND packet the LOAs from FLT manufacturers allowing FDA to reference required parts of their DMF for SMMI's IND; Manufacturers listed under the SNMMI IND must meet the same end-product specifications outlined in the IND application. Therapeutic drug company submits IND of therapeutic drug along with an LOA from SNMMI for IND of FLT (biomarker) and LOA from manufacturers for DMF of FLT.

LOA

(in a drug sponsored study, there is also a IND DMF/CMC for the therapeutic drug)

SNMMI-CTN

NCI’s Shared INDUsed inCancerImaging Program

ManufacturerManufacturer

Among DMF CMC info is included in IND package

Manufacturer

DMF

registration registrationsubmission submission Company ATherapeu�c drug “a”

Fig. 6.2 Shared-IND strategy of NCI and centralized-IND strategy of SNMMI


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by RSNA/QIBA is involving larger stakeholders and the scope of modalities and engaged in the development of standardized protocol of clinical trials, methodolo-gies, and procedures for imaging biomarker validation of each modality. To respond to the demand for more “precise” diagnosis, it is required to achieve higher repro-ducibility of study data, as well as to standardize methodologies and procedures to generate these data. Reproducibility can be assured by means of quality control, documentation, and storage of the study data. Standardization of quality control procedures among variety of communities is a difficult task but necessary to achieve global, simultaneous clinical development [37].

6.3.2 Precision Medicine to Achieve Global Health

6.3.2.1 Precision Medicine Initiative in the Era of GWAS and Big Data

Clinical trial networking and standardization activities have been gaining additional features, advocated by the Precision Medicine Initiative. The attractive challenge of this Initiative is to grant NIH the formulation of million or more population cohorts, which cannot only provide medical records to research communities but also infor-mation of gene profiles and metabolites and microorganisms, environmental and lifestyle data, as well as personal device and sensor data [38]. These active partici-pants are involved in the design of the initiative, which ensures the access to their own health data and empowers them to invest and manage their health. Patient involvement is facilitated with a symbolic campaign at the President’s website, to show photos and names of individual patients who struggle with diseases [1]. They also announced their progress of 6 months to honor people of “Champions of Change,” including not only researchers but also patients who contribute to this initiative [39]. The size of the cohort order is larger than the previously developed by worldwide-known biobank projects such as the United Kingdom (UK) Biobank of 0.5 million; the China Kadoorie Biobank of 0.5 million; Biobank Japan of 0.3 million; and the Taiwan Biobank of 0.2 million. Among them, the UK has launched the next phase project, named 100,000 Genome Project [40], to conduct whole- genome analyses on 100,000 genomes of 70,000 patients of the National Health Service (NHS). Genomics England, a company owned by the Department of Health, is engaged in the genetic sequencing services for this project. The stories of the first benefited family patients and other participants appear with individual names and photos, in the website of this company [41].

Another part of the $70 million grant of the USA to NCI involves large-size clinical trials of new type, such as the Molecular Analysis for Therapy Choice (NCI- MATCH) [42]. This is a nationwide 10-arm clinical trial to recruit 3000 patients of advanced solid tumors of various cancer types and lymphomas. Among these and on the basis of DNA sequence mutation analysis, about 1000 patients would be enrolled and allocated into tens arms of drugs that target distinct molecular biomarkers. This NCI-MATCH and another one called Molecular Profiling-based Assignment of


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Cancer Therapeutics (NCI-MPACT) are genetic testing-based clinical trial strate-gies called “basket” studies, where multiple tumor types with multiple single muta-tions are targeted to evaluate effects of multiple drugs, in a single trial [43]. A second type of new clinical trial design is called the “umbrella” study, where single tumor type is targeted, but multiple therapeutic strategies are evaluated in a single trial. One example is the investigation of serial studies to predict your therapeutic response with imaging and molecular analysis 2 (I-SPY 2) [44]. This is a breast cancer trial to use tissue and magnetic resonance imaging (MRI) biomarkers to test the 12 investigational drugs. These two new types of trials are conducted in collaboration with multiple sponsor companies of these drugs and genetic sequencers.

6.3.2.2 Basket-Type Clinical Trial and Imaging Archive

Again and for the concept of precision medicine, we shall seek knowledge on how the basket-type clinical trials can incorporate additional value by means of molecu-lar imaging. Being driven toward the era of precision medicine, the key message “Precision Medicine… Visualized” from WMIS in 2015 is a simple and compre-hensible catchphrase. Moreover, the European College of Radiology provided more practical and specified concepts in 2014 to argue that “imaging genomics show great potential in precision medicine” [45]. They described “radiomics” as a “high- throughput extraction of large amounts of imaging features” (sometimes from pop-ulation imaging); “imaging genomics” as a “discovery of associations between imaging phenotypes and genotypic information,” to identify imaging characteristics that indicate genetic predispositions; and “theranostics” as an intriguing new field that can correlate the power of the imaging technology with genomic information, which can help to tailor precision therapy. Later on, they developed more detailed statement on this concept [46]. It is prerequisite for imaging technology to play a key role in precision medicine not only to facilitate exploration of human physiol-ogy or drug metabolism at the molecular level (in RDRC framework) but also to expand clinical trial network strategies for drug approval (in the IND framework). This expanded networking strategy should be directed toward larger amount of data sharing to facilitate partnership with patients.

The set of ongoing NCI programs to support Precision Medicine Initiative includes projects of imaging, e.g., “Quantitative Imaging Network (QIN)” and “The Cancer Imaging Archive (TCIA)” [47]. QIN is an initiative to develop quantitative therapeutic outcome measurements among networked institutes of excellence. A recent announcement focuses on “radiomics” to develop standard operating proce-dures (SOPs) to convert descriptive, qualitative imaging techniques into inherently quantitative mineable data to connect with patient demographic, outcome, and gene expression databases. The procedures of data acquisition, segmentation, extraction, and analysis are to be standardized by this initiative [48]. TCIA is another initiative to develop a large archive of cancer imaging data accessible to the public, which includes many study results linked to The Cancer Genome Atlas (TCGA) project. The aim of the TCIA initiative is to generate multidimensional maps of the key


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genomic changes in major types and subtypes of cancer. These archived data can be used for secondary analysis and hypothesis generation, through an open-source software package [49].

6.3.2.3 Expanding the Clinical Trial and Global Standardization Networks

To facilitate precision medicine, global collaboration and standardization are key prerequisites. The rationale stems from the fact that the precise focus on one genetic variability needs to seek out a common population with the same set of characteris-tics all over the world. It is conceivable that a group of individuals who are fit for a specific therapeutic intervention may not be located in the same geographic region, and, hence, the search may need to be expanded beyond jurisdictional boarders. Contrary to that, there would be some kind of ethnic (intrinsic and extrinsic) factors, which may affect the response to some interventions that are globally utilized. This is the basic premise of many of the initiatives from global pharmaceutical compa-nies to facilitate multinational clinical development, aiming at simultaneous approv-als in multiple countries.

For this reason, the Japanese and Chinese Society of Nuclear Medicine agreed in April 2015 to develop an Asian-initiated clinical trial network, as a means to facili-tate the participation of Asian regions into multinational clinical trial initiatives [50]. Development of the Asian network will contribute to (1) more and better par-ticipation in the already existing Western-initiated clinical trial networking and standardization activities and (2) establishment of alternative networks for the development of medical technologies that are highly needed by the Asian popula-tion. The use of a given network by a researcher or a company will depend on the purpose of each study or clinical development. We should realize precision medi-cine for better health in the world. This implies that precision medicine should be for the majority of the world citizens who seek for but have not yet access to their best-fit medicine. To achieve this end, clinical trial networking and data archive along with standardization should be established from various regional perspectives and initiatives.

So, what about ethics? In the latter part of this paper, we reflect upon trend of discussion of research ethics to articulate characteristics in the era of precision med-icine, considering how ethics in molecular imaging science is discussed in these contexts.


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6.4 Ethical Consideration for Precision Medicine (Table 6.1)

Table 6.1 Summary of characteristics of ethical consideration in the era of precision medicine

Individual ethics to assure human dignity

Rigorous privacy protection

Traditional de-identification procedures may not be effective in study settings of data sharing, including whole-genome sequencing. Advancement of information technology on anonymity and informed consent process, with the recognition of the characteristics of a given research setting, is essential prerequisites.

In some settings of brain imaging studies, we need to seek not only for valid informed consent but also for advance-directive and broad consent of study subjects, with the perspective of autopsy reports after the death of the subjects, as well as surrogate consent and permission of their family members.

Right to know and right not to know

Taking a more patient involvement strategy comes with the requirement to assure participants’ right to know the results of the study by means of information sharing, which is not only among the research community but also with the participants of the study. This right also requires the ethical obligation of the researcher of managing incidental findings, e.g., a brain tumor found in the process of brain imaging.

Sometimes, patients do not want to know about the future possibility of a disease, but some of the family members or related community need to be informed. The right not to know of the people at risk of disease should be assured not to coerce diagnostic test on such people.

Presymptomatic diagnosis consultation

Needs for presymptomatic diagnosis consultation are critical when diagnosis is somewhat credible, but there are no therapeutic options.

Traditional ethical issues of diagnostic genetics are common in imaging diagnostics, particularly with regard to how the information of future disease development without therapeutic option can be ethically managed. Collaboration and integration of medical practitioners beyond disciplinary specifications are required to provide care for patients in this setting.

Collective ethics to assure social value of research

Clinical trial registration and data sharing

The initiative aims to ensure collective ethics that generate reliable research results contributing to public health and individual patient in the future. To that end, clinical trial registration has already become a regulatory requirement as well as an ethical obligation.

Data sharing is becoming an ethical obligation of investigators of clinical trial. Initiative of imaging archive is responding to this demand.

Justifiable commercialization

In the setting of clinical trials, more industry-academia collaborations are promoted, along with stricter conflict of interest management. For imaging scientists, decision of installation of costly equipment should be independent from benefit provided by the manufacturer.

In the setting of biobank and health data archive, a critical issue is the separation of the process of informed consent to donate samples or materials and the process of utilizations of these donated samples or materials.

(continued)


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6.4.1 Individual Ethics to Assure Human Dignity

6.4.1.1 Rigorous Privacy Protection

As described above, the Precision Medicine Initiative needs a large amount of data sharing associated with genetic information of individuals. Of course such initiative needs more rigorous privacy protection beyond the traditional way of data protec-tion. Patient participants have been more and more involved in this initiative in terms of the design as well as the appearance in the webpages. Whatever you achieve with their participation, it should be based on traditional ethical obligations to pro-tect the right of privacy derived from the respect to human dignity. The US Health Insurance Portability and Accountability Act (HIPAA) [51] of 1996 and the Privacy Rule [52] of 2000 stipulated the definition of protected health information (PHI) and how it could be “de-identified” (by removing some individual identifiable infor-mation) to be exempted from the regulations. The 2012 report of President’s Commission on ethics of whole-genome sequencing [10] reflected that items defined as de-identification in HIPAA regulation may not be actually enough for anonymity purposes. In such case, one option is to redefine more sophisticated ways of de-identification of the genome-sequenced data by means of information technol-ogy. Another option would be to obtain informed consent of study subjects to use their “de-identified” data, which may not be in complete anonymity. While both options may be needed in certain circumstances, more rigorous data security infra-structure is required. Additionally, some mechanisms may be needed where only a “qualified” investigator can access such information.

In case of brain imaging, there are additional issues. One issue is that MRI brain imaging data may be reconstructed to such image of face, which enable subject’s acquaintances to identify whose image it is. Development of a technology to “deface” brain MRI data has been discussed, but there would be some cases, in which such procedures are not practical. In such a case, the abovementioned informed consent comes to be required. Another issue is arising, regarding studies of Alzheimer’s disease, in which brain imaging is associated with pathological autopsy analysis. This type of studies not only need the “informed consent” of a subject for imaging examination but also “advance directive” of this subject during living time for future autopsy analysis, which has a nature of “broad consent”

Table 6.1 (continued)

Avoid exploitation and stigmatization

To avoid exploitation, we should recognize the issue of distributive justice: Research participants have equal right to access the benefit generated from research. Additionally, abuse of newly developed technologies for getting return of investment should be carefully avoided.

To avoid stigmatization, more in-depth partnership with patients and an empowerment approach are needed. It is a prerequisite to protect the population at risk of a disease found by diagnostic research, from discrimination or stigmatization by their social status and access to health insurance and social security.


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( without complete information of future studies). Furthermore, and upon the death of this subject, “legal permission (or authorization)” for autopsy by a family mem-ber of a deceased is required by law. Similarly, a “surrogate consent” by a family member may be needed to perform association studies and enable to compare brain images with tissue pathological analysis.

6.4.1.2 Right to Know and Right Not to Know

When you take a strategy of patient involvement, expanding awareness, and empow-erment to them, you need further efforts to assure their “right to know” and “right not to know.” The right to know means patients’ right to control, including access to, their own information. This right is derived from fundamental privacy right, a part of a personal right, a corollary of human dignity. A recent discussion on this issue goes beyond the traditional issue of informed consent at the time of inclusion of a study subject. The Declaration of Helsinki [53], international ethical principles for medical research involving human subjects by the World Medical Association, recommends to provide the research subject with an option of being informed about the study results. Also, the European Union’s Clinical Trial Regulation [54], imple-mented in 2016, requires the investigators to provide participants with an identifier number of clinical trial registration by which the ongoing trial information and trial results are open to public. In addition to the study’s information or outcome, there is an issue of “incidental findings.” For example, during the process of brain imag-ing for neuro-imaging study, the researcher may find a brain tumor as an unintended or incidental finding [55]. In this case the ethical question is: does this neuroscien-tist has the obligation to provide care for this brain tumor? A priori, the immediate perception is that the investigator’s ethical obligation is to deal with such incidental finding properly. Logically, if the neuroscientist does not have the expertise to pro-vide proper treatment for the brain tumor, he or she should advise this patient to consult with an oncologist. There may be other cases in which an investigator could provide care for a disease revealed as incidental findings (e.g., comorbidity dis-eases), which is called “ancillary care” [56]. The issue is particularly problematic in the case of epidemiological studies in developing country, where ancillary care may not be available in the ordinary practice [57]. To what extent the investigator has to provide care, or who should pay for this additional medical expenditure, depends on how the study’s protocol was designed and the nature of the particular situation. At the very least, the researchers are obliged to define at the time of protocol develop-ment how this kind of incidental findings should be managed, including who should pay for what.

Meanwhile, some patients do not want to know about the future possibility of serious incurable diseases, whereas family members or surrounding community members want and need to be informed. “Right not to know” is a right of a person who is at risk of future possibility of a disease, which is based on family history information. This “right” is argued in a book [58] by psychologist Nancy Wexler, who was at risk of Huntington’s disease and one of the finders of a genetic marker


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of this disease [59], causing debates on “genetic discrimination.” The UK govern-ment granted insurance companies conditional access to genetic test information of their clients, while the USA established the Act to prohibit health insurance and employment discrimination, based on genetic information. While it is still contro-versial whether the “right not to know” is a part of privacy right, the central issue is to avoid undue influence on a person to undergo certain examination, in case this person does not want to know about their risk of future development of diseases.

6.4.1.3 Presymptomatic Diagnosis Consultation

The issue of presymptomatic diagnosis has been long discussed in the context of genetic research, which is the same in case of imaging. This issue is critical when the diagnosis is somewhat credible, but there is no therapeutic option. This topic is often discussed in the context of amyloid imaging, where it is conceivable to detect Alzheimer’s disease, and there is no therapeutic option. There is the argument that such diagnosis could help the patient and family to develop a life plan. However, there is a likely possibility that the public healthcare insurance will refuse to pay for such social benefit [60]. In case the validity of diagnostics is not enough, it is justifi-able not to provide the patients with such information in which the implication is still uncertain. Meanwhile, validated presymptomatic diagnostic information with-out therapeutic option should be provided to patients, along with careful consulta-tion service.

Traditionally, genetic counseling has been provided for reproductive decision- making. Later on, such service has come to be provided in various settings of the medical practice. In this context, collaboration among a variety of medical profes-sionals is needed, e.g., primary care physicians and specialists, psychologists, and social workers. Similarly, diagnostic radiologists should take a leading role to inte-grate various disciplines related to the imaging results of a patient. The function of a qualified diagnostic radiologist in the era of precision medicine should not be limited to reading the results but rather extend his or her skills to coordinate the necessary steps for the patient’s decision-making of therapeutic options and life-style choices.

6.4.2 Collective Ethics to Assure Social Value of Research

6.4.2.1 Clinical Trial Registration and Data Sharing

Beyond traditional consideration on individual rights, research ethics are more and more expanding their scope to global public health. Three fundamental principles of research ethics were defined in the Belmont Report in 1979 [61]: (1) respect for persons, which means autonomy, derived from human dignity and applied to informed consent; (2) beneficence and non-maleficence, applied to risk-benefit


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assessment; and (3) justice, implying fair distribution of risks and benefits. This set of principles was expanded to eight principles by Emmanuel [62], focusing on global health in developing countries and now including values of “collaborative partnership” and “social value.” Social value means value generated from research and utilized in medical practice and public health. To achieve social value, a better integration of research results is needed. From this demand, ethical obligation of clinical trial registration and data sharing has been required in recent decades. Clinical trial registration is mostly facilitated by the statement of the International Committee of Medical Journal Editors (ICMJE) in 2004 [63], to set a condition of trial registration that considered the publication of trial results. Registration must be in a nonprofit, publicly available registry, and a defined set of information of the trial has to be registered prior to the first subject enrollment. This obligation is derived from reciprocity with altruism of volunteers, who trust that their participa-tion would contribute to improvement of healthcare for others. Another related rea-son of this obligation is to avoid publication bias and make positive and negative trial results publicly available. This requirement is included in the legal system of the USA [64] and EU [54] and in governmental ethical guidelines of Japan [65]. In these three regions, the obligation of registration is expanding from the outline of initial trial information toward information of revisions of the protocols, as well as trial results. Furthermore, the Declaration of Helsinki [58] included this obligation of clinical trial registration in the 2008 revision and then in the 2013 revision to expand the scope of obligated study type, from only clinical trials to any study cov-ered by the Declaration. This means that an observational study including individual- identifiable human tissue or health data initiated by physicians must be registered to some eligible registry.

In addition, responsible data sharing is becoming strengthened as an ethical obli-gation, which entails the storage of anonymous raw data from each study subject in a public repository to be shared by a responsible research community. Associations of pharmaceutical companies of the USA and Europe issued a joint policy statement [66] in 2015, based on a workshop organized in 2013 by the US Institute of Medicine (IOM). Then, IOM issued a comprehensive report in 2015 [67]. Similarly, a recent draft revision of the Ethical Guidelines of Biomedical Research by the Council for International Organizations of Medical Sciences (CIOMS) released in 2015 [68] includes the data sharing and study registration as subcategories of ethical obliga-tion of public accountability. Finally, the ICMJE issued again a statement to set the data sharing as a condition for the publication of study results [69]. The abovemen-tioned imaging archive initiatives are responding to this ethical requirement.

6.4.2.2 Justifiable Commercialization

Increasing demand for precision medicine and generation of “social value” of research requires approval of research products granted by a regulatory authority. This situation needs strengthened partnership between academic research institutes and profit-making companies, which raises the question what is justifiable and


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unjustifiable commercialization in the context of clinical trials, human sample bank-ing, and health data archive? In the context of clinical trials, there is skepticism that a commercial company may influence an academic researcher to generate biased results to benefit company’s product. Meanwhile, involvement of industry is prereq-uisite to achieve quality control or acceptable research levels for product approval. In the recent trend to facilitate industry-academia collaboration, disclosure and management of conflict of interests are becoming an ordinary practice in the medi-cal research community. In some situations, an academic researcher should decline receiving monetary incentives from industry or otherwise limit their involvement in some process of the research. In case of imaging studies, the academic researcher may be engaged in the decision-making process of the installation of costly equip-ment. Independence of such decisions from benefit-taking manufacturers of the equipment is strictly required by laws, and such legal framework is up to various jurisdictions.

In the context of biobanks, there is a question of morality in which the gover-nance framework of the society can justify commercial use of human-derived mate-rials (including samples and information). To say it simply, research use of human material by commercial companies (e.g., pharmaceutical companies or commercial research institutes) with objective of making a profit would be ethically justifiable on conditions that all the legal regulations and ethical norms are followed. On the other hand, commercial trade of human material as it is would not be ethically jus-tifiable with the exception of some conditional cases. Then, what is the borderline between these justifiable and unjustifiable cases? Commercialization of human body and human-derived information would be an infringement of human dignity, according to the Kantian philosophy, which prohibits the utilization of human as a mere tool for other objectives. For this reason, the process of the individual’s per-mission for the use of his/her material and the process of granting permission of each project to use these materials should be separated. In the former process, mon-etary inducement is prohibited to avoid moral corruption to induce an individual to sell his/her own body parts or health information. An individual will trust some nonprofit organization to store their materials, e.g., biobank institute and donate the material without intention of making a profit, by means of selling body parts. Under a strong governance framework, this biobank would permit qualified researcher or commercial entities to make use of materials trusted by individuals. This is the established scheme identified as a justifiable commercialization of human materials, discussed in the context of ethics of the biobank.

6.4.2.3 Avoid Exploitation and Stigmatization

From the abovementioned Kantian philosophy to prohibit utilization of human as a mere tool for other objectives, it is also required not to utilize a human being as a research subject, without allowing them to access the benefits generated from research. This is also a fundamental ethical dilemma, as research is conducted for the goal to generate results and contribute to public health but not for the care of


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individual research subjects. The ethical principle of “justice” in the Belmont Report [61] suggested that it is not ethical if only vulnerable people participated in the research and wealthy people can have access to better care provided as the research results. This issue of exploitation has been long discussed in the process of revisions of the Declaration of Helsinki. The 2013 revision of the declaration recommend researchers and related government to make provisions to grant post-trial access to all research participants who need an intervention identified as beneficial and to provide this information to participants during the informed consent process. This is based on recognition that research participants have equal right to access the ben-efits generated from the research. Imaging studies can be conducted mostly in wealthy people where the majority of the research participants are able to pay for diagnosis and research status not covered by public healthcare insurance. More imaginable ethical infringement in the setting of diagnostic imaging would be abuse of diagnostic services in clinical practice to get a return of investment. To avoid this kind of abuse, health technology assessment for public and private insurance cover-age is strictly demanded.

As often discussed in genetic research, diagnostic research to find mechanisms of a disease and provide some kind of prognosis but not provide care confronts the issue of stigmatization. Especially in case of utilization of the large amount of data of patients or people at risk of serious disease, we need ethical consideration to avoid collective stigmatization of some populations. To avoid stigmatization, more in-depth partnership with patients and an empowerment approach are needed. This implies that the characteristics of a patient involvement strategy must be in agree-ment with President Obama’s Precision Medicine Initiative. Imaging studies for early diagnosis of cancer or neurodegenerative disease associated with genetic information would contribute to the development of better healthcare for these peo-ple at risk. However, it is prerequisite to protect these people from discrimination or stigmatization for their social status and access to health insurance or social security.

6.5 Conclusion

We have discussed the history and evolution of the US regulatory framework and clinical development strategies related to molecular imaging, which focuses on a critical path initiative that aims at personalized medicine and then precision medi-cine to achieve individual well-being and global public health. Next, we reflected upon ethics in the era of precision medicine from both perspectives of individual and collective ethics. Some of the discussions are common in various disciplines of the medical science. In some aspects, the focusing trend is in the era of GWAS and big data analysis, whereas in other parts, the main focus is on how imaging studies play the role of activating these trends.

As for regulatory reformation and initiative of clinical development, more col-laborative approach is required in the era of precision medicine. This should include


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the basket-type clinical trial network, large-scale cohorts associated with imaging archives, and clinical outcome information. Establishment of regulatory framework along with communities’ collaboration toward quality assurance and standardiza-tion are key issues.

