27 nuclear medicine imaging instrumentation

27.1

CHAPTER 27

NUCLEAR MEDICINE IMAGINGINSTRUMENTATION

Mark T. MadsenUniversity of Iowa, Iowa City, Iowa

27.1 INTRODUCTION 27.127.2 SCINTILLATION CAMERAS 27.227.3 SPECT SYSTEMS 27.1427.4 SUMMARY 27.20REFERENCES 27.20

Nuclear medicine is a diagnostic imaging modality that is used to obtain clinical information aboutmost of the major tissues and organs of the body. Diagnostic information is obtained from the waythe tissues and organs process radiolabeled compounds (radiopharmaceuticals). The radiopharmaceu-tical is typically administered to the patient though an intravenous injection. The radiopharmaceuticalis carried throughout the body by the circulation where it localizes in tissues and organs. Images ofthese distributions are acquired with a scintillation camera. Ideally, the radiopharmaceutical would goonly to abnormal areas. Unfortunately, this is never the case and the abnormal concentration of theradiotracer is often obscured by normal uptake of the radiopharmaceutical in the surrounding tissues.Images of higher contrast and better localization can be obtained with tomographic systems designedfor nuclear medicine studies (SPECT systems). These are described in detail below.

The imaging of radiotracers in the body presents special challenges that are unique. The flux ofgamma rays available for imaging is orders of magnitude less than that used in x-ray radiography orx-ray computed tomography (CT). In addition, the high energy of the gamma rays makes detectionmore difficult. As a result, the images produced in nuclear medicine studies are much noisier andhave worse spatial resolution. In order to appreciate these problems and how they affect the design ofnuclear medicine imaging devices, we will briefly review the physics of gamma ray interactions.1

The intensity of gamma rays traveling through material is gradually reduced by absorption orscattering. This loss of gamma