A reflection on ethics articulated that interests of ethical consideration are shift-ing from individual ethics toward collective ethics to achieve social value of research and global health. Traditionally, individual and collective ethics are regarded to be conflicting. However, to explore recent trends carefully, we can say that assurance of collective ethics could be simultaneously assurance of individual ethics in matured scientific communities. Nevertheless, both can still be seen conflicting in some immature research settings. This is also consistent with the theory of precision medicine, which should contribute to the well-being of an individual patient and the entire public health. This can be achieved through global collaborative research initiatives of sound scientific and ethically justifiable project designs, which aim at clinical development, profound understanding of the broad characteristics of sci-ence, clear regulatory definition attachment, and research and moral ethics in the era of precision medicine.

Acknowledgment This paper is based on our studies on PET drug-related regulations in the USA granted by the Japanese Ministry of Health, Labour and Welfare, titled “Regulatory frameworks of the United States and other countries concerning nuclear medicine diagnosis using PET drugs produced by an in-house PET drug synthesizer” in 2011 and 2012 fiscal year and another one titled “Regulatory science concerning clinical application of nuclear medicine diagnosis using PET drugs produced by an in-house PET drug synthesizer” in 2013 and 2014 fiscal year. The material discussed for this paper is based on a presentation by Chieko Kurihara: Ethics and regulation of precision medicine in nuclear medicine: Toward more international collaboration in the world. In: Precision Medicine Forum in Nuclear Medicine. Sept 26 to 27, 2015. Blue Horizon International Hotel, Linyi, Shan dong Province, China, invited by the organizer Professor Yaming Li, President of Chinese Society of Nuclear Medicine; Professor and Director, Department of Nuclear Medicine The first Hospital of China Medical University.

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Chapter 7Optical Imaging: How Far Can We Go

Wenbin Ma, Sara Herrera-Rodriguez, and Hector Martinez-Valdez

Abstract Optical imaging comprises a broad and diverse range of technology modalities, which aim to assess the morphology, dynamics, and pathophysiology state of organs, tissues, cells, organelles, and molecules, in their natural state and real time. The development of pioneer instruments for optical imaging is referenced to the seventeenth- to eighteenth-century era (circa 1632–1723), when it is pre-sumed that Antonie van Leeuwenhoek invented and used a prototype of the light microscope (Wollman AJM, Nudd R, Hedlund EG, Leake MC, Open Biol 5:150019, http://dx.doi.org/10.1098/rsob.150019, 2016). The groundbreaking discovery and its application was reported by Leeuwenhoek and Robert Hooke (Hooke R, Micrographia: or some physiological descriptions of minute bodies made by mag-nifying glasses, with observations and inquiries thereupon. Courier Corporation, New York, 1665; Baker H, Leeuwenhoek M, Philos Trans 41:503–519, doi:10.1098/rstl.1739.0085, 1739), who used a combination of stage, light source, and lenses similar to those currently used (Wollman AJM, Nudd R, Hedlund EG, Leake MC, Open Biol 5:150019, http://dx.doi.org/10.1098/rsob.150019, 2016). Fast-forward to the twentieth to twenty-first century, innovative advances underscore the extraordi-nary progression into imaging technologies, such as 3-D electron microscopy, con-focal fluorescent imaging, mass spectrometry, bioluminescence, and optoacoustics, to name a few (Weissleder R, Nahrendorf M. Proc Natl Acad Sci 112(47):14424–14428, 2015). Collectively, the new imaging modalities enable researchers to reveal complex structures and dynamic interactive processes happening deep inside cel-lular compartments, which can provide invaluable basic and clinical science infor-mation. The present chapter outlines over three centuries of optical imaging technology, as it relates to the rationale that led to the development of innovative methods, which have transformed the means to observe, analyze, study, and

W. Ma Key Laboratory of Gene Engineering of the Ministry of Education, State Key Laboratory for Biocontrol, School of Life Sciences, Sun Yat-sen University, Guangzhou 51006, China

S. Herrera-Rodriguez Centro de Investigación y Asistencia en Tecnología y Diseño del Estado de Jalisco (CIATEJ), Av. Normalista 800, Colinas de la Normal, CP 44270 Guadalajara, Jalisco, México

H. Martinez-Valdez (*) Vyripharm Pharmaceuticals, 2450 Holcombe Boulevard, Houston, TX 77021, USAe-mail: [email protected]

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diagnose the nature of cellular structures and functions. The paradigm shift inherent to the progressive advances of optical imaging and their impact on bench to bedside applications are accordingly discussed.

7.1 Development of Optical Tools

The power to accurately discern between separate objects with maximal precision defines the concept of resolution in light microscopy. However, the notion behind the resolution of the light microscope was not appreciated until the 1800s. For instance, blurring levels created by light diffraction associated to the sample under observation and the imaging capacity of the microscope were defined by George Airy [4, 142] as the point spread function (PSF). The mathematics of the PSF pat-tern, aptly coined the Airy disk, describes a central light peak intensity within feeble rings, which diffuse from the center. Consistent with Airy’s deductions, Abbe was the first to define the limits of resolution, which are equally determined by the sam-ple preparations and the instrument stage, lenses, and light source [1, 136].

When considering the Airy disk in the context of established criteria, the resolu-tion of two image points under the objective is predictable if the distance that sets them apart is equal or bigger than the disk radius [110, 136]. Examples of such reso-lution limits include the dissociation of sister chromosome kinetochores in yeast mitotic spindle [136]. Notably, Abbe went on to create the condenser to increase the illumination of the sample [2, 3, 136], but its application was short-lived as Köhler’s new standard for bright-field light microscopy was subsequently adopted [67].

Consequently, resolution became a major driving force for innovation, which transformed optical imaging, as a science and technology landmarks with broad applications in biomedicine and industry. In the process, phase-contrast microscopy was invented to resolve the low contrast caused by the light refractive index, which is almost equal to that of the water in biological specimens [147, 148]. Conversely, a method known as differential interference contrast (DIC) polarizes light within a prism into two beams at 90°, which traverse the sample and are subsequently recom-bined by another prism to create a phase difference and high-contrast images [95, 123, 142]. The resulting high-resolution images become clear and precise, which has made DIC the standard method to reveal defined cellular limits. Of note, the use of diverse labeling reagents has significantly enhanced light microscopy resolution [80], improved contrast, and broadened the range of applications for optical imag-ing (Fig. 7.1). The remarkable improvements in conventional light microcopy [80, 95, 123, 142] provided exceptional means to assess normal and pathological organ development by implementing high-definition tissue section and staining methods (Fig. 7.2). In subsequent chapters, we describe the emergence of innovative tech-nologies that have shaped the resolution, precision, and dynamics of optical imaging.

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7.2 Electron Microscopy

The aim of this section is to underscore how electron microscopy (EM) has contrib-uted to the advances of optical imaging, particularly in defining cell structure, organelle compartmentalization, protein homing, and molecular interactions and functions.

Based on the “wave of particles” hypothesis by the French physicist Louis-Victor de Broglie [29], which underscored the greater resolution of the shortwave of an electron, the idea of the electron microscope was first conceived [65, 66]. However, the first eukaryote cell EM images are credited to the Canadian-American biologist Keith Porter, despite the notion that the original transmission EM (TEM) of bacteria and virus pictures had been produced by Ruska’s brother Helmut [66, 105, 112].

The groundbreaking EM methods were straightforward and, hence, attracted the interest of researchers in the natural sciences, who wanted to examine cell contours, boundaries, and intracellular compartments. Such a task, which was only possible until resin-embedded tissues and thin section preparations, became available [28, 58, 66, 93, 97, 100, 104].

Fig. 7.1 Immunohistochemistry (IHC) of murine testis. (a) Antibody-specific staining of a protein antigen, which is selectively present in male germ cells with maximal levels (dark-brown staining) at spermatogonia-spermatocyte stage, when cells begin to populate the lumen. (b) Is a sequential IHC section of the same specimen, which was stained with an isotype-matched pre-immune (with-out antigen target) IgG to control for specificity. (c and d) Are identical images as those of a and b, only at a higher amplification as noted by the by the image bars

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Since then, EM technology has experienced a total transformation to adapt to the needs to resolve complex 3-D ultrastructure details of cells, organelles, and mole-cules [68]. Although the basic TEM foundation remained virtually unchanged, key parameters such as sample preparation, image acquisition, and data processing have significantly contributed to the 2-D to 3-D progression [39]. For instance, TEM has played a major role in resolving biological structures, such as mitochondria cristae and intermembrane spaces at sub-nanometer level (Fig. 7.3), a biochemical hub for normal and malignant cell bioenergetics, oxidative stress, and death/survival cues. On the other hand, scanning EM (SEM) has proved to be instrumental in revealing the dynamics of cell mobility and traffic, involving pseudopodia formation, adher-ence, migration, and aggregation (Fig. 7.4), a turning point in the assessment of cancer metastasis. Major advances resulting from enhanced image acquisition and data processing automation include the SEM and volume EM (VEM), which paved the way for a wide range of biomedical applications with higher resolution and

Fig. 7.2 Postnatal development in health and disease. (a) Shows an image of normal murine alveolar development from a paraffin-embedded lung section, stained with hematoxylin and eosin (H&E), as captured by standard light microscopy. (b) Reveals abnormal mouse alveolar develop-ment, characteristic of emphysema. Section preparation, H&E staining, and image capture are the same as in (a) and obtained from littermates at the same postnatal age. (c and d) Depict optical images of paraffin-embedded mouse lung sections stained with Harts solution [80], which compare alveolar elastin organization during postnatal development between normal (c) and emphysema (d) mouse littermates

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speed imaging [70, 71, 99]. As an example of a specific VEM modality, the serial section TEM (ssTEM) is a manual method intended to produce high-resolution 3-D images [49, 99]. The inherent challenge of this technique is the capture of serial snapshots from extremely thin tissue sections, which mandates dexterous skills to achieve precise image succession. On the other hand, SEM technology can generate images with outstanding resolution, given their field electron emissions combined with powerful detectors [99, 137].

Concordant with the 3-D TEM modalities, electron tomography combines the precise angle rotation of relatively thick tissue sections with a coordinate electron beam channeled through a continuous tilt span. Images are serially acquired and instantly converted into a high-resolution 3-D tomography [13, 88, 99]. To optimize

Fig. 7.3 Transmission electron microscopy (TEM) assessment of the intracellular landscape in leukemia cells. (a) A 2 μm image of the PreB acute lymphocytic leukemia cell organization, which displays a prominent nucleus- and mitochondria- rich cytoplasm, a common feature of cancer. (b) A 100 nm image shows the conspicuous presence of an antiapoptotic molecule (silver grain/black dots) within the intermembrane space of the mitochondria

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Fig. 7.4 Scanning electron microscopy (SEM) appraisal of T lymphocyte dynamics. (a) Control resting T-lymphocytes without extracellular stimulation. (b–h) Progressive detection of T lympho-cyte lamellipodia formation, rolling, migration, and aggregation after genetically engineered activation

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the ssTEM, one needs to consider the time that each image acquisition takes in the serial tomography process, to ensure that the entire area of the serial sections are captured, prevent loss of image information, and enhance resolution by deconvolut-ing serial dataset analysis [13, 99, 127]. This means that while ssTEM is a powerful technology with broad scope and depth, there is room for improvement. In agree-ment with the creative progression and in response to technological demands, upgraded hardware and software are constantly emerging to reduce the complexity and time of operation, as thoroughly reviewed earlier [99].

The notable advances in electron microscopy are underscored by the range of automation of image capture, resolution, and data processing, which enable researches of all walks of science to sort out the intricacies, molecular compartmen-talization, and dynamics of distinct tissues, cells, and organelles.

7.3 Fluorescence Microscopy

In more than two decades, fluorescence microscopy has contributed to major advances in optical imaging, and the inherent power of this remarkable technology has enabled researchers to achieve cell analyses beyond compare, in most biomedi-cal fields [40]. Further innovation and expansion of this technology rely on con-certed advances in optical hardware, software development, chemistry, cell and molecular biology, and bioengineering [40].

Early applications of fluorescence microscopy started with the observation that probes conjugated with distinct fluorescent agents could reveal the location of tar-geted molecules and their potential of complex formation, with the aim to image the occurrence of 3-D molecular interactions in real time [40, 59], under physiological and pathological conditions. Historically, the first fluorescent dye used to label cell nucleic acids in the mid-1900s was acridine orange. While the staining with acridine orange was moderately superior to conventional dyes, it would not match the speci-ficity of fluorescence-conjugated antibodies, which arrived in the 1950s along with the dichroic mirrors [25, 26, 40, 103]. This advancement was subsequently followed by the application of mercury and xenon lamps, after which were subsequently upgraded to laser beam excitation. Concomitantly, major improvements in image capture were being achieved by the invention of instant scanning and CCD process-ing [23, 40, 44, 53, 126]. These advancements enabled the exploration of histologi-cal compartments (Fig. 7.5), which then made it possible to assess the efficacy of vaccine-mediated immune responses and to diagnose malignant cell development.

Fluorescence microscopy became a key technology in optical imaging because it provided basic and biomedical researchers a dynamic means to uncover the physiology, biochemistry, and molecular intricacies of the cell [113, 121]. While fluorescence still adheres to basic optical concepts, microscope hardware design, image contrast, level of resolution, and software for data acquisition gradually

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Fig. 7.5 Confocal microscopy delineation of histological compartments. (Upper panel) Shows a strong and selective detection of B220+ B-lymphocytes after reaction with a primary anti- B200- specific monoclonal antibody and revealing the reactivity with a secondary fluorescein-conjugated anti-mouse antibody. (Center panel) Depicts a prominent and specific cell reactivity of CD3+ T cells to a primary rat anti-CD3 antibody, revealed by a secondary rhodamine-conjugated anti-rat antibody. (Lower panel) Reveals the fluorescence merge of B and T cell reactivity to respective anti-B220 and anti-CD3 antibodies, which demonstrates their selective compartments within the spleen microenvironment

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improved [113]. In this section, the principles of broad-field fluorescence micros-copy are specifically defined to underscore how different modalities such as laser scanning confocal microscopy, two-photon microscopy, scanning disk confocal microscopy, total internal reflection, and superresolution microscopy have gradu-ally contributed to the success of optical imaging.

7.4 Multiphoton Confocal Microscopy

After the discovery, purification, cloning, and functional characterization of natu-rally occurring fluorescence proteins [20, 78, 108, 119], optical imaging harnessed the power of broad-spectrum tools to provide the means to study the subtleties of molecular interactions in unparalleled detail [77, 91, 141]. The diverse array of fluo-rescent wavelength use led to a broader range of applications, including a role of the natural fluorescent proteins as molecular reporters of inducible gene expression (Fig. 7.6). The new developments rapidly led to the creation and use of high- resolution multiphoton confocal microscopy with powerful scanning features and versatile data processing software [15, 30, 52, 75, 141].

Fast-forward to decades later, the resolution and scope of fluorescence optical imaging have transformed biomedical research in unprecedented ways, for instance, the stimulated emission depletion (STED) technology capable of reaching a high confocal fluorescence microscopy resolution, by reducing the focal site of light detection to tenths of nm [17, 27, 50, 51, 74]. When data recording became a neces-sity, structured illumination microscopy (SIM) was implemented to lower the gear of high-frequency information, by shooting light in patterns onto the sample [27, 73, 94, 146]. Other upgrades such as photoactivated localization microscopy (PALM) and stochastic optical reconstruction microscopy (STORM) were similarly adapted, when precision and in-depth 3-D superresolution imaging were required [9, 16, 27].

7.5 Total Internal Reflection Fluorescence

Chemical fluidity of the plasma membrane, through the ion channels, is essential to monitor live inter- and intracellular activity with the microenvironment. In response to the demanding research interests, a new microscope with total internal reflection fluorescence (TIRF), increased resolution, and high-speed scanning capabilities was manufactured to quantitatively image, locate, and measure the activity of ion channels [8, 44, 143]. Basically, in a region in which the optical field is short-lived and cannot propagate (evanescent field), light is selectively flashed to excite fluores-cent probes to restrictively image cells in liquid media, which is confined between a glass slide and a cover slip. The overall gain of the TIRF technique is that while the incoming light is reflected by cover slips, the one that penetrates the cell media is

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evanescently absorbed [8, 44, 143]. The transient light phenomena enable the selec-tive imaging snapshot within a very focused area of no more than 200 nm, namely, the cell plasma membrane (roughly 7–10 nm) against the cover slip [8, 44, 143]. Because the probes are not in proximity to the glass and hence less likely to be reflective, the risk of fluorescence background is negligible. Consistent with these features, TIRF microscopy applications are of significant value for high- definition single-molecule imaging in live cells, at the plasma and sub-plasma membrane levels.

Fig. 7.6 Inducible green fluorescence protein (GFP) expression. (a and b) Bright-field and fluo-rescence images demonstrate the strong inhibition of GFP gene expression effect under the control of the tetracycline repressor (TetR). (c and d) The same images reveal that after tetracycline- mediated release of TetR repression, GFP is readily and strongly expressed

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7.6 Förster/Fluorescence Resonance Energy Transfer

With state-of-the-art fluorescent tools, the measurements of the distance between different proteins and their dynamic interactions were then possible to visualize in multiple colors by resonance energy transfer methods postulated nearly 70 years ago [37, 64, 91]. Simply put, Förster or fluorescence resonance energy transfer (FRET) is a nonradioactive technology, which quantitates the fluorescence energy that is exchanged between the emission of a donor molecule and the excitation of an acceptor molecule [23, 64, 92]. To that end, FRET measurements require an overlap between the emission and excitation spectra of respective donor and acceptor mol-ecules, which is key to ensure that enough energy is engaged to support protein interactions [54, 79]. Other parameters for effective FRET imaging include proxim-ity (within 1–10 nm) and orientation of fluorochrome donor-acceptor pairs [54, 64, 92, 120]. Donor molecules include fluorescent proteins, lanthanide group elements of the periodic table, and fluorescence-conjugated nanoprobes [23, 114], whereas metal quencher reagents (with or without fluorescence) and organic chemicals can function as FRET acceptor molecules [23, 72, 114].

7.7 Fluorescence In Situ Hybridization

Among the broad applications of fluorescence in optical imaging, fluorescence in situ hybridization (FISH) is distinguished for its unique property to provide infor-mation in the context of nuclear, cellular, and/or histological microenvironments [14, 62, 63, 122, 135, 138]. This powerful technology employs high-resolution methods and probes to quantitatively image DNA, RNA, and protein targets at intra-cellular compartments [10, 55, 129].

As an example, chromosomal abnormalities are recognized as genetic signa-tures, which are implicated in diverse diseases and known to affect gene dosage, structure, processing, and function. Consistent with this notion, chromosomal trans-locations are behind gene duplications, deletions, and rearrangements (Fig. 7.7), which often result in phenotypes that can either be life-threatening or have severe morphological and physiological consequences [7, 14, 122, 140]. The technology that enabled the visualization, identification, and analysis of chromosomes has been fundamental to enable accurate counts, integrity assessments, and detection of genetic abnormalities [43, 149]. Methods like chromosome G-banding (Fig. 7.7) are routine in most laboratories, where cytogenetics studies are performed with high resolution to reveal both subtle and complex gene abnormalities [131, 140]. Other applications include karyotype, chromosome gene assignment, chromatin structure, DNA recombination, gene expression, and radiation dosimetry assessments [83, 111, 145].

Detection of specific chromosome segments to structurally assess inaccessible gene lesions can be laborious and technically demanding. Hence, the need for

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innovative approaches moved science into the development of technology that made FISH a breakthrough reality [83, 111, 122, 135, 145].

FISH appeared in the laboratory scene over 30 years ago but rapidly evolved into a broad range of basic research and biomedical applications [12, 33, 60, 107, 118, 122, 135]. Progressively, florescent probes were designed by conjugating a compre-hensive array of versatile fluorochromes to achieve multiparameter cytogenetic imaging [61]. Irrespective of the application, FISH technology exploded in the past decade, with fluorochrome reagents, probe engineering, and image-capture hard-ware and software that are diverse and sophisticated [101, 115]. FISH has earned solid credibility for its chromosome/gene mapping capabilities, specificity, preci-sion, flexibility and superb microscopy, and digital imaging support, thus rapidly becoming an indispensable tool in biomedical research [101, 115]. Notably, FISH- dependent genetic queries find widespread use in a variety of scientific fields, including genetics, neurosciences, reproduction, toxicology, ecology, and evolution [101, 122, 138] to name some.

Because of the enormous diversity of the FISH technology, where acronyms are coined for any given application [101, 115, 138], one can justify the inclusion of FISH in the menu of multiple and varied analytical methods. Because the aim of the present section is to underscore the overall contribution of optical imaging to broad diagnostic assessments, only a brief appraisal of FISH applications is herein presented.

For example, when unambiguous karyotype analyses are needed and gene- specific locations, cryptic gene fusions, and intricate chromosome rearrangements

Fig. 7.7 Fluorescence in situ hybridization (FISH). (a) Shows chromosome banding depicts dis-tinctive karyotype features observed in normal leukocytes. Arrows indicate the localization of a targeted gene (yellow dots) on the two chromosome (Chr) 9 alleles, after hybridization with locus- specific genomic DNA. (b) Chr 9 (red) and 22 (yellow) painting reveals classical FISH images of a t(9;22)(q34;q11) translocation, in which segments of the BCR gene of Chr 9 are reciprocally fused with the ABL gene of Chr 22. The fusion resulting from such translocation is known as the Philadelphia (Ph) chromosome, which is commonly found in chronic myelogenous leukemia. The blue fluorescence by DNA dye Hoechst 33342 provides chromosome reference background

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have to be simultaneously defined, multicolor chromosome painting was applied [130]. Accordingly, the design of multiple fluorochromes and the development of broad wavelength range detection systems led the way to multiplex-FISH (M-FISH), which gave a record boost to cytogenetics [57, 63, 132, 138]. This innovative improvement to the FISH technology has been particularly important for cancer cytogenetics (Fig 7.7).

Consistent with the continued need for improvements, reworking of FISH proto-cols is particularly important when genetic alterations result from chromosome multimerization, as it occurs in trisomy syndromes. Because accurate assessment of micronucleation events is not a trivial feat, blocking cytoplasm partition with cyto-chalasin B (CB) in conjunction with FISH (CB-FISH) has become instrumental in assessing most forms of chromosome segregation abnormalities [46, 138].

Quantitative determination of telomere loss in aging can exploit the power of telomere hybridization, using peptide nucleic acid (PNA) FISH combined with the versatility of flow cytometry (flow-FISH), which can measure fluorescent telomere signals in cell suspensions [6, 10, 48, 90, 106]. The approach enables to manage multiple cell analyses with high resolution and has enormous clinical potential.

Assessment of chromosome loci susceptibility and DNA strand breaks is patho-physiologically important in biomedicine. To that end, a detection method known as the comet assay was developed to measure the degree of DNA breaks at the single- cell level by imaging the electrophoretic exit of DNA from the nucleus onto an agarose gel field [24, 31, 98]. When combined with FISH (comet-FISH), the proce-dure reveals the chromosome sites with relevant DNA breakage susceptibility [32, 45, 47, 125].

Similarly, the combination of FISH DNA probes with antibody probes (immuno- FISH) to detect precise gene loci has virtually unlimited potential [41, 84, 128, 144, 150]. Likewise, the accurate capture of aberrant sister chromatid exchanges by com-bining BrdU/cell cycle labeling with FISH (harlequin-FISH) advanced the technol-ogy a significant step farther [62, 86, 96].

Focused cytogenetic analysis on gene fusions resulting from chromosome rear-rangements found a niche that has relevant diagnostic and prognostic value. By using dual-color FISH probes flanking the breakpoint site of chromosomal translo-cations (split-signal FISH), precise identification of rearranged loci can be readily achieved [42, 133, 134, 138].

Notably, in situ mRNA expression using fluorescence-conjugated antisense com-plementary RNA (cRNA) probes opened new means to evaluate gene transcription (Fig. 7.8), mRNA processing, and decay [35, 81, 82, 89, 102, 138, 139]. This expression-FISH method enables to equally assess endogenous transcription, enforced plasmid-mediated overexpression, or retrovirus-dependent transduction in transgenic (Tg) animals [77, 81, 89]. The application of expression-FISH is virtu-ally unlimited, from single-cell gene expression (Fig. 7.8) to phenotypical/patho-logical differential mRNA arrays, transcriptional regulation, nuclear/cytosol traffic, and diagnosis-based transcriptome analyses [35, 77, 82, 102, 138, 139].

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Collectively, in situ hybridization technology in general and FISH in particular evolved into innovative multiparameter applications, which provide technical flex-ibility and predict advances in biomedical research.

7.8 Flow Cytometry

Looking back, the impact of the availability of antibodies has had an extraordinary impact in all fields of biology research. The subsequent progress in quantitative cell imaging by the generation of diverse fluorescence-conjugated antibody probes

Fig. 7.8 Expression fluorescence in situ hybridization (expression FISH). (Upper panel) A con-trol hematoxylin and eosin (H&E)-stained murine testis section to show the basal architecture of the tissue. (Central panel) Hybridization of a serial section of the same tissue with a red fluorescence- labeled antisense complementary RNA (cRNA) probe to assess the expression of a male germ cell-specific gene. (Lower panel) Control hybridization of a serial section of the same tissue with a red fluorescence-labeled antisense cRNA probe to confirm gene expression specific-ity. The blue fluorescence by DNA dye Hoechst 33342 provides tissue reference background

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gave rise to flow cytometry [34, 38, 80, 85, 89]. Briefly, the basic application of flow cytometry involves conjugation of fluorochromes to antibody probes, which target specific cell antigens. Then, laser beams energize the fluorochrome conju-gate, which results in fluorescence discharge that is received and quantitated by a signal detection system [117]. As it could be anticipated, different fluorochromes became readily available, and multiparameter flow cytometry was feasible, in which detectors with specific wavelength spectra concomitantly analyzed the emission of fluorochrome- conjugated antibody probes. As the flow cytometry technology continued to emerge, the applications focused on the identification of cell surface protein biomarkers, known as cluster determination (CD) antigens, which served to identify cell lineages and developmental stages [34, 38, 80, 85, 89]. As such, flow cytometry enabled researchers to accurately achieve compre-hensive phenotype analyses, discover new and unique cell populations (Fig. 7.9), and provide evidence of cell surface assembly of receptor proteins [5, 19, 21, 76, 109]. Notably, flow cytometry has demonstrated the capability to concomitantly

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Fig. 7.9 Multiparameter flow cytometry scheme to sort unique cell subsets. (a) Shows the isola-tion of hematopoietic cell progenitors (HCP) based on identity (ID) cell surface markers CD34 and CD10. (b) After sorting HCP, a specific B-cell progenitor (BCP) subset can be purified by the ID marker CD19. (c) A rare BCP subpopulation (2.9 % of total BCP) can then be isolated by its selec-tive surface expression of a surrogate κ-like ID marker

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analyze cell phenotypes and physiological activities, such as proliferation (Fig. 7.10), survival, and death [22, 34, 38, 69, 76, 80, 85, 87, 89, 109].

Unquestionably, flow cytometry is a powerful optical imaging technology that provides unique biological information to not only assess cell phenotypes and fre-quency or diversity of cell subsets but also to measure intracellular and cell surface protein expression and molecular interactions [22, 34, 69, 80, 87, 89, 109].

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As scientific demands for multicolor flow cytometry applications grew, complex-ity of wavelength spectra increased, due to high noise/signal rates from fluoro-chrome overlap and the inherent cell autofluorescence, thus limiting the number of fluorochromes that could be simultaneously used [18]. Importantly, innovation is the norm, when out-of-the-box ingenuity is at work and to unambiguously demon-strate how far optical imaging can go. Thus, in keeping with this reasoning and the need to refine flow cytometry capabilities, a mass spectrometry/flow cytometry hybrid (mass cytometry) or cytometry time-of-flight (CytTOF) technology was cre-ated, which ingeniously uses mass reporters instead of fluorochromes [116].

Basically, the CytTOF innovation entails chelator reactions, which conjugate stable metal isotopes from the lanthanide group of the periodic table of elements to as many as 40 different antibodies. The multiple probe conjugation enables the simultaneous visualization of cell surface receptor/transporters and intracellular signaling molecules, which in turn can quantitatively reveal the dynamics of bio-chemical pathways [18, 36, 116, 124]. Given the differential energy of the isotope elements, CytTOF can accurately assess ten times more parameters in a single cell than conventional flow cytometry methods, it has higher detection sensitivity, and the scope of CyTOF applications is as broad as it is versatile [18, 36, 116, 124].

Overall, the access to the innovative and continually evolving antibody probes, tagged recombinant proteins, and stable isotope mass reporters permits to undertake more challenging cell and molecular biology tasks. The range of simultaneous veri-fication of protein identity, molecular mass, and covalent protein interactions, using CyTOF’s multidimensional imaging analyses, is virtually unlimited [18, 36, 124].

7.9 Concluding Remarks

The progress attained thus far in optical imaging hardware, software, and data min-ing is a clear indication of the continued evolution of this extraordinary technology and the foundation of its impressive developing future. Accordingly, one can fore-see unprecedented opportunities for creative design of novel probes and innovative lasers, which can ultimately enable fluorescence microscopy to improve both fun-damental science and personalized precision imaging to diagnose, treat, and cure challenging pathologies, such as neurological disorders and cancer.

It is conceivable that forthcoming new instrumentation, 3-D scanning image cap-ture, and data processing will completely transform the means of quantitative opti-cal image analyses in health and disease. At present, there is no doubt that the recent advances in fluorescence microscopy have already paved the way for basic research and biomedicine to achieve detailed physiology and biochemical assessments in live cells with maximal depth and precision. It is therefore safe to predict that inno-vation will continue to lead the development of even more powerful tools and meth-ods, which will provide even more diverse means to explore in vivo intricate cell, organelle, and molecular dimensions with virtually no limits.

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Part IVClinical Aspect


Chapter 8Applications of UIH High-Resolution PET/CT in Zhongshan Hospital

Jun Zhou and Hongcheng Shi

Abstract With the advent of new generation high-resolution PET/CT (Shanghai United Imaging Healthcare Co., Ltd., UIH), the molecular imaging modality employed the cutting- edge lutetium-yttrium oxyorthosilicate (LYSO) crystal with size of 2.35 × 2.35 mm, which is of high stopping power, high light yield, and fast decay time, to gain adequate uniformity, excellent linearity, high resolution, and perfect alignment. When a PET/CT scanner is produced with high-resolution per-formance and time-of-flight technique, its clinical significance could be expressed as doctors can use a fine-resolution but sharp image to diagnose a fine and small tumor (e.g., to evaluate submillimeter pulmonary nodule) in the human body. The high-resolution time-of-flight PET/CT is a major advance with promise, which is very good news for nuclear medicine physicians and will enable us to respond to more clinical concerns with the superior technology.

With the advent of new generation high-resolution PET/CT (Shanghai United Imaging Healthcare Co., Ltd., UIH), the molecular imaging modality employed the cutting-edge lutetium-yttrium oxyorthosilicate (LYSO) crystal with size of 2.35 × 2.35 mm, which is of high stopping power, high light yield, and fast decay time, to gain adequate uniformity, excellent linearity, high resolution, and perfect alignment. When a PET/CT scanner is produced with high-resolution performance and time-of-flight technique, its clinical significance could be expressed as doctors can use a fine-resolution but sharp image to diagnose a fine and small tumor (e.g., to evaluate submillimeter pulmonary nodule) in the human body. The high- resolution time-of-flight PET/CT is a major advance with promise, which is very good news for nuclear medicine physicians and will enable us to respond to more clinical concerns with the superior technology.

UIH spared no effort to produce this high-resolution time-of-flight PET/CT scan-ner, which was soon installed in Zhongshan Hospital, Fudan University, in December 2014. Till to March 2016, 5000 patients performed examination using this high-

J. Zhou • H. Shi (*) Department of Nuclear Medicine, Zhongshan Hospital, Fudan University, Shanghai, China

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resolution time-of-flight PET/CT scanner. Its clinical effect has been widely acclaimed for high temporal resolution (faster scan speed), high spatial resolution, high-quality image, and lower radiation dose. As a matter of fact, to be able to yield a high-resolution PET images, one PET/CT scanner needs (a) high- resolution detec-tor, (b) appropriate physics corrections, and (c) advanced image reconstruction methods. The process of this scanner was validated before clinical application, and clinical evaluation was introduced as below.

8.1 Phantom Study

The image quality of UIH PET/CT was tested to surpass the industry’s highest spa-tial resolution and the standard of the National Electrical Manufacturers Association (NEMA) of 2.8 mm in the spatial resolution. In the meantime, the temporal resolu-tion achieved the industry-leading 486 picosecond. The uMI S-96R PET/CT align-ment testing data and results showed that the dual modalities had a perfect match. The detector intrinsic uniformity testing suggested that both the percent variation and the asymmetry of the frequency distribution were very low. The spatial linearity test-ing demonstrated that the absolute linearity and the differential linearity are within perfect ranges. The spatial resolution and the energy resolution were tested with good results. The Mini Deluxe Phantom was used to evaluate ultrahigh- resolution PET, which showed that the rod diameter of 2.4 mm was clearly depicted (Fig. 8.1). The Hoffman brain phantom was performed to investigate acquisition and recon-struction methods by uMI S-96R PET/CT with good image quality (Fig. 8.2).

Fig. 8.1 Mini Deluxe Phantom tested by UIH PET/CT scanner

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8.2 Animal Study

Animal studies were applied to focus on using high-resolution detector with appro-priate physics corrections and advanced image reconstruction methods to get high- quality images of great practical importance. For instance, a normal mouse scanned by uMI S-96R PET/CT with 0.5 mCi 18F–FDG and 10 min of acquisition time was demonstrated good anatomic details and FDG uptake distributions (Fig. 8.3a–d). With the same imaging parameters, a tumor-bearing nude mouse with hepatocellu-lar carcinoma showed a large, well-circumscribed soft tissue mass with heteroge-neous FDG uptake (Fig. 8.4a–b). The high-resolution modality was also used to scan dog with 18F–FDG whole body PET/CT imaging (Fig. 8.5a–d). In particular, the non-gating heart of dog with rest or stress was displayed via a cardiac scan with brilliant metabolic PET imaging (18F–FDG, 5 mCi; scanning time, 3 min; LYSO, 2.4 × 2.4 mm; and matrix, 256 × 256) and high-resolution anatomical CT imaging (kV,120 and mAs, 70). The fusion tools are helpful for displaying powerful PET/CT fusion images, which can locate papillary muscles of the heart of a dog with meta-bolic profile (Figs. 8.6a–b and 8.6c–e). F-18 NaF whole body bone imaging can display the entire skeletal system with PET, CT multiplanar reformation (MPR), PET/CT fusion, and CT 3D reconstruction modes (Fig. 8.7a–i). In a comparative study, the image quality of whole body bone imaging of a dog scanned by UIH F-18 NaF PET/CT was superior to that of the same dog by the same type of PET/CT from another manufacturer (Fig. 8.8a–f). In addition, UIH F-18 NaF PET/CT imaging

Fig. 8.2 Hoffman brain phantom tested by UIH PET/CT scanner

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can easily detect fractures of the bilateral distal femoral bones of a dog prior to and after operation (Fig. 8.9a–d).

8.3 Clinical Evaluation

The UIH high-resolution time-of-flight PET/CT provided more doctor-friendly information in oncology, infection, and cardiovascular diseases. The clinical oncol-ogy applied PET/CT to diagnose tumor, to search for unknown primary malignant tumor, to perform tumor staging and restaging, and to evaluate therapy response.

Fig. 8.3 (a–d) Normal mouse 18F–FDG PET/CT imaging

Fig. 8.4 (a–b) Tumor-bearing nude mouse 18F–FDG PET/CT imaging

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(1) There are several cases (including lung cancers, gastric cancer, and pleural mesothelioma) to show the clinical implication of the high-resolution time-of- flight PET/CT.

Fig. 8.5 (a–d) 18F–FDG dog whole body imaging

Fig. 8.6a–b 18F–FDG dog heart imaging


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Fig. 8.6c–e Fusion tools for 18F–FDG PET/CT – display with dog heart

Fig. 8.7 (a–i) 18F–NaF dog whole body bone imaging

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Fig. 8.8 (a–f) 18F–NaF dog whole body bone imaging

Fig. 8.9 (a–d) 18F–NaF dog bone imaging with the bilateral distal femoral fractures


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Fig. 8.10a–b CT images on lung window showed an oval, spiculated, mixed ground-glass opacity nodule at the middle lobe of the right lung

Fig. 8.10c–f 18F–FDG PET/CT was performed for differential diagnosis of a solitary pulmonary nodule. PET/CT images with MPR demonstrated that the solitary pulmonary nodule was of marked FDG uptake with 2.37 of SUVmax and 0.58 of SUVmean. The nodule was pathologically confirmed to be a Grade II alveolar type lung carcinoma

Case 1: A 51-year-old woman presents with no fever and cough (Figs. 8.10a–b and 8.10c–f).

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Case 2: A 74-year-old man presents with stimulating dry cough, chest tightness, and shortness of breath for a week (Figs. 8.11a–c and 8.11d–l).

Fig. 8.11a–c Axial non-enhanced chest CT images on lung window (2A) and mediastinal window (2B) showed marked tapering and obstruction in the initial segment of the bronchus of the right upper lobe. Axial contrast-enhanced chest CT scan (2C) showed a segmental atelectasis of the right upper lobe caused by a mass in the right hilum. 18F–FDG PET/CT was performed for differential diagnosis of the mass

Fig. 8.11d–l PET/CT images with MPR demonstrated that the right hilar mass was of marked FDG uptake. Lung cancer was confirmed by fiber-optic bronchoscopy with biopsy. Then, a right upper lobectomy was performed, and the mass was pathologically verified to be squamous cell carcinoma. Two lymph node metastases were found in a regional lymph node dissection


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Fig. 8.12 (a–c) An axial non-enhanced upper abdominal CT image with prone position (a) dis-played a focal thickening with ulcerative appearance in the lesser curvature of gastric wall. An axial contrast-enhanced upper abdominal CT scan with supine position (b) showed marked enhancement of the mass in the lesser curvature of gastric wall. Then, the patient underwent gas-troscopy and three ulcer lesions were found within the wall of stomach. PET/CT scan was per-formed for tumor staging. PET/CT fusion image with prone position (c) showed the ulcerative mass was of marked FDG uptake. The mass was pathologically verified to be a poorly differenti-ated ulcer-type adenocarcinoma

Fig. 8.13 (a–d) A coronal non-enhanced CT image (a) demonstrated diffuse left pleural thicken-ing with a small amount of effusion. 18F–FDG PET/CT was performed for differential diagnosis and staging. Coronal PET image (b), PET/CT image (c), and PET image with maximum intensity projection (d) demonstrated that marked diffuse FDG uptake lesions were noted within the left pleural cavity and peritoneal cavity. Malignant mesothelioma of pleura and peritoneum was pre-sumed, and the outpatient patient rejected a further diagnosis

Case 3: A 66-year-old man presents with a persistent stomachache for 2 months (Fig. 8.12).

Case 4: A 46-year-old man presents with serious chest pain (Fig. 8.13).

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(2) Searching for unknown primary malignant tumor is another wonderful clinical application for PET/CT. We can take another two cases for example.

Case 5: A 62-year-old man who has a lump with gradual enlargement in his right neck for 3 months presents with no history of malignancy. Chest CT scan showed nothing significant (not shown). Gastroscopy was performed and no malignant evidence was found. PET/CT scan was applied for searching primary tumor (Fig. 8.14).

Fig. 8.14 (a–f) The top three UIH coronal PET/CT images (a–c) showed that multiple lymphade-nopathies with marked FDG uptake were well displayed with specific morphological features, which is superior to the bottom three coronal PET/CT images (d–f) with 1 h delay scan from another scanner. The right upper mediastinal mass was pathologically verified to be a poorly dif-ferentiated thyroid carcinoma


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Case 6: A 65-year-old man with unexplained syncope underwent brain MRI exam-ination. Multiple metastatic lesions were founded on MRI (not shown). PET/CT was performed for searching for an unknown primary tumor (Fig. 8.15).

(3) Tumor staging (e.g., breast cancer, ovarian cancer, and liposarcoma) and tumor restaging after therapy are one of the most important clinical applications for PET/CT.

Fig. 8.15 (a–e) PET/CT and PET MIP images showed that multiple marked FDG uptake mass was identified in the bilateral occipital lobes, the bilateral necks, the right axillary fossa, the bilat-eral hilar regions, the mediastinum, the mid-lower esophagus, and the peritoneal cavity. The lesions were pathologically verified to be small cell carcinoma and mixed squamous cell carcinoma

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Case 7: A 55-year-old woman presents with lower abdominal pain and increased stool frequency recently. The patient also reports about 3 kg weight loss within last 1 month. Abdominal ultrasonography demonstrates bilateral ovarian masses, uterine leiomyoma, and a small amount of fluid accumulation in the pelvic cavity. PET/CT scan was applied for diagnosis and staging (Fig. 8.16).

Additional Findings Laparoscopic examination showed a 4 cm*4 cm*3 cm soft tissue mass in the left ovary. The right ovary adhered to adjacent organs, the greater omentum looked like biscuits, and diffuse miliary nodules were seen on the surface of the liver.

Fig. 8.16 (a–g) Axial lower abdominal CT image (a) and PET/CT image (b) demonstrated that an enlarged lymph node (within white circle) with increased FDG uptake was noted in the ileocecal region. Axial contrast-enhanced lower abdominal MRI image (c) and PET/CT image (d) showed that an enlarged lymph node (within white circle) was seen adjacent to the right anterior perito-neum. Coronal non-enhanced CT image and PET/CT image (e–f) showed that a large low density mass with markedly increased FDG uptake was identified in the right lobe of liver. Coronal PET image with maximum intensity projection showed that multiple masses with lymphadenopathies were revealed in the entire abdomen


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Case 8: A 35-year-old man underwent a mass resection of the right leg for suffer-ing from myxoid liposarcoma. Recent CT imaging found suspected metastatic lesions in the bilateral lungs. PET/CT scan was performed for staging (Fig .8.17).

Case 9: A 50-year-old woman presents with a right breast mass (Fig. 8.18).

(4) The tumor therapy response evaluation based on PET/CT is a more feasible and reliable method at the present day. Especially, surveillance of molecular tar-geted tumor treatment is one of the most important clinical significance for PET/CT, which is a prognostic tool to tumor management in the past decade.

Fig. 8.18 (a–j) PET/CT scan was applied for diagnosis and staging. PET/CT images with multi-planar reformat (a–j) demonstrated an ill-defined, high FDG uptake, and soft tissue mass in the right breast with retraction of the adjacent skin, a moderate FDG uptake pulmonary nodule, mul-tiple high FDG uptake masses in the liver, and multiple high FDG uptake osteolytic lesions within the thoracic vertebrae

Fig. 8.17 (a–h) PET/CT images (a–h) demonstrated a well-defined, homogeneous high FDG uptake, and mixed solid and cystic myxoid iposarcoma behind the right distal femur prior to tumor resection, and an ill-defined, heterogeneous high FDG uptake, and low density soft tissue mass in the left pelvic cavity and multiple high FDG uptake pulmonary nodules in the bilateral lungs with bilateral pleural effusions

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Case 10: A 59-year-old man who was diagnosed with lymphoma half a year ago. PET/CT was performed again for therapy response evaluation and restaging after systemic chemotherapy (Fig. 8.19).

Case 11: A 38-year-old man was diagnosed with diffuse large B cell type non- Hodgkin’s lymphoma of the left tonsil. 18F–FDG PET/CT was performed for stag-ing before therapy and for therapy response evaluation after three-cycle chemotherapy (Fig. 8.20).

Fig. 8.19 (a–f) Coronal PET/CT images demonstrated that multiple high FDG uptake soft tissue masses or lymph nodes were noted in the bilateral axillary fossae and the bilateral inguinal fossae on September 26, 2014 (a), and then lymphoma progressed to the bilateral neck, mediastinum, peritoneal cavity, and the surface of liver on March 4, 2015 (b), although several lesions in the bilateral inguinal fossae disappeared after treatment (10B). The coronal PET/CT images (c–f) showed that multiple lymphadenopathies with marked FDG uptake before systemic chemotherapy. The UIH PET MIP image (f) well displayed multiple lymphadenopathies with specific morpho-logical features, which is superior to the PET MIP image from another scanner

Fig. 8.20 (a–j) PET/CT images with MPR demonstrated that a large, high FDG uptake, soft tissue mass was seen in the left tonsil (a–e), and the mass vanished after treatment for non-Hodgkin’s lymphoma (f–j)


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(5) Infection is a common disease in clinical settings. Most of infectious diseases may manifest as marked high FDG uptake such as tuberculosis, lung abscess, aspergillosis, bacterial pneumonia, and so forth.

Case 12: A 61-year-old man accepted aortic valve replacement operation in April 2014. The patient felt swelling and pain in the region of chest incision 4 months later after operation. Incision and drainage operation was performed in October 2014. The patient felt even worse swelling and pain in the region of chest incision again 1 year later after operation (Fig. 8.21a–c). Skin ulceration appeared and pus outflow was found soon. PET/CT scan was performed for evaluating infection lesions (Fig. 8.21b–d). Infection focus removal operation was performed. Acute and chronic inflammatory cell infiltrations as well as microabscesses were found in the removal chest wall.

Case 13: A 62-year-old woman accepted total arch replacement and descending aorta intraluminal stent grafting for suffering type A dissecting aneurysm. The swelling and pus outflow in the region of chest incision were found repeatedly. Exploration of the infection of the aortic vascular prosthesis and reconstruction of

Fig. 8.22 (a–h)The top four PET/CT images (a–d) showed that multiple marked FDG uptake lesions, which severity of infection is lower than that of the bottom four PET/CT images after intervention (e–h), were seen in the region of chest incision and mid-anterior mediastinum

Fig. 8.21 (a–d) Pre-contrast CT scan images (a–b) with mediastinal window showed persistent infection in the region of lower sternum within 1 year after operation. PET/CT images revealed the severity and the extent of infection, which were displayed hot spots

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chest wall with muscle flap transfer were performed in May 2015. PET/CT scan was performed for evaluating infection lesions before therapy and evaluating the therapy response after therapy. Pathology was verified to be chronic inflammatory cell infil-tration and giant cell reaction was found in the removal chest wall (Fig. 8.22).

(6) As a very important method for diagnostics and assessment of prognosis in ischemic heart disease, the clinical value of PET/CT is gradually getting much attention in cardiovascular disease.

Case 14: A 54-year-old man was performed PCI in ER for suffering from ACS. Coronary artery angiography showed that 95 % stenosis was noted in the middle segment of the left anterior descending branch with blood thrombosis, 75 % stenosis in the middle segment of the first diagonal branch, and 80 % stenosis in the proximal segment of the second diagonal branch of the LAD coronary artery. A stent was input in the middle segment of the LAD after thrombus aspiration. Myocardial perfusion and FDG imaging were performed for evaluating myocardial viability (Fig. 8.23).

Summary UIH PET/CT scanner in Zhongshan Hospital is widely used in clinical settings. The high-resolution time-of-flight PET/CT scanner can provide high- quality diagnostic imaging. By now it is stable for use, and its scan time is also reasonable. Moreover, the HyperView (23.6 cm) is convenient to use for patients. In addition, the scanner for small animal study is extra benefits. We think that the UIH high-resolution time-of-flight PET/CT scanner will be widely used in clinical prac-tice and benefit more patients with promise.


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Fig. 8.23 (a–b) Resting myocardiac 18F–FDG PET perfusion imaging perfectly matched with 99mTc-sestamibi cardiac scan images (a). The defects in the anterior and inferior wall, anterior and septal walls at the apex, and apex were seen in the left anterior descending territory (b)

J. Zhou and H. Shi


Chapter 9Clinical Evaluation of Focused High- Resolution Breast PET

Kanae Kawai Miyake and Yuji Nakamoto

Abstract Breast cancer has high incidence among women worldwide. Previous studies indicate that conventional whole-body positron emission tomography (PET) with 18F-fluorodeoxyglucose (FDG) can be used to detect metastasis in patients with breast cancer. However, it may not perform well in the assessment of the primary site, mainly due to limited spatial resolution. To circumvent this limitation, some groups have developed high-resolution PET systems that are specifically designed for breast evaluation. In this chapter, we review features of dedicated breast PET systems and present examples of clinical studies performed thus far. These include our clinical experiences with a comprehensive breast PET system, using a ring- shaped scanner. Future developments related to specific breast PET systems are also discussed.

Keywords Dedicated PET • High resolution • Breast cancer

9.1 Introduction

Breast cancer is one of the most common malignancies among women worldwide, with estimates of 232,340 new cases of invasive breast cancer in the United States and 1.8 million new cancer cases from 188 countries in 2013 [13, 17]. Several ran-domized, controlled trials in Europe and the United States showed that mammogra-phy screens contributed to a reduction of breast cancer mortality by 25–48% [9, 35, 37], underscoring the benefits of early detection using imaging techniques. Individualized minimal invasive multidisciplinary approach, in conjunction with lumpectomy, sentinel lymph node biopsy, radiation therapy, and/or systemic ther-apy, is one of the current treatment trends for breast cancer [18, 19]. Thus, imaging techniques that allow early detection to tailor personalized therapy may play an important role in breast cancer diagnosis and prognosis.

K.K. Miyake • Y. Nakamoto (*) Department of Diagnostic Imaging and Nuclear Medicine, Kyoto University Graduate School of Medicine, 54 Shogoinkawahara-cho, Sakyo-ku, Kyoto 606-8507, Japane-mail: [email protected]


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Positron emission tomography (PET) is a useful technique, which enables to visualize biologic functional processes and molecular features within the human body. Potential clinical roles of PET techniques for breast cancer may include the early detection, problem solving in difficult mammographic cases, staging, restag-ing, treatment monitoring, and prediction of treatment efficacy. However, whole- body PET imaging results with 18F-fluorodeoxyglucose (FDG) have revealed the limited diagnostic performance of conventional PET imaging in the evaluation of primary breast cancers [10], mainly due to the low spatial resolution of PET scan-ners. For instance, sensitivity for the visualization of primary tumors smaller than 1 cm was shown to be low [2]. To overcome this problem, high-resolution PET systems for focal assessment of breast lesions have been developed by several groups.

9.2 Types of Focused Breast PET Scanners

A dedicated breast PET system provides high-resolution PET images with a limited field of view (FOV). In comparison to conventional whole-body PET, focused breast PET systems offer unique structural and computing features, such as close proxim-ity to the breast, small detector elements, depth-of-interaction (DOI) measurement capability, and sophisticated reconstruction algorithms, which provide higher spa-tial resolution and higher photon sensitivity. In keeping with this notion, the design of dedicated breast PET scanners that have been developed so far can be comprised into two groups [8]: positron emission mammography (PEM) (Fig. 9.1a) and fully tomographic dedicated PET scanner (Fig. 9.1b, c). A list of manufactures is pro-vided in Table 9.1.

PEM has two detector heads integrated with planar or curved breast compression paddles, which acquire limited-angle tomographic images from incomplete 3D data obtained with a mildly compressed breast (Fig. 9.1a). The positioning of the breast is similar to that in mammography, and two projections (craniocaudal and mediolateral oblique) are usually obtained. The Flex Solo II scanner (Naviscan, San Diego, USA)

A B C

tumortumor

tumor

detector

detectordetector

detector

Ring-shaped detector

Fig. 9.1 Types of dedicated breast PET scanners. (a) Positron emission mammography (PEM). (b) Fully tomographic dedicated breast PET with dual-plate detector rotating around the breast. (c) Fully tomographic dedicated breast PET with ring-shaped detector

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is a PEM device that was cleared for marketing by the US Food and Drug Administration (FDA) in 2009 [24] and is currently the most common and com-mercially available dedicated breast PET system in the United States. Usually, a scanning time for one projection is 7–10 min [25, 34]. The in-plane spatial resolu-tion was reported to be 2.4 ± 0.3 mm full width at half maximum (FWHM) for images reconstructed with a three-dimensional (3D) list-mode maximum likelihood expectation maximization (MLEM) algorithm. On the other hand, the cross plane spatial resolution is low, with a reported FWHM of 8.0–8.2 mm [24, 25], but attenu-ation and scatter corrections are available. This system is equipped with quantitative metrics to measure quantitative values, called PEM uptake values (PUV) [36]. A biopsy capability is included in this model.

The fully tomographic dedicated PET systems are newer generations of PET devices that acquire complete 3D data from an uncompressed breast, and there are several variations in the detector design. One is a dual- or multi-plate detector rotat-ing around the breast (Fig. 9.1b), such as Clear-PEM developed by the Portuguese consortium under the framework of the Crystal Clear Collaboration at CERN [1] and PEM/PET system developed at West Virginia University [32]. Another type of the fully tomographic dedicated PET systems is a ring-shaped detector encircling the breast (Fig. 9.1c), which includes MAMMI (Oncovision, Valencia, Spain) and Elmammo (Shimadzu, Kyoto, Japan).

In Kyoto University, authors examined a scanner performance of the Elmammo prototype (Shimadzu, Kyoto, Japan) and have been performing human imaging with this system since 2009 [26]. Elmammo has a complete ring-shaped detector, consisting of 36 detector blocks arranged in three contiguous rings with 12 detector modules, which have a transaxial diameter of 185 mm and an axial FOV of 155.5 mm (Fig. 9.2). Each detector block has four layers with a 32 × 32 array of 1.4 × 1.4 × 4.5 mm3 lutetium gadolinium oxyorthosilicate (LGSO) crystals, coupled to a 64-channel position-sensitive photomultiplier tubes (PSPMT). This system has also DOI measurement capability. Elmammo is one of those systems that can achieve the highest spatial resolution among PET systems for human applications. With the Elmammo prototype, minimal FWHMs in the radial, tangential, and axial directions are 1.6, 1.7, and 2.0 mm and 0.8, 0.8, and 0.8 mm, for filtered back projection (FBP) and 3D dynamic row-action maximum likelihood algorithm (DRAMA) reconstruc-tions, respectively [26]. These values are much smaller than the FWHM obtained with whole-body PET or PET/CT systems of approximately 5–7 mm [7, 12]. Elmammo has capabilities of attenuation and scatter corrections and holds a

Table 9.1 Examples of dedicated breast (db) PET systems

PEMFully tomographic dedicated breast PETRotating planar head type Ring-shaped detector type

PEM Flex Solo II (Naviscan)High-resolution PEM (Stanford U)

Clear-PEM (Hospital of Portuguese Institute of Oncology)PEM/PET (West Virginia U)dbPET/CT (UC Davis)

MAMMI (Oncovision)Elmammo (Shimadzu)Partial ring dbPET (Pennsylvania U)

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quantitative metrics to obtain standard uptake value (SUV). A breast is usually scanned for 5 min with a patient lying in a prone position. Because the axial FOV of the ring-shaped scanner is designed to be large enough to cover the entire depth of a breast, the whole image is obtained at once without changing the position of the detector. Elmammo is commercially available since 2014 in Japan.

9.3 Clinical Values of Focused Breast PET Systems

In comparison to conventional whole-body PET, the potential advantages of the high-resolution PET systems may include (1) improvement of detection of breast cancer uptake, (2) visualization of detailed distribution of PET tracers accumulated in breast lesions, and (3) quantification of tumor uptake with less biased interpreta-tion, resulting from the partial volume effect. Although available published data are still limited, there have been several published studies that addressed the diagnostic values of dedicated breast PET with 18F-FDG in the detection of breast cancers in patients with suspicious lesions or established cancers. However, no or very few studies have addressed the diagnostic values of dedicated PET systems in the screening of asymptomatic women or assessment of treatment efficacy. In the fol-lowing section, we summarize results of reported studies and describe potential advantages and disadvantages for the clinical use of dedicated breast PET.

Fig. 9.2 Elmammo prototype (Shimadzu, Kyoto, Japan). This is a fully tomographic dedicated breast PET system with ring-shaped detectors. A patient lies in the prone position with her breast in the aperture of the ring-shaped detector (Cited from Miyake et al. [26])


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9.3.1 Diagnostic value in Patients with Confirmed Breast Cancers

Several studies have addressed the diagnostic values of dedicated breast PET with 18F-FDG in the detection of breast cancers in patients with known or suspected cancer lesions. As summarized in Table 9.2, most of these studies were performed with PEM scanners, which showed sensitivities of 80–95% and varied specificities that ranged from 33 to 100% in the detection of breast cancers [4, 14, 23, 28, 33]. There has been one study for a ring-shaped dedicated PET device (Elmammo prototype) that showed a sensitivity of 82% and a specificity of 50% [21]. Recently, a metaanalysis was performed for a total of eight PEM articles published between 2000 and 2012, which showed pooled sensitivities and specificities on lesion-basis analysis of 85% (95% CI, 83–88%) and 79% (95% CI, 74–83%), respectively, with area under the curve (AUC) of 0.88 [11]. These results are highly similar or slightly better than those of whole-body PET, with a sensitivity of 83% (95% CI, 73–89%) and a specificity of 74% (95% CI, 58–86%) [10].

However, these studies enrolled patients with confirmed breast cancers or suspi-cious lesions diagnosed based on conventional breast examinations, such as clinical diagnosis and mammography. Hence, overall sensitivities could have been elevated due to the high ratio of relatively large cancers. From a clinical point of view, in patients with recently diagnosed breast cancers, the main role of the dedicated breast PET may be the detection of additional lesions, because the new findings may change the treatment strategy, particularly in cases in which a conservative therapy is envisioned. Thus, some studies provided data with stratifying lesions to known index cancers and additional lesions in the ipsilateral or contralateral breasts.

For known index breast cancers, reported sensitivities of 18F-FDG PEM are uni-formly high, ranging from 92 to 95% (Table 9.3) [4, 5, 22, 34]. Two studies showed the sensitivities of PEM for index cancers were significantly higher than those of whole-body PET (56–68%) [22, 34]. In addition, Kalinyak et al. showed PEM (95%) had higher sensitivity than whole-body PET/CT (87%, p = 0.03) [22].

Table 9.2 Overall diagnostic performance of 18F-FDG dbPET in detection of cancers in patients with suspected or known breast cancer

Authors Year dbPET type No. of patients Sensitivitya Specificitya

Murthy et al. 2000 PEM 16 80% 100%Levine et al. 2003 PEM 16 86% 91%Rosen et al. 2005 PEM 23 86% 33%Berg et al. 2006 PEM 94 90% 86%Eo et al. 2012 PEM 101 95% NAIima et al. 2012 Ring shaped 69 82% 50%

dbPET dedicated breast PET, NA not availableaLesion-basis analysis


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For additional ipsilateral breast cancer, PEM had sensitivity of 41–85% and high specificity ranging from 74 to 91% (Table 9.4) [5, 22, 34]. Although the sensitivity of PEM was still limited, Kalinyak and colleagues demonstrated that PEM (47–57%) was more sensitive than either whole-body PET (6.7%, p < 0.05) or whole- body PET/CT (13%, p < 0.01) [22] (Table 9.4).

In the detection of additional breast cancers in the contralateral breasts, Berg et al. showed sensitivities of 18F-FDG PEM were only 20% (3 out of 15) in the pro-spective reading session and 73% (11 out of 15) in the retrospective reading ses-sions [6].

Detection and visualization of sub-centimeter breast cancers (≤10 mm) have been challenging issues for conventional PET imaging. Whole-body PET has a lim-ited sensitivity for sub-centimeter cancers, with reported sensitivities of 0% for T1a (>1 mm and ≤5 mm) invasive cancers and 13–39% for T1b (>5 mm and ≤ 10 mm) invasive cancers [3, 22]. However integrated whole-body PET/CT could provide improved sensitivities of 0–40% for T1a and 71–83% for T1b [21, 22]. In dedicated breast PET systems, according to data from small subpopulations, sensitivities of dedicated breast PET range 25–100% (average 46% for 28 reported lesions) for T1a

Table 9.3 Sensitivities of 18F-FDG PET systems for index breast cancer

Authors Year No. of patients Analysis PEM wbPET wbPET/CT

Berg et al. 2006 77 Lesion basis 93%Kalinyak et al. 2014 69 Breast basis 92%a 56%

109 Breast basis 95%b 87%Schilling et al. 2011 208 Lesion basis 93%a 68%Berg et al. 2011 388 Lesion basis 93%

wbPET whole-body PET, NA not availableaSignificantly superior compared to whole-body PET (p < 0.05)bSignificantly superior compared to whole-body PET/CT (p < 0.05)

Table 9.4 Diagnostic performance of 18F-FDG PET systems in detection of additional ipsilateral breast cancer

Authors YearNo. of patientss Analysis

Sensitivity SpecificityPEM wbPET wbPET/CT PEM wbPET wbPET/CT

Kalinyak et al.

2014 69 Breast basis

47%a 6.7% 91% 96%

109 Breast basis

57%a 13% 91% 95%

Schilling et al.

2011 208 Lesion basis

85% NA 74% NA

Berg et al.

2011 388 Breast basis

51% 91%

Lesion basis

41% 80%

wbPET whole-body PET, NA not available (although scan was performed)aSignificantly superior compared to the other modality (p < 0.05)


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and 46–86% (average 81% for 86 reported lesions) for T1b (Table 9.5) [4, 5, 21, 22, 34]. Collectively, these data suggest that dedicated breast PET scanners are more sensitive than conventional PET scanners in the detection of sub-centimeter tumors. Kalinyak and colleagues demonstrated that in T1b cancers of index tumors, PEM has a significantly higher sensitivity than whole-body PET in the same popu-lation (95% versus 37%, n = 19, p = 0.002) [22].

In summary, current data may indicate that dedicated PET systems could yield better ability to visualize small cancers and additional ipsilateral breast cancers in patients with known breast cancer, compared to conventional whole-body PET systems. In our experience, we encountered several cases in whom PET with ring- shaped scanner (Elmammo prototype) performed better than conventional whole-body PET, in visualizing a sub-centimeter breast cancer (Fig. 9.3) and mul-tiple additional cancers in the affected breast (Fig. 9.4).

It remains to be determined if dedicated breast PET practically contributes to the breast cancer detection in the current clinical setting, in which breast MRI is usually the second-line imaging after the application of conventional mammography and ultrasonography to diagnose breast cancers. Breast MRI is known to have high sen-sitivity for primary breast cancers. Notably, several studies that compared the diag-nostic performance between PEM and breast MRI in patients with known or suspected breast cancer showed no difference in sensitivity [5, 6, 34]. Thus, it might be too optimistic to think dedicated PET with 18F-FDG can replace breast MRI. Some practitioners suggest the clinical role of dedicated breast PET with 18F-FDG may be an alternative examination to breast MRI in patients who are not able to tolerate breast MRI [6, 11, 15, 16].

Emerging technologies, such as multimodality imaging and application of new PET tracers, potentially allow further improvement of the breast cancer detection with dedicated PET systems. Berg et al. demonstrated that integration of PEM and

Table 9.5 Sensitivities of 18F-FDG dbPET for sub-centimeter invasive cancer

Authors Year dbPET type

T1a (>1 and ≤5 mm) T1b (>5 and ≤10 mm)No. of lesions

Sensitivitya (%)

No. of lesions

Sensitivitya (%)

Berg et al. 2006 PEM 2 50 6 67Schilling et al.

2011 PEM 4 100 21 86

Berg et al. 2011 PEM 16 25 13 46Kalinyak et al.

2014 PEM 4 75 40 95

Iima et al. 2012 Ring- shaped

2 50 86 81

Total 28 46 86 81

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MR imaging increased cancer detection, compared to MR imaging alone (74% vs. 60%, p < 0.001) [5]. With the development of PET agents, dedicated breast PET could be used for the evaluation of expression status of molecules for targeted ther-apy, such as hormone receptors and human epidermal growth factor receptor 2 (HER2), and the functional status, such as tumor hypoxia. Making the most of the unique features of PET as functional imaging may provide clues to identify and establish a role of dedicated breast PET, which may be distinctive from that of breast MRI.

9.3.2 Screening of Breast Cancer in Asymptomatic Females

To the best of our knowledge, there have been no published studies investigating the diagnostic performance of dedicated PET systems in screening of asymptomatic females for the diagnosis of breast cancer. Given that dedicated breast PET can

Fig. 9.3 A 55-year-old female with a 6-mm breast cancer in the right breast: Histology, invasive ductal carcinoma. While no apparent abnormal uptake is observed on whole-body PET/CT (a; maximal intensity projection PET image) after the administration of 18F-FDG, dedicated breast PET (Elmammo prototype.; standard reconstruction mode) shows focal uptake (arrows in b [coro-nal image] and c [axial image]) corresponding to the cancer which enhances on MRI (arrow in d).


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visualize small cancers better than whole-body PET, the former may improve the detection rate of breast cancers in broad screening populations. However, radiation exposure and high operating cost could hinder the widespread use of dedicated breast PET systems as a screening modality. It has been estimated that common dose of 18F-FDG, used in either dedicated breast PET or whole-body PET, is associ-ated with higher radiation exposure than screening mammography, which creates a higher risk of radiation-induced cancer [20, 29]. Further discussions are needed to assess the benefits and the risks that are inherent to the application of dedicated PET systems for the breast cancer screening program. Recently, private cancer screening projects with whole-body PET/CT or PET/MRI are getting common in Japan. Exceptionally in this setting, dedicated breast PET can be easily used as an addtional study without further radiation exposure.

9.3.3 Distribution Assessment

Intratumoral heterogeneity is one of the hot topics in the PET oncology field. Like in other tumors, breast cancer may exhibit intra-tumor structural and functional heterogeneity, which may also be influenced by distinct gene expression patterns. Such heterogeneity could conceivably lead to altered distribution of PET tracers within the tumor. Dedicated breast PET may allow the visualization of more detailed distribution of PET tracers accumulated in the breast, compared to conventional whole-body PET. Figure 9.5 is an example in which dedicated breast PET visualized

Fig. 9.4 A 58-year-old female with multifocal breast cancer: Histology, invasive ductal carci-noma. Compared to 18F-FDG whole-body PET (a, coronal image), more cancer foci are visualized on 18F-FDG dedicated breast PET (b, Elmammo prototype; enhanced-resolution reconstruction mode; coronal image)


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detailed distribution of 18F-FDG within a tumor. A high-resolution distribution map of PET tracers may be useful to determine appropriate biopsy sites. In manner, mor-phological PET findings that are specific for breast cancer could be identified and thus generate precise and detailed PET-MRI fusion images, which may provide a better understanding to personalized physiopathology of breast cancers.

9.3.4 Quantitative Analysis

Quantitative values such as PUV and SUV can be obtained with some dedicated breast PET systems. Increase of spatial resolutions potentially contributes to the reduction of the bias caused by the partial volume effect, which commonly under-estimates quantitative values in smaller lesions on whole-body PET images. Reliability of quantitative values is under investigation in some systems.

9.3.5 Dedicated Breast PET-Guided Biopsy

Some of dedicated PET devices have been equipped with a biopsy capability. Biopsy capability has a merit in enabling histologic evaluation of each suspicious or equivocal findings on dedicated breast PET. Because the half-life of 18F-FDG is 110 min, 18F-FDG accumulated in a target lesion, which is visualized on dedicated breast PET, serves as a marker during the biopsy procedure.

Fig. 9.5 A 73-year-old female with a solid and cystic breast cancer: Histology, invasive ductal carcinoma. Maximal intensity projection (a) of dedicated breast PET (Elmammo prototype; stan-dard reconstruction mode) shows heterogeneous accumulation of 18F-FDG to the tumor. On sagit-tal image of dedicated breast PET (b), intense uptake is observed corresponding to the enhancing solid parts of the tumor on MRI (c), indicating that detailed geographic distribution of FDG accu-mulation can be visualized on dedicated breast PET


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9.4 Future Visions

One of the future trends in dedicated breast PET imaging may be multimodality imaging. Integration of this device to the other imaging modalities may hold advan-tages to make it easier to correlate dedicated PET images to other findings and provide anatomical maps that may be helpful for the accurate interpretation of func-tional information on dedicated PET. These maps could also be instrumental to compensate for the technical and diagnostic limitations of dedicated PET, such as FOV at the deep breast near the chest wall or the occurrence of false negatives. There have been several ongoing projects to generate integrated systems. In the University of California, Davis, a hybrid system of dedicated PET and CT scanning (dedicated breast PET/CT) has been developed [1]. In Washington University, a high-resolution PET insert that works in conjunction with a whole-body PET/CT system based on the “virtual pinhole PET” concept is under development [38]. With this system, whole-body PET images with higher-resolution images of the area close to the insert can be obtained. Under the framework of the Crystal Clear Collaboration at CERN, the integrated system of Clear-PEM with ultrasound is under development [27]. A ring-shaped PET scanner for simultaneous breast PET/MR imaging has been developed in the Brookhaven National Laboratory [31]. This MRI-compatible PET scanner is designed to be placed within the breast radio- frequency coil of a 1.5 T MRI scanner.

Development and broad access to additional PET tracers are one of the biggest issues that may be associated to the future evolution of dedicated breast PET imag-ing. There are various PET tracers that have been developed and can be used to characterize breast cancers (Table 9.6). However, clinical application of PET tracers other than 18F-FDG is still limited. Breast cancer is one of the tumors in which

Table 9.6 Potential PET tracers for breast cancers

PET tracer Target18F-fluoroestradiol (FES) Estrogen receptor18F-fluorofuranyl norprogesterone (FFNP) Progesterone receptor89Zr-trastuzumab HER264Cu-DOTA-trastuzumab HER218F-ZHER2:342-Affibody HER264Cu-TP3805 VPAC118F-2-fluoropropionyl-labeled PEGylated dimeric RGD peptide (FPPRGD2)

αvβ3 Integrin

89Zr-bevacizumab VEGF-A18F-fluorothymidine (FLT) Proliferation11C-choline Membrane synthesis11C-methionine Protein synthesis18F-fluoromisonidazole Hypoxia18F-annexin V Apoptosis

HER2 human epidermal growth factor receptor 2, VEGF vascular endothelial growth factor


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several key molecules linked to the effectiveness of systemic hormone therapy and chemotherapy in individual tumors have been already identified [18, 19, 30]. Noninvasive methods that can visualize the expression status of these molecules may be useful to evaluate the spatial heterogeneity of the molecules within tumors and to monitor the dynamics of temporal expression changes. There is an urgent need for development and broad access to additional PET tracers.

9.5 Conclusions

Technological innovation has led to the development of dedicated breast PET sys-tems with high spatial resolutions. Although available clinical data are still limited, the ones available suggest that dedicated PET scanners could improve the detection of small breast cancers, as compared with the conventional PET scanners. Increased spatial resolution may also help to visualize detailed distribution of PET tracers within a tumor and could provide quantitative values with less bias caused by the partial volume effect. With the development of the multimodality techniques and new PET tracers for breast cancers, dedicated breast PET imaging may potentially allow further evolution in future breast cancer diagnosis and prognosis.

References

1. Abreu MC, Aguiar JD, Almeida FG, et al. Design and evaluation of the clear-PEM scanner for positron emission mammography. IEEE Trans Nucl Sci. 2006;53:71–7.

2. Avril N, Dose J, Jänicke F, et al. Metabolic characterization of breast tumors with positron emission tomography using F-18 fluorodeoxyglucose. J Clin Oncol. 1996;14:1848–57.

3. Avril N, Rosé CA, Schelling M, et al. Breast imaging with positron emission tomography and fluorine-18 fluorodeoxyglucose: use and limitations. J Clin Oncol. 2000;18:3495–502.

4. Berg WA, Weinberg IN, Narayanan D, et al. High-resolution fluorodeoxyglucose positron emission tomography with compression (“positron emission mammography”) is highly accu-rate in depicting primary breast cancer. Breast J. 2006;12:309–23.

5. Berg WA, Madsen KS, Schilling K, et al. Breast cancer: comparative effectiveness of positron emission mammography and MR imaging in presurgical planning for the ipsilateral breast. Radiology. 2011;258:59–72.

6. Berg WA, Madsen KS, Schilling K, et al. Comparative effectiveness of positron emission mammography and MRI in the contralateral breast of women with newly diagnosed breast cancer. AJR Am J Roentgenol. 2012;198:219–32.

7. Bettinardi V, Danna M, Savi A, et al. Performance evaluation of the new whole-body PET/CT scanner: discovery ST. Eur J Nucl Med Mol Imaging. 2004;31:867–81.

8. Bowen SL, Wu Y, Chaudhari AJ, et al. Initial characterization of a dedicated breast PET/CT scanner during human imaging. J Nucl Med. 2009;50:1401–8.

9. Broeders M, Moss S, Nyström L, et al. The impact of mammographic screening on breast cancer mortality in Europe: a review of observational studies. J Med Screen. 2012;19:14–25.

10. Bruening W, Uhl S, Fontanarosa J, et al. Noninvasive diagnostic tests for breast abnormalities: update of a 2006 review. Comparative Effectiveness Review No. 47. (Prepared by the ECRI Institute Evidence-based Practice Center under Contract No. 290- 02-0019.) AHRQ Publication No. 12-EHC014-EF. Rockville: Agency for Healthcare Research and Quality; 2012.


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11. Caldarella C, Treglia G, Giordano A. Diagnostic performance of dedicated positron emission mammography using fluorine-18-fluorodeoxyglucose in women with suspicious breast lesions: a meta-analysis. Clin Breast Cancer. 2014;14:241–8.

12. De Ponti E, Morzenti S, Guerra L, et al. Performance measurements for the PET/CT Discovery-600 using NEMA NU 2-2007 standards. Med Phys. 2011;38:968–74.

13. DeSantis C, Ma J, Bryan L, Jemal A. Breast cancer statistics, 2013. CA Cancer J Clin. 2014;64:52–62.

14. Eo JS, Chun IK, Paeng JC, et al. Imaging sensitivity of dedicated positron emission mammog-raphy in relation to tumor size. Breast. 2012;21:66–71.

15. Fowler AM. A molecular approach to breast imaging. J Nucl Med. 2014;55:177–80. 16. Glass SB, Shah ZA. Clinical utility of positron emission mammography. Proc (Baylor Univ

Med Cent). 2013;26:314–9. 17. Global Burden of Disease Pediatrics Collaboration, Fitzmaurice C, Dicker D, et al. The Global

Burden of Cancer 2013. JAMA Oncol. 2015;1:505–27. 18. Goldhirsch A, Wood WC, Coates AS, et al. Strategies for subtypes – dealing with the diversity

of breast cancer: highlights of the St. Gallen International Expert Consensus on the Primary Therapy of Early Breast Cancer 2011. Ann Oncol. 2011;22:1736–47.

19. Goldhirsch A, Winer EP, Coates AS, et al. Personalizing the treatment of women with early breast cancer: highlights of the St Gallen International Expert Consensus on the Primary Therapy of Early Breast Cancer 2013. Ann Oncol. 2013;24:2206–23.

20. Hendrick RE. Radiation doses and cancer risks from breast imaging studies. Radiology. 2010;257:246–53.

21. Iima M, Nakamoto Y, Kanao S, et al. Clinical performance of 2 dedicated PET scanners for breast imaging: initial evaluation. J Nucl Med. 2012;53:1534–42.

22. Kalinyak JE, Berg WA, Schilling K, et al. Breast cancer detection using high-resolution breast PET compared to whole-body PET or PET/CT. Eur J Nucl Med Mol Imaging. 2014;41:260–75.

23. Levine EA, Freimanis RI, Perrier ND, et al. Positron emission mammography: initial Clinical Results. Ann Surg Oncol. 2003;10:86–91.

24. Luo W, Anashkin E, Matthews CG. Performance evaluation of a PEM scanner using the NEMA NU 4—2008 small animal PET standards. IEEE Trans Nucl Sci. 2010;57:94–103.

25. MacDonald L, Edwards J, Lewellen T, et al. Clinical imaging characteristics of the positron emission mammography camera: PEM Flex Solo II. J Nucl Med. 2009;50:1666–75.

26. Miyake KK, Matsumoto K, Inoue M, et al. Performance evaluation of a new dedicated Breast PET Scanner using NEMA NU4-2008 Standards. J Nucl Med. 2014;55:1198–203.

27. Moadel RM. Breast cancer imaging devices. Semin Nucl Med. 2011;41:229–41. 28. Murthy K, Aznar M, Thompson CJ, et al. Results of preliminary clinical trials of the positron

emission mammography system PEM-I: a dedicated breast imaging system producing glucose metabolic images using FDG. J Nucl Med. 2000;41:1851–8.

29. O’Connor MK, Li H, Rhodes DJ, et al. Comparison of radiation exposure and associated radiation-induced cancer risks from mammography and molecular imaging of the breast. Med Phys. 2010;37:6187–98.

30. Perou CM, Sørlie T, Eisen MB, et al. Molecular portraits of human breast tumours. Nature. 2000;406:747–52.

31. Ravindranath B, Junnarkar SS, Purschke ML, et al. Results from prototype II of the BNL simultaneous PET-MRI dedicated breast scanner. IEEE Nucl Sci Symp Conf Rec. 2009;2009:3315–7.

32. Raylman RR, Majewski S, Smith MF, et al. The positron emission mammography/tomography breast imaging and biopsy system (PEM/PET): design, construction and phantom-based mea-surements. Phys Med Biol. 2008;53:637–53.

33. Rosen EL, Turkington TG, Soo MS, et al. Detection of primary breast carcinoma with a dedi-cated, large-field-of-view FDG PET mammography device: initial experience. Radiology. 2005;234:527–34.


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34. Schilling K, Narayanan D, Kalinyak JE, et al. Positron emission mammography in breast can-cer presurgical planning: comparisons with magnetic resonance imaging. Eur J Nucl Med Mol Imaging. 2011;38:23–36.

35. Shapiro S. Periodic screening for breast cancer: the HIP Randomized Controlled Trial. Health Insurance Plan J Natl Cancer Inst Monogr. 1997;22:27–30.

36. Springer A, Mawlawi OR. Evaluation of the quantitative accuracy of a commercially available positron emission mammography scanner. Med Phys. 2011;38:2132–9.

37. Tabár L, Vitak B, Chen HH, et al. The Swedish two-county trial twenty years later. updated mortality results and new insights from long-term follow-up. Radiol Clin N Am. 2000;38:625–51.

38. Tai Y-C, Wu H, Pal D, O’Sullivan JA. Virtual-Pinhole PET. J Nucl Med. 2008;49:471–9.



Chapter 10Recent Developments with Large-Bore PET/CT

Tomohiro Kaneta and Tomio Inoue

Abstract Large-bore positron emission tomography and computed tomography (PET/CT) combo scanners have recently been developed and released for use. The large bore not only provides a spacious and comfortable environment for the patient, but it also allows the operator to observe and interact with the patient more easily during the examination, thus maximizing safety and minimizing patient anxiety. However, the increased size of the bore may affect image quality and the sensitivity of the PET scanner. To compensate for the diminished image quality, it is necessary to develop different technologies, including a time-of-flight (TOF) function. In this chapter, we introduce recent developments in the field of large-bore PET/CT, using the Celesteion as an example.

Keywords Large bore • PET/CT • Time of flight • TOF

10.1 Introduction

Ever since the emergence of the combination of positron emission tomography (PET) and computed tomography (CT) into one system, the size of the common bore (patient port) of the PET/CT scanner has been approximately 700 mm (Table 10.1). This size is adequate for most patients; however, when the patient is over-weight or devices such as holding fixtures for radiotherapy and artificial ventilators are present, a larger bore is needed. Large-bore PET/CT scanners have recently been developed, including the Celesteion by Toshiba Medical Systems (Tochigi, Japan), which has bores of 900 mm for the CT section and 880 mm for the PET section. In addition, the scanner has a wide-scan field of view (FOV) of up to 700 mm. This large bore size not only provides the patient with a spacious and

T. Kaneta (*) Department of Radiology, Yokohama City University, Yokohama, Japane-mail: [email protected]

T. Inoue School of Medicine, Yokohama City University, Yokohama, Japan


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comfortable environment, but it also allows the operator to observe and interact with the patient more easily during the examination, thus maximizing safety and minimizing patient anxiety (Fig. 10.1). However, the larger bore size may affect image quality, especially the sensitivity of the PET scanner. Thus, Toshiba incorporated a time-of- flight (TOF) feature into the Celesteion. In this chapter, we discuss recent developments in the field of large-bore PET/CT, using the Celesteion as an example.

10.2 The Celesteion

The Celesteion PET/CT scanner combines a high-speed helical 16-slice CT scanner with a newly designed PET scanner, which harbors a lutetium–yttrium oxyorthosili-cate (LYSO) scintillator block detector. The CT scanner has 16 rows of tube detec-tors with 994 detector cells per row, and each revolution of a detector can cover up to 32 mm. The scintillator block detector system of the PET scanner contains LYSO crystals arranged in 48 rings. The crystals have a transaxial length of 4 mm and an axial length of 4 mm, which are arranged in detection units (blocks) consisting of 48 × 16 crystals coupled to a photomultiplier. The 48-ring system can obtain 95 PET

Table 10.1 Size of PET/CT scanners. The Celesteion has the largest patient port size

Model name Company

Patient port size [mm]

PET detector ring

FDA cleared

PET CT Diameter [mm] [year]

ECAT Accel CTI (Siemens) 562 N/A 824 2000Discovery LS GE Healthcare 600 700 927 2001Biograph Duo Siemens Healthcare 700 700 824 2002Discovery ST GE Healthcare 700 700 886 2003Aquiduo Toshiba Medical

Systems700 720 830 2005a

GEMINI TF Big Bore Philips Healthcare 850 850 900 2009Biograph TruePoint 16 Siemens Healthcare 700 700 842 2009Ingenuity TF Philips Healthcare 700 700 900 2011TruFlight Select Philips Healthcare 700 700 2011Discovery PET/CT 600 GE Healthcare 700 700 810Discovery PET/CT 610 GE Healthcare 700 700 886 2012Discovery PET/CT 690 GE Healthcare 700 700 886Discovery PET/CT 710 GE Healthcare 700 700 886 2012Biograph mCT Siemens Healthcare 780 780 842 2012Celesteion Toshiba Medical

Systems880 900 907 2014

Data were obtained from Imaging Technology News Online and Journal of Nuclear MedicineaJapanese FDA only

T. Kaneta and T. Inoue

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10 Recent Developments with Large-Bore PET/CT

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images (48 direct and 47 cross plane), separated by 2 mm and covering an axial FOV of 19.6 cm. The PET scanner can acquire data in three-dimensional (3D) con-figurations. The energy window of the system is set to 425–650 keV, and the coin-cidence time window is set to 1.6–4.2 ns, depending on the size of the FOV. The temporal resolution of the TOF is <450 ps.

10.2.1 Spatial Resolution

The National Electrical Manufacturers Association (NEMA) NU 2 2012 protocol [2] recommends using an F-18 fluorodeoxyglucose (FDG) point source (size <1×1 ×1 mm) inside a 75-mm-long glass capillary tube with an inner diameter of 0.9–1.0 mm and a 0.4-mm-thick wall. The total activity should be low enough to keep dead-time losses and randomness below 5% of the total events. Data were acquired at three transaxial locations (x,y), i.e., (1,0, 10,0, 20,0) cm, and at two axial positions (z) within the PET FOV, i.e., center FOV. At least 2×106 coincidence events were acquired at each position. TOF information was not used in this measurement. No attenuation and scatter correction and no post-smoothing filter were applied. The transverse and axial resolutions at the different positions of the point source are sum-marized in Table 10.2, which lists the full width at half-maximum (FWHM) values at 1, 10, and 20 cm. These data are provided by Toshiba Medical Systems.

10.2.2 Sensitivity

A 70-cm-long polyethylene tube (inner diameter = 1 mm; outer diameter = 3 mm) was filled with 4.5 MBq of FDG (at the start of acquisition) and placed inside of five concentric aluminum sleeves all with the same length and known diameters [2]. Data were acquired for each aluminum sleeve for 300 s (11.9 × 106 prompts in the

Table 10.2 Spatial resolution measured for the PET component of the Celesteion system (NEMA NU2-2012)

Spatial resolutionDistance (cm) FWHM(mm)

Transverse radial 1 4.61Transverse tangential

1 4.71

Axial 1 4.02Transverse radial 10 4.63Transverse tangential

10 4.85

Axial 10 4.52Transverse radial 20 4.73Transverse tangential

20 5.38

Axial 20 5.42

All data was provided by Toshiba Medical Systems


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first acquisition). The measurements were taken at the center of the FOV and at a 10-cm radial offset. Online random subtraction was applied using a delayed coinci-dence window. The corrected true coincidence count rate was recorded as a function of sleeve thickness and extrapolated to a zero-thickness sleeve. The system sensitiv-ity was then computed as the ratio of the true count rate with no absorption to the initial activity.

The Celesteion system has sensitivities of 3.90 and 3.71 kcps/MBq for the 0- and 10-cm off-center positions, respectively (from Toshiba Medical Systems data release). These sensitivities are lower than those of standard PET/CT scanners, probably because the detectors are farther apart in the large-bore scanner. To com-pensate for this potential disadvantage of the Celesteion system, certain steps were taken, as outlined below.

10.2.2.1 Increase the Axial FOV

The Celesteion was given more detector rings than previous scanners to increase the axial FOV. This increases the number of counts per bed position, while the larger bore decreases the number of bed positions required for the same axial coverage.

10.2.2.2 Time-of-Flight (TOF) Technique

While a conventional PET system can only localize the target to a straight line between the two detectors, a TOF PET system can determine more accurately the location of the target, by measuring the difference in the arrival times of two gamma rays. The Celesteion system has a temporal resolution of <450 ps, which enables the diameter of the positron source to be limited to <6.7 cm.

10.2.2.3 Scintillator

The advantages of the LYSO crystals contained in the multiple-ring scintillator block detector system of the Celesteion are the high light output and density, quick decay time, excellent energy resolution, and low cost. Table 10.3 compares the

Table 10.3 Properties comparison for LYSO, GSO, and BGO

Properties LYSO GSO BGO

Density (g/m3) 7.3 6.7 7.13Attenuation (cm−1) 0.87 0.67 0.96Decay Constant (ns) 50 50 300Light yield (%) NaI (Tl) 75 35 25Photofraction (%) 30 25 40Energy resolution (511 kev, %) 20 12 16


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properties of LYSO crystals with those of two other widely used crystals, bismuth germinate oxide (BGO) and gadolinium oxyorthosilicate (GSO). While the thick-ness of the scintillator affects temporal and spatial resolution, a thinner crystal has better temporal resolution but lower sensitivity due to the increased number of pen-etrating gamma rays. On the other hand, a thicker scintillator has decreased spatial resolution due to the increased variability in the location of scintillation. In general, non-TOF PET systems do not need good temporal resolution but have a thick scin-tillator to increase the sensitivity. Toshiba Medical Systems has not released the exact thickness of the crystal used in the Celesteion, but it appears to be thin.

10.2.2.4 Arrangement of Photomultiplier Tubes (PMTs)

Column-shaped PMTs are used to cover square-shaped crystals so it is impossible to completely cover the entire area of the crystals. The Aquiduo, an earlier PET/CT system from Toshiba Medical Systems, had PMTs of only one size that covered 69.1% of the area of the crystals. The Celesteion has PMTs of two sizes that cover 79.4% of the area of the scintillator array (Fig. 10.2). The increased coverage increases the sensitivity of the system.

10.2.2.5 Design of Electronic Circuits

The electronic circuits detect the coincidence events, record time information, con-figure the energy window, discriminate the pulse shape, and so forth. The Celesteion required fewer PMTs to detect a gamma ray, enabling an accurate count of coinci-dence events in a high count rate setting.

Fig. 10.2 Array consisting of two types of PMTs with different apertures so that they cover 79.4% of the area of the scintillator array


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10.2.3 Image Quality

A NEMA image quality phantom (International Electrotechnical Commission (IEC) body phantom) containing six spheres with internal diameters of 10, 13, 17, 22, 28, and 37 mm was used to evaluate the image quality of the Celesteion. A cylin-drical insert (diameter = 5 cm), containing a low-density material with an average density of 0.3 g/ml, was positioned in the center of the phantom to simulate lung tissue and provide a nonuniform attenuation distribution. The phantom was filled with a solution of water and 18F-FDG with a background activity concentration of ~5.3 kBq/ml. The four smallest spheres were filled to a target-to-background ratio (TBR) of 8:1 (for the first set of scans) and 4:1 (for the second set of scans). The remaining two largest spheres were filled with nonradioactive water. The phantom was positioned so that all the spheres were aligned within the same transaxial image plane in the center of the FOV.

For a PET scan, the matrix size was 144, the pixel size was 4 mm, and the slice thickness was 4 mm. CT scanning was performed using a 120-kV tube voltage. The FOV was set to 550 mm.

PET images were reconstructed using TOF list-mode ordered subset expectation maximization (TOF-LM-OSEM) with a 450-ps TOF temporal resolution kernel [3]. The TOF-LM-OSEM method is a TOF-OSEM algorithm that uses area-simulating volume (ASV) and calculates the geometric probabilities in the system matrix of 3D PET systems. The PET images were Gaussian filtered. For a 2-min PET scan, we varied the number of iterations and the parameters for the Gaussian filter to obtain the most suitable reconstruction conditions, which were then used in reconstructing PET images with other scan times.

Figure 10.3 shows the images of the IEC body phantom obtained with and with-out the TOF function. The TOF image from a 2-min scan had better spatial resolu-tion than the non-TOF image from a 5-min scan. The smallest sphere (diameter = 10 mm) had a higher maximum standardized uptake value (SUVmax) and looked clearer on the TOF image.

Figure 10.4 shows the clinical images obtained with and without the TOF func-tion. In our institution, patients received an intravenous injection of 3.5–5.0 MBq/kg of 18F-FDG after at least 6 h of fasting, followed by an uptake phase of 60 min. The image in Fig. 10.4a was obtained without the TOF function from the scan, at a rate of 4 min/bed position. The images in Fig. 10.4b–d were obtained with the TOF function from the scans at rates of 1, 1.5, and 2 min/bed position. In our opinion, the image obtained with the TOF function at a scan rate of 1.5 or 2 min/bed position (Fig. 10.4c and d) corresponds to the image without the TOF function from the scan at a 4 min/bed position rate (Fig. 10.4a).


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Fig. 10.4 Clinical FDG PET images obtained with and without the TOF function. (a) Image obtained without the TOF function from the scan with 4 min/bed position. (b–d) Images obtained with the TOF function from the scans with 1, 1.5, and 2 min/bed position. The images in (c) and (d) appear to be similar to the image in (a)

Fig. 10.3 Images of the IEC body phantom obtained with and without the TOF function. (a) The image of the 2-min scan obtained with the TOF function has better spatial resolution than that of (b) the 5-min scan without TOF. The smallest sphere (diameter = 10 mm) had the higher SUVmax and looked clearer on the TOF image


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Fig. 10.5 The frame format of STONES. After preprocessing the input image with overall smoothing, the image undergoes iterative comparisons and smoothing of adjacent voxels. After the iterative process, the final image and the input image undergo post-processing. They are weighted and blended to produce an output image

Fig. 10.6 Sample FDG PET images obtained with and without STONES. Compared to the non- post- filtered image, the image obtained with a Gaussian filter (b) is smoother than the image with no post-processing filtering (a) but has low contrast. (c) The STONES image has low noise and high contrast


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10.2.4 Statistical Transfer with Optimizing Noise and Edge Sensing (STONES)

Toshiba Medical Systems developed a new anisotropic diffusion filter, Statistical Transfer with Optimizing Noise and Edge Sensing (STONES), to reduce the statis-tical background in PET images while preserving the contrast and edges of the structures. This should improve the quality of images and make the acquisition time shorter. Figure 10.5 presents the frame format of STONES. The loop in the center indicates that the adjacent voxels are iteratively compared and smoothed, which decreases noise without loss of contrast. Before the selective smoothing in the itera-tive loop, a preprocessing filter performs overall smoothing. After selective smooth-ing, the weighted input image and the weighted output image of the loop are blended in post-processing procedure. Toshiba Medical Systems is in collaboration with Yokohama City University (Yokohama, Japan) to optimize the STONES for use in the Celesteion. Figure 10.6 presents sample FDG PET images obtained with and without STONES. Compared to the non-post-filter image (Fig. 10.6a), the image with the Gaussian filter (Fig. 10.6b) is smooth but with low contrast. The image taken with STONES (Fig. 10.6c) has low noise and high contrast.

10.3 Conclusion

We introduced a recently developed large-bore PET/CT system. The large bore pro-vides a spacious and comfortable environment for the patient, but it may affect image quality, especially via the sensitivity of the PET scanner. To compensate for these deficiencies, it was necessary to develop different technologies, including a TOF function.

References

1. National Electrical Manufacturers Association. NEMA standards publication NU 2–2007: performance measurement of positron emission tomographs. National Electrical Manufacturers Association: Rosslyn; 2007.

2. National Electrical Manufacturers Association. Performance measurements of positron emission tomographs. NEMA Standards Publication NU 2–2012. National Electical Manufacturers Association: Rosslyn, USA; 2012.

3. Ye H, Niu X, Wang W. Improved list-mode reconstruction with an areasimulating-volume projector in 3D PET: IEEE Medical Imaging Conference Record 2012.



Chapter 11Copper-Labeled Radiopharmaceuticals in Oncology

Hiroaki Kurihara

Abstract Among copper (Cu) isotopes, radioactive 60Cu, 61Cu, 62Cu, 64Cu, and 67Cu have potential to be used in biomedical research. 60Cu, 61Cu, and 62Cu can be used in positron emission tomography (PET) diagnostic imaging. Similarly, 67Cu can be used in gamma imaging, and it is potentially suitable for diagnostic PET imaging and radiotherapy. These five radioactive Cu isotopes can be produced in a cyclotron. 62Cu and 64Cu are the most frequently used Cu radioisotopes because of their avail-ability. Recently, interest in Cu radiopharmaceuticals has been increasing because they have been conjugated to antibody proteins, peptides, nanoparticles, and small molecular compounds for preclinical and clinical studies. This chapter provides an overview of the preparation, chemical, and clinical applications of Cu-labeled radiopharmaceuticals.

11.1 Introduction

Copper (Cu) is a transition metal with the atomic number 29, which is an important and widespread element in the world. It has been known since ancient times, and it is used as currency in many countries. Cu has 2 stable isotopes, 63Cu and 65Cu, and 27 unstable radioisotopes with decades of research. Most radioactive Cu isotopes have half-lives of less than 1 min; however, 60Cu, 61Cu, 62Cu, 64Cu, and 67Cu have relatively longer half-lives ranging from 10 min to 62 h, and they are suitable for potential use in biomedical research. For example, 64Cu was used to evaluate Cu metabolism and to assess patients with Wilson disease in 1982 [1]. In the last few decades, the development of imaging modalities and radiolabeling techniques has permitted to overcome the limits in anatomical imaging, allowing for a shift toward molecular imaging. By not only generating high-resolution images of the human body but also assessing the distribution of specific molecules noninvasively, molec-ular imaging helps physicians to diagnose illness and choose appropriate treatment regimens [2]. This new approach with radioactive Cu offers further opportunities for

H. Kurihara (*) National Cancer Center, Tokyo, Japane-mail: [email protected]

https://en.wikipedia.org/wiki/Copper

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clinical research, diagnostic imaging, and targeted radiotherapy. Copper isotopes with atomic masses below 63 tend to undergo β+ decay and can be used in PET diagnostic imaging. In contrast, isotopes with atomic masses above 65 tend to undergo β− decay and can be used in gamma imaging and radiotherapy. Moreover, radioactive 64Cu undergoes both β+ and β− decay, so 64Cu may be suitable for use in diagnostic PET imaging, gamma imaging, and radiotherapy (Table 11.1).

Availability is an important consideration for the widespread use of any radioiso-tope in medicine. In this regard, the half-lives of 60Cu and 61Cu seem to be too short for commercial delivery of routine clinical isotopes in nuclear medicine practice. 62Cu also has a short half-life, 10 min, but it can be easily produced on-site using a 62Zinc (Zn)/62Cu generator and readily available for clinical studies. For example, 62Cu(II)-diacetyl-bis(N4-methylthiosemicarbazone) (62Cu-ATSM), an imaging agent targeting hypoxia, can be prepared for clinical research using a simple procedure with generator-produced 62Cu [3]. 64Cu has also been available for clinical research because it can be produced in a baby cyclotron located in a hospital. It has a rela-tively longer half-life (12.7 h), which makes it possible to prepare 64Cu-labeled radiopharmaceuticals. Despite the fact that 64Cu has a low positron branching ratio (17.6%), some researchers cite 64Cu-labeled radiopharmaceuticals, which can be used as PET probes that provide better intrinsic image resolution with low β+ maxi-mal energy [4, 5]. 67Cu undergoes β− decay and is potentially useful for targeted radiotherapy, but due to limited availability, few researchers have reported its use thus far [6, 7].

In keeping with this notion, Cu-labeled radiopharmaceuticals may have clinical impact in oncology. In this chapter, an overview of the production and properties of

Table 11.1 Decay characteristic and production property of copper radiopharmaceutical

IsotopeNuclear production reaction Half-life β+ MeV (%) β− MeV (%) EC (%) γ MeV (%)

60Cu 60Ni(p,n)60Cu 23.4 min 3.92 (6%) – 7.4% 0.85 (15%)59Co(3He,2n)60Cu 3.00 (18%) 1.33 (80%)

2.00 (69%) 1.76 (52%)2.13 (6%)

61Cu 61Ni(p,n)61Cu 3.3 h 1.22 (60%) – 40% 0.284 (12%)59Co(3He,n)60Cu 0.38 (3%)

0.511 (120%)

62Cu 62Ni(p,n)62Cu 9.7 min 2.91 (98%) – 2% 0.511 (194%)60Cu(p,2n)62Zn:

62Zn/62Cu64Cu 64Ni(p,n)64Cu 12.7 h 0.655

(19.3%)0.573 (39.6%)

41% 1.35 (0.6%)0.511 (38.6%)

67Cu 68Zn(p,2p)67Cu 62.0 h – 0.577 (20%) – 0.184 (40%)67Zn(n,p)67Cu 0.484 (35%) 0.092 (23%)

0.395 (45%)

H. Kurihara

https://en.wikipedia.org/wiki/Positron_emission

https://en.wikipedia.org/wiki/Beta_decay

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radioactive Cu, the chemistry of Cu radiopharmaceuticals, and clinical applications for Cu radiopharmaceuticals is provided.

11.2 Production and Properties of Radioactive Cu

With advances in the medical sciences, Cu has gained a lot of attention, especially in molecular imaging. Several radioactive Cu isotopes can be obtained by bombard-ing nickel (Ni), cobalt (Co), or Zn with proton (p), neutron (n), or helium (He) in a cyclotron or reactor, but proper selection of the radionuclide to generate radiophar-maceuticals is critical. It depends upon several factors: half-life, radionuclide energy, cost, and availability. The half-life of the radionuclide should allow for suf-ficient uptake and distribution to yield high image quality. The energy of the radio-nuclide emission should be appropriate for proper detection by the equipment [8]. This section provides an overview of the production and properties of five radioac-tive Cu isotopes because they have the most potential for molecular imaging appli-cations (60Cu, 61Cu, 62Cu, and 64Cu) and in vivo targeted radiotherapy (64Cu and 67Cu) [8].

11.2.1 60Cu

60Cu can be produced in a medical cyclotron at a relatively low cost, using proton or deuteron bombardment of enriched 60Ni targets [9, 10]. Other methods of produc-tion, such as the 59Co(3He,2n)60Cu reaction, have been reported [11, 12]. 60Cu is a proton-rich nuclide with a half-life of 23.7 min to its stable 60Ni isotopes, through a combination of positron decay and electron capture processes. Thus, 60Cu is a potential candidate tracer for PET imaging. However, its relatively high-energy positron (maximum energy, 3.92 MeV) and numerous gamma concurrent emissions in cascades (most notably 0.85 MeV, 15%; 1.33 MeV, 80%; 1.76 MeV, 52%; and 2.13 MeV, 6%) affect image quality. These additional γ-photons can downscatter into the acceptance energy window of the scanner and supply incorrect positional information, which results in higher background activity and poorer image quality and accuracy.

11.2.2 61Cu

61Cu can be produced by proton bombardment of 61Ni targets or 3He bombardment of 59Co targets, using a medical cyclotron. The half-life of 61Cu (3.3 h) is longer than that of 60Cu and 62Cu, which makes 61Cu a better choice for prolonged imaging with slower kinetics. However, this isotope is not currently popular because it requires

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highly enriched Ni targets or high-energy 3H beams, which limits the accessibility of 61Cu [8, 12–15]. More economical production methods should be developed before clinical use [16, 17].

11.2.3 62Cu

62Cu is a proton-rich nuclide that decays with a half-life of 9.7 min to its stable 62Ni isotope, through a combination of positron decay (98%) and electron capture pro-cesses (2%). It can be produced using a 62Zn/62Cu generator system. Current 62Zn/62Cu generators achieve high elution efficiency, approximately 96%, using a small volume (approximately 3 ml) of eluate, with very low breakthrough of 62Zn [18]. One disadvantage of 62Zn/62Cu generators is that they can be used for only 1–2 days because of the relatively short half-life of parent 62Zn (half-life, 9.2 h). However, this method has been proven to be highly useful as a source of 62Cu for the synthesis of 62Cu-labeled compounds [18, 19]. 62Cu can also be produced with a medical cyclotron, using proton- or deuteron-induced reactions on enriched 62Ni targets. However, this cyclotron method is not popular in nuclear medicine because its short half-life limits the final yield of 62Cu-labeled radiopharmaceuticals pre-pared from the irradiated target [9]. 62Cu from a 62Zn/62Cu generator is currently the most intensively studied Cu radioisotope after 64Cu [8, 18, 19].

11.2.4 64Cu

64Cu decays via three processes: positron decay, electron capture, and beta decay. This property allows 64Cu isotopes to be used for both PET imaging and radiother-apy. With a half-life of 12.7 h, 64Cu is ideally suited for PET studies that require a longer-lived nuclide. Distribution of 64Cu radiopharmaceuticals from the production site to other facilities is possible, and PET imaging can be conducted up to 48 h after tracer administration. Moreover, because 64Cu has a maximum positron energy of 0.655 MeV, similar to that of 18F, the resulting PET images are of high quality and are the best obtainable with any of the positron-emitting radioactive Cu isotope. Since 39.6% of 64Cu decay occurs by β− emission, there is a possibility of therapeu-tic applications with this nuclide.

64Cu can be produced using either a cyclotron or a reactor. Currently, the most common 64Cu production method utilizes the 64Ni(p,n)64Cu reaction, which can be conducted even with a medical cyclotron [20, 21]. The target for producing 64Cu is enriched 64Ni as follows: a 64Ni target is electroplated on a gold disk at a thickness of 50–100 μm and electroplating is performed at 2.5 V, at currents ranging from 5 to 15 mA, which is completed in 12–24 h. The 64Ni target is then bombarded with a 50 ± 3 μA proton current and after bombardment, 64Cu can be purified from the 64Ni target and other contaminants using an anion exchange column. Target 64Ni can be recovered and reused [22] with this method, and sufficient quantities of purified

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64Cu for diagnostic imaging and therapeutic application can be obtained. The radio-nuclide purity of the 64Cu generated using this process is over 99%. The specific activity of 64Cu may range widely because ubiquitous cold Cu may contaminate the procedure. Producing high levels of 64Cu-specific activity requires careful manage-ment to maintain a metal-free environment for 64Cu preparation.

Another method of 64Cu production using the 64Zn(n,p)64Cu reaction in a nuclear reactor has been described elsewhere [23, 24]. However, this reaction has waste concerns and undesirable byproducts such as 65Zn (half-life, 245 d), which limits the practicality of this production method.

11.2.5 67Cu

67Cu is the longest living Cu radioisotope. It undergoes β− decay with a half-life of 62 h. The growing interest in targeted radiotherapy has increased the demand for this β− emitting isotope. 67Cu can be produced via several reactions with Zn, but it is one of the most difficult radioactive Cu isotopes to produce. A 68Zn(p,2p)67Cu nuclear reaction, the most popular reaction for 67Cu production, requires a high- energy proton beam (20–70 MeV) [25]. The other method to produce 67Cu is via the 67Zn(n,p)67Cu reaction. Although this appears to be simple, it requires a fast neutron flux reactor, and the use of this reactor reaction for extended productions and medi-cal applications is associated with waste concerns and undesirable side reactions. Recently, a novel larger-scale production method for 67Cu using a high-energy pro-ton accelerator has been described, which may allow for commercial delivery of 67Cu to become available in the near future [26].

11.3 Chemistry of Cu Radiopharmaceuticals

Radioactive Cu offers the ability to exploit unique aspects of Cu radiopharmaceutical chemistry. In aqueous solution, Cu is mainly restricted to oxidation states I and II. Cu(I) generally exists only in aqueous solution as a strong complex, since the free ion is disproportionate to Cu2+ and Cu0 metal. A Cu3+ ion may exist under certain conditions, but it is not stable in the biochemical environment because of its strong oxidizing properties. On the other hand, Cu(II) prefers a coordination number of 4. In general, it can be chelated by ligands arranged in a square planar configuration.

So far, many ligands that can chelate Cu(II), such as thiosemicarbazones and macrocyclic bifunctional chelators, have been previously reported. Cu(II)-thiosemicarbazones were first designed in 1964 as anticancer agents and evaluated as radiopharmaceuticals since 1987 [27]. Thiosemicarbazones labeled with radioac-tive Cu have been developed to image blood flow and hypoxia. Other bifunctional chelators are known to be suitable for conjugating radioactive Cu to antibody pro-teins or peptides. In order for these radioactive Cu isotopes chelated with biomole-cules to be effective, high thermodynamic and in vivo stability of the radiolabeled


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biomolecules are required. For example, complexes of Cu(II) and bifunctional che-lators such as ethylenediaminetetraacetic acid (EDTA) and diethylenetriaminepen-taacetic acid (DTPA) were not stable enough in vivo because they rapidly dissociate in human serum and Cu(II) binds to human serum albumin [28]. Other macrocyclic chelators have varying degrees of biological stability. For instance, Cu-labeled 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid (DOTA) is more stable in serum than Cu-labeled 1,4,8,11-tetraazacyclotetradecane-N,N′,N′′,N′′′-tetraacetic acid (TETA) compounds [29]. In addition, the charge of the Cu(II) complex has an important effect on biodistribution. Negatively charged complexes are excreted from the body more rapidly than positively charged ones [30] Fig. 11.1.

O

O

O

O

OO

O

O

O

OH

OH

OH

HOHO

HO

HO

HO

EDTA DTPA

DOTA

TETAcyclam

cyclen

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CB-cyclen

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NH

NH

NHNH

HN

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NH HN

HN

HOOC

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COOH

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NN

NN

HOOC

HOOC

COOH

COOH

NN

NN

HN

N

N

N

NH

HNN

N

N

NOH

Fig. 11.1 Chemical structures of several bifunctional chelators. EDTA ethylenediaminetetraacetic acid, DTPA diethylene triamine pentaacetic acid, cyclam 1,4,8,11-tetraazacyclotetradecane, cyclen 1,4,7,10-tetraazacyclododecane, TETA 1,4,8,11-tetraazacyclotetradecane-N,N′,N′′,N′′′-tetraacetic acid, DOTA 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid, CB-cyclam cross-bridged cyclam, CB-cyclen cross-bridged cyclen

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11.4 Clinical Applications of Radioactive Cu

11.4.1 Cu-Labeled Agents to Assess Blood Flow

The design of radiopharmaceutical agents for diagnostic imaging is dependent on the desired characteristics of the molecular targets. Due to the high rates of morbid-ity and mortality associated with cardiovascular and cerebrovascular diseases, myo-cardial and cerebral perfusion tests are important clinical applications. To date, there are single photon emission computed tomography (SPECT) agents that rou-tinely use current clinical nuclear medicine procedures. However, PET has the potential to improve diagnostic accuracy in these applications, which has led to an interest in developing radioactive Cu-labeled blood flow agents for myocardial and cerebral perfusion imaging. Blood flow agents have been evaluated as ligands that form lipophilic uncharged complexes with Cu. These Cu complexes have been designed to be sufficiently stable to clear the blood and diffuse into tissues of inter-est, such as the brain, heart, kidneys, and even tumors. Cu complexes that release the Cu radionuclide after reaching the tissue of interest are advantageous, because the radioactive Cu can then be trapped in the tissue of interest without the need to washout.

Cu-bis(thiosemicarbazone) (Cu-BTS) complexes have been extensively evalu-ated for their use as Cu radiopharmaceuticals in PET perfusion imaging. Pyruvaldehyde bis(N4-methylthiosemicarbazone)Cu (Cu-PTSM) is the first clini-cally used and widely studied Cu-BTS complex. Human PET studies with 62Cu-PTSM have demonstrated that the tracer provides high-quality PET images of the heart, in which regional myocardial perfusion is accurately delineated [31]. In addition, 62Cu-PTSM is a sufficiently sensitive tracer for regional cerebral blood flow. It has been used to detect regional cerebral perfusion impairment in patients who suffered a stroke [32]. However, 62Cu-PTSM binds to human serum albumin reversibly, which impairs its ability to quantify myocardial perfusion under hyperemic conditions. To overcome this problem, next-generation Cu-BTS com-plexes that may be suitable for quantification of myocardial perfusion with PET have been developed.

11.4.2 Cu-Labeled Hypoxia Imaging Agents

Hypoxic tissue in the brain, heart, and tumors is considered an important imaging target. In oncology, it frequently suggested that a hypoxia imaging technique may help select cancer patients, who would benefit from treatments that overcome, cir-cumvent, or take advantage of the hypoxic environment. This is because tumor hypoxia is an important biological characteristic that leads to radioresistance during cancer treatment. Imaging could also be used to demonstrate the degree to which reoxygenation of tumors occurs during radiotherapy. Many challenges in hypoxia


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imaging with magnetic resonance, optical, and nuclear imaging have been reported [33]. For instance, magnetic resonance imaging (MRI) methods are attractive because they avoid the use of radioactivity and MRI equipment is widely available, but hypoxia-reporter molecules for MRA require relatively large quantities of the reporter. PET has high sensitivity and spatial resolution, which has the advantage of visualizing molecular events in living human tissue, which makes it a leading tool for imaging hypoxia. Hypoxia-reporter molecules used in PET can be divided into two groups based on chemical structure: nitroimidazole compounds and non- imidazole imaging agents [34]. The nitroimidazole compounds include 18F-Fluoromisonidazole (18F-FMISO), 18F-fluoroerythronitroimidazole (18F-FETNIM), and 18F-fluoroazomycin-arabinofuranoside (18F-FAZA). Non- imidazole imaging agents contain a metal complex of radioactive Cu with diacetyl- bis(N4-methylthiosemicarbazone) (ATSM). The following section will focus on agents with radioactive Cu-ATSM, which are the leading contenders for human applications [35].

11.4.2.1 Radioactive Cu-ATSM

The proposed mechanisms of Cu(II)-ATSM accumulation in hypoxic cells include Cu(II)-ATSM, an uncharged lipophilic, highly membrane-permeable molecule, which can undergo reduction by cellular-reducing equivalents and be converted to [Cu(I)-ATSM]−. Then tetrahedral Cu(I) can be easily released from ATSM to form strong Cu(I) complexes or be converted to Cu0 (Cu metal) or Cu2+, which subse-quently forms complexes with proteins and is trapped in hypoxic cells (Fig. 11.2).

Radioactive Cu-ATSM can be readily synthesized with a one-step simple reac-tion using radioactive Cu and the substrate ATSM [36]. In brief, 10 μl of ATSM dissolved in dimethyl sulfoxide solution (1 mg/ml) is added to hydrochloride- buffered radioactive Cu chloride for 2 min. Next, radioactive Cu-ATSM is eluted with ethanol using a C18 Sep-Pak cartridge. After approximately 80% of the ethanol has been evaporated with blown argon gas, the remaining radioactive Cu-ATSM is diluted in saline for injection. Quality analyses showed high radiochemical purity (>95%) and a high yield (>95%) of the final radioactive product [37, 38].

For PET imaging, there are four different positron-emitting Cu isotopes: 60Cu (half-life, 0.39 h; β+, 93%), 61Cu (half-life, 3.3 h; β+, 60%), 62Cu (half-life, 0.16 h; β+, 98%), and 64Cu (half-life, 12.7 h; β+, 19.3%). 60Cu, 61Cu, and 64Cu can be pro-duced in small cyclotrons, and 62Cu can be obtained via a 62Zn/62Cu generator system.

Since Fujibayashi et al. have reported the use of 62Cu-ATSM as a hypoxia imag-ing agent in 1997 [38], numerous studies have demonstrated that 60Cu-ATSM uptake can predict tumor behavior and response to therapy in patients with non-small cell lung cancer [39], cervical cancer [40], colorectal carcinoma [41], and malignant glioma [3]. Chao et al. demonstrated the feasibility of using 60Cu-ATSM imaging to identify hypoxic tumor sub-volume, through co-registration of computed tomogra-phy and 60Cu-ATSM PET images in order to plan a patient’s course of radiotherapy

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and perform intensity-modulated radiation therapy [42]. When Lewis et al. com-pared image quality and tumor uptake of 60Cu-ATSM and 64Cu-ATSM in ten patients with cervical carcinoma, they concluded that both were safe radiopharmaceuticals and that image quality with 64Cu-ATSM is better than that with 60Cu-ATSM due to reduced background [43]. In a study by Dehdashti et al., 60Cu-ATSM uptake (tumor- to- muscle ratio threshold of 3.5) in 38 patients with cervical cancer was inversely related to progression-free survival and cause-specific survival. Similarly, 62Cu-ATSM PET can visualize lung adenocarcinoma and enlarged mediastinal lymph node [44]. Moreover, 62Cu-ATSM PET can help identify highly malignant gliomas, while 62Cu-ATSM T/B ratio may predict hypoxia-inducible factor-1α (HIF-1α) expression. This suggests that 62Cu-ATSM is a suitable biomarker for pre-dicting high-grade malignancy and tissue hypoxia in patients with glioma [45]. An example of a clinical 62Cu-ATSM PET/CT study is shown in Fig. 11.3.

11.4.2.2 Clinical Role of Hypoxia Imaging in Oncology

Tumor hypoxia is commonly present in tumor tissues, where it poses a significant challenge to the curability of human tumors, which can lead to resistance to therapy and enhanced tumor progression. Imaging tumor hypoxia has two major clinical uses: (a) select patients, who may show a better response to curative treatments designed to overcome the limitations of hypoxia and (b) reveal by serial imaging that the treatment strategy reduce the extent of regional hypoxia. Hypoxia imaging may also allow for better identification of a subpopulation of cancer patients, who would benefit from novel targeted anti-hypoxia therapies.

Easy penetration

Cu(II)-ATSM

Cu(II)-ATSM

Hypoxia

retention

Reducedmitochondria

Cu-(I)

e-

Fig. 11.2 Proposed retention mechanism of Cu-ATSM in a hypoxic cell. Cu(II)-ATSM is a neutral lipophilic molecule, which can easily penetrate cell membrane. In normoxic cells, the neutral lipophilic Cu(II)-ATSM can be easily washed out (solid arrow). In hypoxic calls, Cu(II)-ATSM can be converted with electron (e−) supplied from abnormally reduced mitochondria to Cu-(I), which is entrapped in the cells


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11.4.3 Cu-Labeled Antibodies

Many novel molecular targets to treat cancer have been discovered, which has led to the development of molecularly targeted drugs. These drugs are cornerstones of precision medicine that use information from a patient genetic to diagnose, treat, and prevent the disease. Therapeutic monoclonal antibodies have played a major role in molecularly targeted therapy, and there are demands for the use of radiola-beled antibodies in molecular imaging, which can identify the presence of specific targets throughout the body, in a noninvasive approach.

For example, trastuzumab, a humanized monoclonal antibody against human epidermal growth factor receptor 2 (HER2), is a widely used therapeutic antibody. The application of trastuzumab-mediated technology is a well-established strategy to target HER2-positive breast cancer [46–50]. HER2 is overexpressed in 25–30% of patients with breast cancer and is directly involved in tumor cell survival, prolif-eration, maturation, metastasis, and angiogenesis [51, 52]. HER2 expression is rou-tinely determined using immunohistochemistry (IHC) or fluorescence in situ

Fig. 11.3 A clinical case of the 62Cu-ATSM PET/CT application (Courtesy of Hirofumi Fujii, M.D., National Cancer Center Hospital East). A 60-year-old man suffering from lung adenocarci-noma was examined with 62Cu-ATSM PET/CT study. (a, b) CT images show lung tumor in right upper lobe and mediastinal lymphadenopathy. (c, d) 62Cu-ATSM PET/CT fusion image and PET image revealed high accumulation in the lung tumor and the enlarged lymph node

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hybridization [53]. However, technical problems can arise when lesions cannot be easily accessed for core needle biopsy [54]. In addition, HER2 expression can vary during the course of the disease [55], even between tumors in the same patient [56]. To overcome these problems, a novel noninvasive technique such as HER2 PET imaging is desirable for conclusive assessment of HER2 expression.

So far, molecular PET imaging for HER2 has been studied intensively with trastuzumab labeled with 124I, 86Y, 76Br, 89Zr, or 64Cu [4, 57]. For PET imaging, 64Cu can be a useful radioactive Cu nuclear agent because of its relatively longer half- life. The following section will focus on anti-HER2 molecular imaging with 64Cu-labeled trastuzumab in humans.

Briefly, the preparation of 64Cu-labeled trastuzumab occurs as follows: after puri-fication of the trastuzumab IgG (Herceptin®) by ultrafiltration with phosphate- buffered saline (PBS), the PBS-filtered trastuzumab is added to 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid (DOTA) mono-N- hydroxysuccinimide ester and dissolved in water. After incubation at 40 °C for 3 h, crude DOTA-trastuzumab is purified with PBS by using a PD-10 column. The PBS buffer, including DOTA-trastuzumab, is exchanged for a sodium acetate buffer (100 mM, pH 6.5) by filtration. 64Cu-DOTA-trastuzumab is then prepared by adding previously purified 64Cu (see “Production of Radioactive Cu”) to the DOTA- trastuzumab acetate solution, followed by incubation for 1 h at 40 °C. Lastly, the reaction mixture is sterilized by filtration through a 0.22-μm filter [4]. In sum, radio-labeling achieves a specific activity of approximately 350 GBq/μmol and a 98% radiochemical purity. Using the 64Cu-trastuzumab PET imaging technique, primary tumor lesions larger than 2 cm in diameter and metastatic brain lesions larger than 1 cm in diameter can be visualized [4, 58]. Typical 64Cu-DOTA-trastuzumab PET images in patients with HER2-positive breast cancer are shown in Fig. 11.4. Remarkably, HER2 specificity of 64Cu-DOTA-trastuzumab has been confirmed in human samples, prepared from surgically removed tumor specimens by autoradiography, IHC, and liquid chromatography-tandem mass spectrometry (LC-MS/MS) [4, 58]. As an example, Fig. 11.5 shows an autoradiograph of a frozen section prepared from a tumor specimen, in which high signal accumulation can be observed in the area where HER2-positive cells were seen by IHC.

Another HER2 PET imaging with a 89Zr-trastuzumab probe can successfully visualize HER2-positive tumors in humans. Due to the relatively longer half-life of Zr-89, 89Zn-trastuzumab provides clearer images. However, it involves higher radia-tion exposure [59]. In contrast, the shorter half-life of 64Cu involves lower radiation exposure but provides images with non-specific activity in the blood [4, 60]. Although improvements are still needed, these imaging techniques can be used to serially monitor HER2 tumor status, during HER2 targeted therapy and also serve to evaluate patients with tumors that are not easily accessible with core needle biopsies.

Recently, several novel HER2 inhibitors have been developed and approved, such as lapatinib, pertuzumab, and trastuzumab emtansine (T-DM1). Monitoring changes in HER2 expression at tumor sites may help physicians determine which HER2 inhibitor should be used during different phases of treatment or if non-HER2


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inhibitors should be used instead. Preclinical results with HER2 imaging are prom-ising, but clinical data are still limited. Clinical HER2 PET imaging findings may support the further development and exploration for the potential of this technique.

11.4.4 Cu-Labeled Peptides

Most current Cu radiopharmaceutical research focuses on 64Cu-labeled peptides for targeted cancer therapy or imaging. They consist of a targeting peptide such as bombesin or an octreotide analogue, a linker, and a bifunctional chelator such as TETA or DOTA. The peptide binds to a specific receptor expressed by cancer cells, while the 64Cu moiety allows for localization of the tumor by PET. For example,

Fig. 11.4 64Cu-DOTA-trastuzumab PET images of metastatic brain tumors in patients with HER2- positive primary breast cancer. The white arrows show the metastatic brain tumors. Upper panels, 64Cu-DOTA-trastuzumab PET images; lower panels, Gd-DTPA-enhanced T1-weighted MRI images. White arrows indicate metastatic brain lesions detectable by both MRI and 64Cu-DOTA- trastuzumab PET, and red arrow indicates a lesion detectable by MRI but not by 64Cu-DOTA- trastuzumab PET

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octreotide, a peptide of eight amino acids, is a somatostatin analogue that has affin-ity for somatostatin receptors. Targeting of somatostatin receptors in tumors has been a goal of cancer treatment and diagnosis. In the first-in-human study, 64Cu-DOTATATE PET was found to be useful to image the somatostatin receptor [61]. Compared with 111In-DTPA-octreotride SPECT, 64Cu-DOTATATE PET pro-vided superior image quality, detected more true-positive lesions, and was associ-ated with a lower radiation burden.

11.5 Summary

Radioactive Cu isotopes are playing a larger role in both PET imaging and targeted radiotherapy. The versatility of Cu isotopes gives them a strong position in the development of new pharmaceuticals. With a 62Zn/62Cu generator system, 62Cu can be obtained easily in hospital settings for use in PET imaging. 64Cu can be used in hospitals, since its longer half-life allows for production at regional or national cyclotron facilities and distributed to local nuclear medicine departments. In addi-tion, high doses of 64Cu-labeled radiopharmaceuticals can be used for targeted radiotherapy. Radioactive Cu could be suitable for theranostics by replacing 62Cu or 64Cu with high-dose 64Cu or 67Cu in Cu-labeled radiopharmaceuticals.

Fig. 11.5 Histological distribution of 64Cu-radioactivity in HER2-positive tumors. Left column, HE staining; middle column, IHC; right column, autoradiography. Loupe images (upper panels) show identical distribution of radioactivity and location of HER2-positive tumor cells for HE stain, IHC, and autoradiography samples. Magnified images (lower panels, ×200) confirmed the radio-activity and HER2-positive status of tumor cells


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31. Green MA, Mathias CJ, Welch MJ, McGuire AH, Perry D, Fernandez-Rubio F, Perlmutter JS, Raichle ME, Bergmann SR. Copper-62-labeled pyruvaldehyde bis(N4- methylthiosemicarbazonato)copper(II): synthesis and evaluation as a positron emission tomography tracer for cerebral and myocardial perfusion. J Nucl Med. 1990;31:1989–96.

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39. Dehdashti F, Mintun MA, Lewis JS, Bradley J, Govindan R, Laforest R, Welch MJ, Siegel BA. In vivo assessment of tumor hypoxia in lung cancer with 60Cu-ATSM. Eur J Nucl Med Mol Imaging. 2003;30:844–50.

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Chapter 12PET Radiopharmaceuticals in Oncology Beyond FDG

Shao-Li Song, Pan-Li Li, and Gang Huang

Abstract Several imaging modalities to diagnose cancer, which include computed tomography (CT), magnetic resonance imaging (MRI), ultrasound, optical imaging, and gamma scintigraphy, have been previously used. For instance, CT and MRI provide considerable anatomic information about the location and the extent of tumors, but do not adequately differentiate residual or recurrent tumors from edema, radiation necrosis, or gliosis. On the other ultrasound images provide limited infor-mation about local and regional morphology with blood flow. Similarly optical imaging showed promising results, but did not demonstrate the ability to detect deep tissue penetration. Notably, radionuclide imaging modalities are diagnostic cross- sectional imaging techniques that map the location and concentration of radionuclide- labeled compounds. Moreover, molecular imaging agents are making it possible to “see” the molecular makeup of the tumor and its metabolic activity beyond tumor location, size, shape, and viability. Other technological limitations include nuclear images, which can provide a very accurate definition of metabolically active areas, but miss anatomic features. As a result, new imaging modalities have combined nuclear images with CT scans for treatment planning. The hybrid scanners combine anatomic and functional images taken during a single procedure, without having to reposition the patient between scans. In this chapter, multiple ligands beyond clini-cal standard 18F-fluorodexoyglucose are reviewed.

S.-L. Song • P.-L. Li Department of Nuclear Medicine, RenJi Hospital, School of Medicine, Shanghai JiaoTong University, Shanghai, Chinae-mail: [email protected]; [email protected]

G. Huang (*)Renji Hospital / Department of Nuclear Medicine, Shanghai Jiao Tong University, Shanghai, China e-mail: [email protected]




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12.1 Molecular Imaging in Oncology

CT and MRI assess the effectiveness of cancer therapy (e.g., volumetric and mor-phological changes), whereas the endpoints rely almost exclusively on the analysis of biopsies by molecular and histopathological methods. While these methods pro-vide microscopic data of the general heterogeneous process, nuclear imaging mea-sures blood flow and metabolic patterns of the target organs. Therefore, a predictive nuclear biomarker is needed to assess clinical endpoints adequately and allow pre-cise measurements of tumor targets on a whole-body image, upon administration of a functional radiolabeled agent. These mechanism-based agents provide image- guided therapy, which may help to decide whether it is better for the patient to dis-continue an ineffective treatment at an earlier phase. Molecular imaging agents play a major role in drug discovery and development because of their ability to quantify drug properties in vivo. For example, positron emission tomography (PET) agents show high sensitivity and specific activities, since they are made through a nuclear transformation and use carrier-free forms of isotopes. Thus, PET agents do not pro-duce detectable pharmacologic effects but provide important information concern-ing the characterization of various diseases such as the central nervous system (epilepsy, psychosis, dementia, Alzheimer’s disease) [1–4] and cardiovascular sys-tem diseases (myocardial viability) [5]. At the same time it also helped determcine cancer staging, cancer restaging, and treatment planning for malignant diseases [6]. In addition, molecular imaging helps to control and monitor dosage for increased safety and effectiveness. The trends for PET agent’s development in oncology are to assist in determining optimal therapeutic dose, to delineate differential diagnosis between inflammation/infection and recurrence, to determine sensitive or resistant to treatment responses, and thus to select the patients who may be good candidates for therapy. Below are the molecular targets for pathway- activated imaging systems.

12.2 Glucose Transporter Target

2-Deoxy-2-[18F]fluoro-D-glucose (18F-FDG)-PET has been used to diagnose [7–9] and stage tumors [10–20], myocardial infarction [21], and neurologic diseases [22, 23]. 18F-FDG was developed in 1976 for the specific purpose of mapping brain glu-cose metabolism in living humans. After the first synthesis of 18F-FDG via an elec-trophonic fluorination with 18F gas, small-volume enriched water targets were developed, which made it possible to produce large quantities of [18F] fluoride ion via high yield 18O(p,n)18F reaction. This was followed by a major milestone, result-ing in the development of a nucleophilic fluorination method that produced 18F- FDG in very high yields. These advances and the remarkable properties of 18F-FDG have largely overcome the limitations of the 110-min half-life of 18F. Although 18F- FDG has been successfully used to image tumors with high glycolytic activity [24]

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in the past two decades, it has several limitations that give rise to false-positive/false-negative diagnosis and poor predictive value of tumor response to chemora-diation therapy [25]. For instance, FDG has poor contrast in brain tumor due to the high uptake of glucose in normal brain tissue [26, 27], and exhibits poor differentia-tion between tumors and inflammatory tissue due to the high uptake of FDG by granulocytes and macrophages [28] .

To improve the diagnosis, prognosis, planning, monitoring, and prediction of cancer treatment results, radiolabeled N-acetylglucosamine of the hexosamine bio-synthetic pathway (HBP) has been used as an alternative to PET FDG [29–32]. Glucose and glutamine are substrates in the synthesis of GlcNAc via the HBP and key nutrients for cancer cells. Glutamine is a precursor amino acid for the synthesis of glucosamine, a prominent initiator in the HBP [33]. Briefly, fructose-6-phosphate from the glycolytic pathway combines with glutamine in the presence of the initiat-ing enzyme glutamine–fructose-6-phosphate amidotransferase (GFAT) to synthe-size glucosamine-6-phosphate. Subsequent enzymatic steps lead to the production of uridine diphosphate N-acetylglucosamine (GlcNAc), a substrate for O-linked glycosylation that is regulated by the terminating enzyme O-linked GlcNAc (O-GlcNAc) transferase (OGT). OGT is the enzyme responsible for the addition of a single GlcNAc residue to the hydroxyl groups of serine and/or threonine residues of target proteins. The HBP terminating in O-GlcNAc cycling (O-GlcNAcylation) has been implicated in cellular signaling cascades and regulation of transcription factors involved in cancer biology [34–37]. In fact, O-GlcNAc glycosylation on the serine or threonine residues of nuclear and cytosolic proteins is common and occurs frequently in all multicellular eukaryotes [38]. O-GlcNAc is part of posttransla-tional modification reactions and appears to modify many nucleocytoplasmic pro-teins, such as the transcription factors NF-κB and NFATc1. NF-κB and NFATc1 are both downstream targets of O-GlcNAc modification during T- and B-cell activation [39]. However, these transcription factors are constitutively activated in cancers and play an important role in growth and survival [40]. The chelator ethylenedicysteine (EC) was conjugated to glucosamine to mimic a GlcNAc analog called ethylenedicysteine- N-glucosamine (ECG). Radiolabeled ECG has been reported as a promising imaging agent in various tumor models in rodents [29, 31], as well as in humans [30].

12.3 Amino Acid Transporter Target

Radiolabeled amino acids are an alternative in characterizing tumors because of their much higher accumulation in tumors than in normal tissues and their rapid blood clearance. In keeping with the higher uptake, the expression of amino acid transporters is upregulated in cancer cells [41, 42], which indirectly measure cell proliferation [43], and assessment of their activities provides the potential of dif-ferential diagnosis and early prediction of treatment responses. Although 11C-methionine is useful for metabolic imaging of tumors by PET [44], it has too

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many metabolic pathways, which makes it difficult to obtain a rate constant [20, 21]. Furthermore, it is difficult to image tumors that have slow uptake because of its short half-life. Among all radiolabeled amino acids, aromatic amino acids are more suitable due to the easier chemistry alteration and characterization. However, aromatic amino acids tend to decarboxylate [45], which reduce its ability to get into cells via amino acid transporters. To circumvent these potential limitations, placement of a methyl group at the alpha position could retard metabolism, and hence, greater effort could be then directed toward the 18F- and 123/124I-labeled amino acids such as 2-[18F] fluoromethyl-L-phenylalanine [46], position 2- and 3-L-[18F] flour- α- methyl tyrosine (FAMT) [47–49], and 123I-labeled L-3-[123I] iodo-α-methyltyrosine (123I-AMT) [50, 51]. L-type amino acid transporter system (LAT), especially LAT1 subtype, is the only system that can transport large neutral amino acids with aromatic rings such as tyrosine, phenylalanine, and tryptophan [52, 53]. LAT family is known to form heterodimers, which contain a chaperone-like heavy chain 4F2hc, and a 12-time-folding transmembrane light chain, which is unique to each subtype [54]. Previous clinical studies showed that uptake of FAMT in the tumor was closely correlated with LAT1 expression and cellular pro-liferation [41, 48]. Its analog, 123I- AMT, has also been used for SPECT studies on brain and pancreatic tumors [55–57]. High accumulation of 123I-AMT in tumors was reported upon PET examination with 124I-AMT, which was carried out in patients with brain tumors [57]. Though AMT shows very promising clinical results, the cost of using such isotope is prohibitive. Therefore, it is desirable to develop a radiotracer with easier chemistry and more affordable isotope that can be used clinically in most major medical facilities. Few promising chelator–amino acid conjugates in animal studies have been reported [58–60].

12.4 Estrogen Receptor Target

The presence of sex hormone receptors in both primary and secondary breast tumors is an important indicator for both prognosis and choice of therapy of the disease. Currently, receptors are determined by in vitro analysis of biopsy specimens and the use of antiestrogen antibodies. Tamoxifen is the therapy of choice for estrogen receptor-positive (ER+) tumors. The detection and measurement of ER+ tumors by the use of a radiolabeled ligand should provide a useful tool for the diagnosis of primary and secondary tumors. This approach may assist in selecting and following the most favorable therapy, as well as predicting its outcome. To this end a number of variations of substituted estradiol [61, 62] and tamoxifen [63] have been pre-pared. These compounds have been relatively successful in detecting ER-rich tissue in vivo, such as breast cancer, ovarian cancer, endometriosis, uterine carcinoma, and meningioma. These compounds may predict the response to anticancer therapy agents.

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12.5 Tumor Hypoxia Target

The success to the noninvasive detection of the hypoxic fraction of tumors by nuclear molecular imaging is that it allows physicians to select patients for addi-tional or alternative treatment regimens that may circumvent or overcome the omi-nous impact of tumor hypoxia and improve disease control [64]. Misonidazole (MISO) is a hypoxic cell sensitizer, and labeling it with different radioisotopes could be useful for differentiating a hypoxic but metabolically active tumor from a well-oxygenated active tumor by PET or planar scintigraphy. The nitro group of nitroimidazole (NIM) is enzymatically reduced by ribonucleoside reductase within viable hypoxic cells [65, 66]. 18F-FMISO and clinically user-friendly chelator-based hypoxia tracers have been used to assess the hypoxic component in brain ischemia, myocardial infarction, and various tumors [51, 67]. Moreover, the assessment of tumor hypoxia prior to radiation therapy would provide a rational means to select patients for treatment with radiosensitizing or bioreductive drugs (e.g., mitomycin C). It is also possible to select proper modalities of radiotherapy (proton, neutron vs. photon) by correlating tumor uptake results with tumor response.

12.6 Lipid Metabolism Target

An elevated level of phosphatidylcholine has been found in tumors, which is the most abundant phospholipid in the cell membranes of all eukaryotic cells and pro-vides a potential target for tumor imaging. This elevation is thought to be the result of increased uptake of choline, a precursor of the biosynthesis of phosphatidylcho-line. Malignant tumors show a high proliferation and increased metabolism of cell membrane components that will lead to an increased uptake of choline [54]. Thus, radiolabeled choline can be used as a PET marker for imaging cell proliferation in prostate cancer, brain tumors, and many other types of tumors that lack the urinary radioactivity seen with 18F-FDG [68, 69].

12.7 Tumor Cell Proliferation Target

Noninvasive imaging assessment of tumor cell proliferation could be helpful in the evaluation of tumor growth potential and the degree of malignancy and treatment response in the early assessment, prior to changes in tumor size. Radiolabeled nucleoside/nucleotide analogs should provide proliferative imaging information of primary and secondary tumors. They may also assist in selecting and following the most favorable choice of nucleoside/nucleotide therapy and in following its out-come. 18F-Fluorothymidine (FLT) was developed in an attempt to improve the understanding of the biologic behavior of malignant tumors, which should lead to


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better prognostic evaluation, treatment follow-up, and patient management. However, 18F-FLT shows a lower accumulation in tumors than 18F-FDG, since it only accumulates in the cells that are in the S phase of the cell cycle. This demon-strates a low sensitivity for nodal staging [70], which indicates that there is still room to develop better markers of cell proliferation biomarkers.

12.8 Gene Expression Target

Radiolabeled pyrimidine and purine probes to image herpes simplex virus type 1 thymidine kinase (HSV-1-tk) expression and other reporter genes by PET have been developed [71–73]. For example, pyrimidinenucleoside (e.g., FIAU, 2′-fluoro-2′-deoxy-5-iodo-1-β-D-ribofuranosyl-uracil [FIRU], 2′-fluoro-2′-5-methyl- 1-β-D-arabinofuranosyl-uracil [FMAU], 2′-fluoro-2′-deoxy-5-iodovinyl-1-β-D- ribofuranosyluracil [IVFRU]), acycloguanosine([9-[(2-hydroxy-1-(hydroxymethyl)ethoxy)methyl]-guanine (GCV) and 9-[4-hydroxy-3-(hydroxymethyl)butyl] guanine (PCV)]), and other 18F-labeled acycloguanosine analogs such as 8-fluoro-9-[(2-hydroxy-1-(hydroxymethyl)ethoxy)methyl]guanine (FGCV), 8-fluoro-9-[4-hydroxy- 3-(hydroxymethyl)guanine (FHBG), were developed as reporter substrates to image wild-type and mutant HSV-1-tk expression. Recently, imaging, pharmacokinetics, and dosimetry of 18F-FHBG were reported in healthy volunteers as a first step to image the HSV-1-tk reporter expression in clinical gene therapy trials. The difficulty with these probes is that HSV-1-tk enzyme expression depends on HSV-1-tk gene transduction with adenoviral vectors. The level of HSV-1-tk enzyme expression is likely to be different in diverse types of transuded cells and tissues, thus limiting the application. Understanding the proliferative activity of tumors could aid in the selec-tion of optimal therapy, by estimating patient prognosis and selecting the proper management.

In summary, in order to improve the diagnosis, prognosis, planning, and moni-toring of the cancer treatment, it is important to characterize tumor tissue exten-sively by the development and application of more tumor-specific pharmaceuticals. Radiolabeled ligands as well as radiolabeled antibodies have opened a new era in scintigraphy detection of tumors and have undergone extensive preclinical evaluation.

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Chapter 13Molecular Imaging in Oncology: Challenges and Opportunities in the Twenty-First Century

Ya-Ming Li, David Yang, and Xue-Na Li

Abstract Each improvement of radionuclide imaging modalities (positron emis-sion tomography, PET; single-photon emission computed tomography, SPECT) is dependent upon the development of tumor-specific radiopharmaceuticals with enhanced tumor specificity. Applications of molecular targets for cancer imaging, therapy, and prevention that use generator-produced isotopes are the major focus of ongoing research projects. 99mTc (technetium-99 m,), 68Ga (gallium-68), and 188Re (rhenium-188) are generator-produced isotopes, which are readily accessible and affordable. While 99mTc (t1/2 = 6 h, 140 keV) is used for SPECT, 68Ga (t1/2 = 68 min, 511 keV) labeling is for PET, and both are intended to image molecular targets to determine therapeutic dosage, monitor tumor response to treatment, and establish differential diagnoses. 188Re (t1/2 = 16.9 h, 2.1 MeV) is a therapeutic radionuclide, which can be used to target tumors and deliver lethal radiation resulting from high- energy β− emissions. The present chapter discusses challenges and opportunities in molecular imaging, involving drug discovery, validation, intellectual property, licensing, and regulatory compliance.

Keywords Molecular imaging • 99mTc • 68Ga • 188Re

Y.-M. Li (*) • X.-N. Li Department of Nuclear Medicine, The First Hospital of China Medical University, No.155 North Nanjing Street, Heping District, Shenyang, Liaoning 110001, Chinae-mail: [email protected]

D. Yang Vyripharm Biopharmaceuticals, University of Texas Health Science Center, Houston, TX, USA


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13.1 Molecular Imaging: New Medical Field

Although traditional imaging modalities such as computed tomography (CT) [1, 2], magnetic resonance imaging (MRI) [3, 4], and ultrasound (US) [5] provide tumor volumetric measures and morphological values, they lack diagnostic precision, spe-cific therapeutic guidance, and prognostic information. Conversely, molecular imaging overcomes the limitations inherent to conventional imaging, by measuring tumor attributes such as metabolism, proliferation, or surface receptor expression. To that end, molecular imaging modalities include optical, positron emission tomography (PET), and single-photon emission computed tomography (SPECT) [6–8]. Optical imaging uses specific fluorescent or near-infrared probes to target tumor, but is limited by tissue penetration of the signal, and it is not currently useful for deep tissue lesions. On the other hand, nuclear imaging modalities such as PET and SPECT are most commonly used for molecular imaging and use radiotracers to image, map, and measure target site activities (e.g. angiogenesis, metabolism, apop-tosis, and proliferation). Advances in PET and SPECT technology and the discovery of new molecular probes have increased the interest in developing new radiotracers. Notably, radiotracers have no pharmacologic effect due to the negligible amounts of the ligand, and thus, toxicity is not a concern. In addition to specific target assess-ment, whole-body images can be obtained to examine the pharmacokinetic pattern of radiotracers. Radiotracers provide an avenue for noninvasive assessment of treat-ment endpoints, which, in the past, could only be obtained by histopathological methods. Unlike conventional imaging methods, which are ill suited for early pre-diction of therapeutic responses, radiotracers can provide early information about the efficacy or failure of a given treatment, potentially within a few days, since biological features typically precede anatomical tumor changes. In addition to ther-apeutic monitoring, molecular imaging modalities can be incorporated into all phases of drug discovery and development to quantify drug properties in vivo. Presently, hybrid imaging modalities PET/CT [9, 10] and SPECT/CT [11] have been developed to allow for simultaneous detection of functional and anatomical tumor information. In this report, we describe the pros and cons of molecular imag-ing using nuclear imaging modalities.

13.2 Application of Molecular Imaging in Oncology

Molecular imaging research can be classified as either basic molecular characteriza-tion or translational imaging. In basic molecular characterization, the research focus is to reveal a mechanism and as a proof of concept. In translational imaging, the research focus leads to clinic trials and marketing applications. While the focus of molecular imaging research has traditionally been on identification of tumor- specific markers, translational imaging research has focused mostly on the applica-tion of these markers for evaluation of patient response to radiation therapy, chemotherapy, or chemo-/radiotherapy. Recent studies have focused on the

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identification of certain aspects of therapeutic resistance, in an effort to select patients for molecular-based therapies such as anti-angiogenic [12, 13], thrombo-lytic [14], radiation, and chemo-sensitizing agents [15], as a means to improve radi-ation or chemotherapeutic response. Hence, there is a need for the development of clinically useful radiotracers, which provide (1) proof of concept for target assess-ment, (2) probes to monitor and predict the efficacy of therapy (differentiate tumor grading, inflammation, and infection versus tumor recurrence), (3) rationale to expand to targeted radiotheranostic applications, and (4) a platform to accelerate targeted drug development. Moreover, a successful molecular imaging program must lead to opportunities for grant funding through academic and industrial col-laborations, which result in the generation of intellectual property.

13.3 Opportunities in Molecular Imaging

In 2006, there were an estimated 1.5 million new cases of cancer and 56,000 cancer- related deaths, and as a result of these statistics, the market for diagnostic agents and targeted therapies began to expand. Drug discovery and development are now accel-erating due to the rapid synthesis of potential agents and the development of high- throughput in vitro tests. Drug development can be classified into five phases: (1) identification of a potential drug, (2) proof of concept, (3) clinical assessment of safety and effectiveness of the agent, (4) application for product approval, and (5) regulatory institutional approval. In keeping with these phases, molecular imaging can accelerate drug approval through compelling data that assess metabolism, bio-distribution, volume distribution (Vd), half-life, and pharmacokinetics, which are required by regulatory statute 21CFR 505b1, b2, and j [http://www.fda.gov/cder/index.html under regulatory guidance column]. Table 13.1 summarizes the applica-tion process for drug approval of molecular imaging (NDA, new drug application; ANDA, abbreviated new drug application).

Table 13.1 Opportunity of molecular imaging in drug development

CFR Agent Formulation Preclinic Clinic

505 (b)1 New New Metabolism BioavailabilityBiodistribution MetabolismBioavailability t1/2

PF, Vd, t1/2 VdPk

505 (b)2ANDA

Known New Same as 505 (b)1 Same as 505 (b)1Plus API identity and bioequivalence

Plus API identity and bioequivalence

505 (j)ANDA

Known Known Same as 505 (b)1 Same as 505 (b)1New indication New indication (e.g.,

pamidronate in osteoporosis and anticancer use)

13 Molecular Imaging in Oncology: Challenges and Opportunities…

http://www.fda.gov/cder/index.html

http://www.fda.gov/cder/index.html

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13.4 Development of Molecular Imaging Technology

Successful technologies in molecular imaging in oncology should be able to pro-vide information about tumor molecular attributes (angiogenesis, tumor growth, apoptosis) in order to assess tumor staging, grading, and differentiation and to mon-itor the response to therapy. This implies that in order for this technology to achieve clinical impact, the new tracers must be suitable for easy scale-up and cost-effective.

In certain applications, the development of novel radiotracers utilizes coordina-tion chemistry, in which a chelator is required to trap metal isotopes. This type of chemistry is simple and can be performed with high yield. The radionuclides may be obtained from generators, which provide a low-cost source of isotopes. The clinic-grade prodrug can be formulated as a kit product, which is attractive to phar-maceutical companies involved in the manufacturing and standardization of the agent for future multicenter clinical studies, in a manner that strengthens its applica-tion for the regulatory approval process. The potential benefit of a kit format in clinical oncology is to enable physicians to see, treat, and monitor the progression of the tumor, by defining its shape, size, viability, and differentiation. In this con-text, the kit technology provides a better radiation dosimetry for therapy and offers a simple “shake-and-shoot” chemistry for quick manufacturing and accessibility.

Among all SPECT radioisotopes, 99mTc has been the preferred label for radio-pharmaceuticals because of its favorable low energy (140 keV), inexpensive isotope cost, and easy chemistry. For instance, L,L-ethylenedicysteine (EC) is a successful example of N2S2 chelation [16], which has been extensively labeled with 99mTc with high radiochemical purity and stability of several hours [17–27]. In addition to assessing molecular targets, development of 99mTc agents may be useful for plan-ning specific internal radionuclide therapy with beta-emitting radionuclides such as 188Re. 188Re has attractive imaging characteristics and potential therapeutic use because of its high β− energy (2.1 MeV), short physical half-life (16.9 h), and 155 keV γ-ray emissions, which are useful for dosimetric and imaging applications [28]. The short half-life of 188Re allows for higher doses, compared with long-lived radio-nuclides. Furthermore, it reduces the problems of radioactive waste handling and storage. Similar to 99mTc, 188Re is usually available from in-house generator systems. As a matching pair for imaging and therapy, 99mTc- and 188Re-based agents represent an exciting strategy for personalized medicine and can drive the discovery and development of new molecular imaging agents.

PET imaging beyond 18F-fluorodeoxyglucose (FDG) has sought to use different isotopes, which are more accessible and cost-effective. One example is 68Ga, which possesses significant commercial potential because the isotope can be produced from a 68Ga generator (270-day half-life) on site and thus serve as a convenient alternative to cyclotron-based PET isotopes. 68Ga has a high positron emitting quan-tity (89% of its total decay), and therefore, the main consideration is its spatial reso-lution, which depends on the positron range (energy); the non-colinearity of annihilating photons, intrinsic properties, size, and geometry of the detector; and

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the selection of the reconstruction algorithm. The features of the detector design, physical properties, and the influence on the system’s spatial resolution have been extensively addressed by many authors, leading to a continuous optimization of the hardware. Although the maximum positron energy of 68Ga (max, 1.90 MeV; mean, 0.89 MeV) is higher than that of 18F (max, 0.63 MeV; mean, 0.25 MeV), a study using Monte Carlo analysis on spatial resolution revealed that under the assumption of 3 mm spatial resolution of PET detectors, the conventional full width at half maximum (FWHM) of 18F and 68Ga is indistinguishable in soft tissue (3.01 mm vs. 3.09 mm) [29–31]. As a result, 68Ga-based PET agents are viewed as a convenient, low-cost alternative to cyclotron-produced PET isotopes, which may be patterned after the well-established 99mTc scheme: (1) development of chelator-based radiotracers in a kit format, (2) cost-effective approach to functional imaging, and (3) ability to provide high-quality diagnostic imaging to all clinics equipped with appropriate cameras, irrespective of the presence of a cyclotron.

13.5 Intellectual Property in Molecular Imaging

The development of a new agent inherently leads to a new intellectual property and inevitably to patent and licensing issues. There are three categories of patents: the first one is a utility patent that covers animals, machines, articles of manufacture, composition of matter, and methods of making and doing things; the second one is a simple design patent that protects only the ornamental appearance of the device and may contain a single claim; the third type is a plant patent that includes any plant produced by budding or grafting. In order to be eligible for biotech-based util-ity patent, the patent must include the composition of matter (structure), manufac-turing process (technological methods), the application for new use, or an improvement of the process. The claims of a patent must be novel (unknown or not described in a literature), nonobvious to a person having ordinary skill in the rele-vant art, and useful. To confirm that the invention does not potentially infringe on other similar patents, it is necessary to review the relevant literature for “prior art” before filing the disclosure on the intellectual property.

A patent application from a single institute involves filing an invention disclosure report (IDR) or filing a provisional patent, followed by the filing of an international patent, also known as patent cooperation treaty (PCT). The text of the PCT applica-tion is usually filed 1 year after the provisional application, and it is based on the same text, although it can incorporate updates and revisions of data. Eighteen months after PCT is filed, national stage applications that claim priority to it can be filed, such as non-provisional applications in the USA and other jurisdictions like Canada, China, Europe, etc. There must be sufficient disclosure in the PCT applica-tion, so that someone with ordinary skills in the field of the invention can make or use the invention without an undue amount of experimentation. If there are any preferred or best ways of practicing the invention, that method (if known) must be disclosed in the application. Failure to do so may cause the issued patent to be


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invalidated. After the US patent is filed, either a continuation-in-part (CIP), which must be filed within a year of the original application, or a new patent may be filed if new data arise.

Intellectual property rights, such as the ownership of patents, control, and respon-sibility for patent litigation process, technology transfer, and the right to sublicense, are important to a research institutions and biotechnology companies. Sublicenses typically result in negotiations between the patent holder and licensee, regarding clinical trial and distribution rights, such as data management and control of clinical trials, control of initial manufacturing process, and marketing rights to the product by region and milestone payments. The licensing process is summarized by a term sheet, which covers the license fee, milestone payments (determined by research progress), and royalty streams.

13.6 Challenges in Molecular Imaging

Key milestones in the clinical development of imaging products are (1) chemistry, manufacturing, and control (CMC); (2) pharmacology and toxicology; and (3) phase I clinical studies.

13.7 Chemistry, Manufacturing, and Control (CMC)

The CMC section of a clinical protocol requires information about the chemical formulation and production of the drug. The Food and Drug Administration (FDA) permits radiopharmaceutical production in central commercial facilities under well- controlled conditions and distribution to local clinics, where they are administered. Similarly, radionuclide generator systems that can be produced in a well-controlled facility are embraced by current FDA procedures and have a long history of suc-cessful clinical applications. The regulatory compliance from FDA’s perspective includes good laboratory practice (GLP), current good manufacturing practice (cGMP), and good clinic practice (GCP). GLP covers pharmacology and toxicology (known as safety and effectiveness assays). cGMP covers CMC, batch record, facil-ity, and personnel. GCP covers protocol, data management, and informed consent. Overall, GLP, cGMP, and GCP include organized personnel, equipment, facilities, reproducibility, batch and lot records, accountability for manufacturing, reprocess versus rework, and quality assurance. General criteria of radiotracer development from bench to clinic are shown in Table 13.2.

In addition to the information in Table 13.2, the CMC section also includes a suitable reference standard for the active pharmaceutical ingredient (API), valida-tion of radiochemical purity, stability of the agent, and description of the site per-forming the cGMP processes. Two certificates of analysis (COA) are generally

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required for each new radiotracer. One is for precursor (prodrug) and another one is for final drug product. Two analytical tools, such as HPLC and radio-TLC, are com-monly used to determine the purity of the radiolabeled drug.

13.8 Pharmacology Validation and Toxicology

Preclinical pharmacology and toxicology studies must also be included in clinical protocols. These include studies on drug uptake and target validation, biodistribu-tion and dosimetry, image consistency (software for analysis), and toxicology stud-ies under GLP guidelines.

13.9 Clinical Studies

GCP covers the organization, personnel, facility, clinic protocol, informed consent source document, clinical research agreement, data management, and institutional review board (IRB) approval. The guidance outlined in the GCP presents several challenges for the partnership between academic research institutions and the bio-tech industry. These challenges involve issues of relationship maintenance such as “What are the basic ground rules of the business arrangement?”; intellectual prop-erty such as “Who owns any resulting inventions or discoveries?”; and academic freedom such as “Who controls the data?” and “What are the publication rights?”; and liability exposure.

Table 13.2 General criteria of radiotracer development from bench to clinic

Toxicology (safety)LD-50Tolerated doseAcute toxicity, chronic toxicityPharmacology (effectiveness)PharmacokineticsMetabolismDispositionChemistry, manufacturing, and controlsIngredient (active, inactive),Strength (mCi/ml)Specific activity, purity (HPLC, TLC for chemical, radiochemical osmolality, pH)Sterility, pyrogenicityGMP or GLP process


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13.10 Conclusions

Molecular imaging presents broad opportunities and challenges. Agent develop-ment using novel isotopes and ligands may lead to more personalized and efficient patient care and disease management, while also producing new intellectual prop-erty. This may open new collaborative efforts between multidisciplinary groups to address regulatory concerns and streamline the approval process.

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25. Zheng YY, Saluja S, Yap GP, et al. Gallium and Indium Complexes of Bis(amino thiol) (N(2)S(2)) Ligands. Inorg Chem. 1996;35:6656–66.

26. Anderson CJ, John CS, Li YJ, et al. N,N′-ethylene-di-L-cysteine (EC) complexes of Ga(III) and In(III): molecular modeling, thermodynamic stability and in vivo studies. Nucl Med Biol. 1995;22:165–73.

27. Sun Y, Anderson CJ, Pajeau TS, et al. Indium (III) and gallium (III) complexes of bis(aminoethanethiol) ligands with different denticities: stabilities, molecular modeling, and in vivo behavior. J Med Chem. 1996;39:458–70.

28. Azhdarinia A, Yang DJ, Yu DF, et al. Regional radiochemotherapy using in situ hydrogel. Pharm Res. 2005;22(5):776–83.

29. Sanchez-Crespo A, Andreo P, Larsson SA. Positron flight in human tissues and its influence on PET image spatial resolution. Eur J Nucl Med Mol Imaging. 2004;31:44–51.

30. Grove NB, Martin MJ. Log-f tables for beta decay. Nucl Data Tables. 1971;10:205–19. 31. Schechter NR, Yang DJ, Azhdarinia A, Chanda M. Technologies for translational imaging

using generators in oncology. Recent Patents Anti-cancer Drug Discovery. 2007;2:1–8.



Chapter 14Visualization of Apoptosis: Annexin V Imaging

Skye Hsin-Hsien Yeh, Fan-Lin Kong, and Ming-Hsien Lin

Annexin V, a human protein with nanomolar affinity for cell membrane-bound phosphatidylserine (PS), is the most widely used conjugate for the detection of apoptosis by using the imaging modalities such as SPECT, PET, MRI, and optical imaging. This chapter will initially focus on the most recent reports on annexin V-conjugated imaging agents in both animals and humans, followed by conclusions and the possible future directions of annexin V imaging.

14.1 Introduction

Apoptosis or programmed cell death plays a critical role in normal physiology and pathology of numerous disease states [1]. Therefore, the in vivo visualization of apoptosis would allow for both early detection of therapy efficiency and evaluation of disease progression. Several agents have been developed and investigated for apoptosis imaging by using different imaging modalities such as single-photon emission computed tomography (SPECT), positron emission tomography (PET), optical imaging (OI), and magnetic resonance imaging (MRI). Each imaging modal-ity has certain advantages as well as limitations. The choice of the right imaging modality or hybrid scanner depends on the parameter of interest under consider-ation (i.e., anatomical structure, functional metabolism, etc.).

S.H.-H. Yeh (*) Brain Research Center, National Yang-Ming University, Taipei, Taiwan

Aging and Health Research Center, National Yang-Ming University, Taipei, Taiwane-mail: [email protected]

F.-L. Kong Department of Cancer Systems Imaging, The University of Texas MD Anderson Cancer Center, Houston, TX, USA

M.-H. Lin Division of Nuclear Medicine, Taipei City Hospital Zhongxiao Branch, Taipei, Taiwan


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14.2 Annexin V-Phospholipid Complex

The externalization of the phosphatidylserine (PS) on the cell membrane has been identified as a major biochemical marker of apoptosis and could in principle be exploited for the detection of apoptosis [2]. Annexin V (36 kDa), which interacts strongly and specifically with phosphatidylserine residues, has been the most stud-ied imaging probe for apoptosis [3] (Table 14.1).

14.3 SPECT and PET Imaging

14.3.1 Single-Photon Emission Computed Tomography (SPECT) Imaging

Since the first apoptotic imaging reported in 1999 [4, 5], in vivo imaging of cell death with radiolabeled annexin V has been widely used in animal studies and clini-cal trials [6]. The unique advantages of radiotracers include their high sensitivity and the translational potential. Among various SPECT radionuclides, technetium- 99 m (99mTc) is the most prominent isotope for the nuclear imaging because of its ideal nuclear properties and easy availability at low cost [7, 8].

Based on the previous study [9, 10], Blankenberg et al. reported an improved 99mTc-annexin V radioprobe using the bifunctional agent hydrazino nicotinamide (HYNIC) [5]. 99mTc-HYNIC-annexin V showed the greatest uptake in the kidneys, liver, and urinary bladder; however, it was devoid of any bowel excretion, resulting in excellent signal to background ratio in the abdominal region. With the modified procedure of preparation, 99mTc-HYNIC-annexin V could be synthesized efficiently with high yield. By far, 99mTc-HYNIC-annexin V has been extensively investigated in animal models [11–15]. Multiple clinical trials have confirmed the clinical utility of 99mTc-HYNIC-annexin V in determining the efficacy of chemotherapy in the patients of non-small cell lung cancer for the detection of apoptotic regions [16, 17].

In 2000, Tait et al. reported 99mTc-HYNIC-cys-annexin V117, which revealed site-specific labeling; however, it showed similar apoptosis avidity when compared to the previous version of 99mTc-HYNIC-annexin V [18]. Similar radiolabeled annexin V probes such as 99mTc(CO)3-HIS-cys-AnxV [19, 20] and 99mTc-His10- annexin V [21] demonstrated improved sensitivity for detecting dead or dying cells.

Yang et al. reported the use of 99mTc-EC-annexin using ethylenedicysteine (EC) as a chelator to assess the level of apoptosis of tumor cell [26]. The preclinical data of breast cancer patients showed the total effective dose equivalent for 99mTc-EC- annexin V of 6.80–7.89 mSv could be reasonable and allow it for clinical use and it could be a predictor for evaluating the treatment-related apoptosis after induction of chemotherapy [27].

Recently, 99mTc-C3(BHam)2-annexin V was developed using a bis(hydroxamide) derivative [C3(BHam)2] as a bifunctional chelating agent [28]. In vivo evaluation of

S.H.-H. Yeh et al.

235

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14 Visualization of Apoptosis: Annexin V Imaging

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S.H.-H. Yeh et al.

237

99mTc-C3(BHam)2-annexin V showed decreased uptake and retention in nonspecific tissues and much lower kidney accumulation of radioactivity when compared to 99mTc-HYNIC-annexin V. Their findings also indicated that 99mTc-C3(BHam)2- annexin V could be a potential candidate as a predictor for response to chemotherapy.

Additional to 99mTc, 67Ga [22, 23] and 111In [24, 25] were also used to label annexin V or its mutants for site-specific detection of apoptosis.

14.3.2 Positron Emission Tomography (PET) Imaging

The major advantages of PET imaging over SPECT are its much higher sensitivity, spatial resolution, and quantitative imaging; therefore, annexin V has been radiola-beled with fluorine 18 (18F) and many other isotopes for positron emission tomography.

18F-labeled annexin V with N-succinimidy-4-18F-fluorobenzoate (18F-SFB) has been investigated by several groups [29, 30] [31]. These studies of 18F-SFB annexin V demonstrated comparable apoptotic imaging feasibility to 99mTc-labeled annexin V and a fast clearance [31]. Moreover, 18F-SFB annexin V showed a significant higher accumulation in the mice treated with doxorubicin when compared to the control group [30].

Annexin V can also be labeled with thiol-reactive agents such as N-substituted maleimides, and iodoacetamide can be used to modify proteins at cysteines at spe-cific sites [32].18F-N-[2-(4-fluorobenzamido)ethyl]maleimide (18F-FBEM) was used to label thiol-containing proteins as a novel site-specific labeling prosthetic group [33, 34]. Compared to the previous generation of 18F-SFB-labeled annexin V, the novel 18F-FBEM-cys-annexin V showed faster renal and a lesser extent of hepa-tobiliary excretion in normal mice and more sensitivity of site-specific detection in the rats of hepatic apoptosis model [35].

14.4 MRI Imaging

One of the main differences between magnetic resonance imaging (MRI) scan and other imaging modalities like PET is that MRI scan which reveals high spatial reso-lution allows scientists to navigate through the entire living organism, down to the cellular level. Several annexin V-based contrast agents have been developed. However, due to the fundamentally low sensitivity of MRI, how to deliver sufficient contrast agents safely and acquire sufficient imaging signals in vivo is definitely the concern.


238

14.4.1 T-Positive Images: Gadolinium-Labeled Annexin V

To access the redistribution of phosphatidylserine in the event of apoptosis, annexin V was linked to gadolinium diethylenetriamine pentaacetate (Gd-DTPA)-coated liposomes [36]. A significant increase in signal intensity was visible in those regions containing cardiomyocytes in the early stage of apoptosis. The in vivo Gd-DTPA- annexin V MRI imaging provided a rapid targeting of apoptotic cells in the ischemic and reperfused myocardium. Moreover, van Tilborg and his colleagues reported Gd-DTPA-bis(stearylamide) (Gd-DTPA-BSA)-labeled annexin V, the multiple functions of lipid-based bimodal contrast agent, enables the detection of apoptotic cells with both MRI and optical techniques [37]. Gd-DTPA-BSA was covalently coupled multiple human recombinant annexin V to introduce specificity for apop-totic cells. The imaging results showed a significant increase of the relaxation rates of apoptotic cell pellets when compared to the untreated control cells, which may have applications for the in vivo detection of apoptosis. In 2010, the same group developed a small micellar annexin A5-functionalized nanoparticle for noninvasive MRI and fluorescent imaging of PS exposing cells in atherosclerotic lesions [38]. In vivo MRI images of the abdominal aorta in atherosclerotic ApoE(−/−) mice revealed enhanced uptake of the annexin A5-micelles as compared to control micelles, which was corroborated with ex vivo near-infrared fluorescence images of excised whole aortas.

14.4.2 T2-Negative Images: Iron Oxide-Labeled Annexin V

Compared to T1 agents, superparamagnetic iron oxide nanoparticle-based T2 agents are assumed to be the preferred MRI contrast agents for evaluating apoptosis due to their high sensitivity. Up to date, the common labeling approach for apoptotic imag-ing is based on cross-linked derivative of monocrystalline iron oxide (MION), also known as cross-linked iron oxide (CLIO) [39].

Annexin V-CLIO allowed the identification of cell suspensions containing apop-totic cells by MRI even at very low concentrations of magnetic substrate [40]. Van Tilborg et al. investigated the internalization of, when co-exposed to apoptotic stim-uli, annexin A5 was shown to internalize into endocytic vesicles by using annexin A5-functionalized iron oxide particles [41].

Recently, our group present annexin V conjugated with superparamagnetic iron oxide nanoparticles (USPIO-annexin V) to the mice with Fas-induced hepatic apoptosis (data unpublished). The results showed that USPIO-annexin V accumu-lated in the region of hepatic apoptosis significantly decreased in comparison with control group (p< 0.05) (Fig. 14.1). USPIO-annexin V MRI may provide useful properties such as quantitative pharmacologic hepatic apoptosis that can be used as an indicator for hepatitis or liver injury induced by chemotherapy or after radiation exposure.

S.H.-H. Yeh et al.

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14.5 Optical Imaging

Petrovsky et al. first demonstrated that annexin V-labeled fluorophore Cy5.5 could be used as a nonradioactive probe for apoptosis [42]. Later in 2004, the modified annexin V-Cy5.5 conjugate was used to measure the tumor response to chemother-apy by fluorescence molecular tomography (FMT). This probe provided higher quantification accuracy validated by histology when compared to the traditional planar illumination methods [43]. The quantitative results also showed tenfold increase of fluorochrome intensity in cyclophosphamide-sensitive tumors and a sev-enfold increase of resistant tumors compared with controls. Smith et al. developed a fluorescent imaging probe conjugating zinc(II)-dipicolylamine (Zn-DPA) with annexin V [44]. In vivo studies demonstrated that the fluorescent Zn-DPA targeting ligand selectively targeted to the apoptotic tumor cells was consistent with ex vivo biodistribution and histological analyses [45].

14.6 Multiple Imaging Modality

Multiple imaging modalities generate more informative and effective imaging in the diagnosis and treatment of a large number of diseases, particularly if the machine combines both functional and anatomical imaging modalities. By using multiple imaging instruments, researchers can track multiple molecular targets

Fig. 14.1 T2-weighted in vivo MR images at the abdominal region. Images acquired at 1 hour after administration USPIO. Control mice (upper panel, a–c) and anti-Fas-induced hepatic apop-tosis (lower panel, e–g). li liver, ki kidney, sp spleen, pa pancreas (Adapted from Yeh et al. [unpub-lished data])


240

simultaneously and obtain more accurate localization and precise expression of bio-markers [46].

AnxCLIO-Cy5.5, the first magneto-optical nanoparticle, can be used as a bifunc-tional tracer in MRI and fluorescence imaging [47]. The in vivo images demon-strated that myocardial T2 signals of AnxCLIO-Cy5.5 were significantly lower in the mice receiving transient coronary artery (LAD) occlusion, and fluorescence tar-get to background ratio was significantly higher when compared to the controls [48]. In addition, annexin V-conjugated quantum dots with a paramagnetic lipidic coating (Gd-DTPA) for MRI and fluorescent imaging showed high specificity for detecting apoptotic cells [38, 49].

As an alternative to MRI/optical imaging, nuclear/optical imaging was also developed for the detection of apoptosis. Zhang et al. evaluated 111In-labeled annexin A5-conjugated core-cross-linked polymeric micelles (CCPM) for micro-single- photon emission tomography/computed tomography (μSPECT/CT) and fluores-cence molecular tomography (FMT) imaging in various disease models including tumor apoptosis, hepatic apoptosis, and inflammation. [50] [51]. Zhang et al. pro-vided the clue that multiple imaging techniques should be advantageous in assess-ing and validating early diagnosis and therapeutic responses in diseases associated with apoptosis.

14.7 Conclusions and Perspectives

Over the past two decades, there have been many tracers proposed by using differ-ent modalities for apoptosis imaging, but none of them yet has achieved fully the validation for the differential localization or biochemical cellar progression of apop-tosis. In this review, we focus on imaging agents conjugated with annexin V by using different imaging modalities such as single-photon emission computed tomography (SPECT), positron emission tomography (PET), optical imaging (OI), and magnetic resonance imaging (MRI). Each modality allows for the in vivo non-invasive detection of apoptotic cells and cell products. Not surprisingly, multimodal imaging, combining two or more of these techniques (PET/MRI or SPECT/CT or optical/CT), will become a key player for basic and translational medicine in humans and animals in the future, despite the challenges when considering acquir-ing and combining nonredundant images as well as imaging time, throughput, and cost of technology.

However, for the development of apoptosis-detecting imaging agents, there are several concerns that should be taken in mind such as the pharmacokinetic/pharma-codynamics of new agents, signal to background ratio in the abdominal region, and differentiation of apoptosis and necrosis. Consequently, we believe that all of these factors will be integrated and clear obstacles to introduce a successful apoptosis imaging agent.

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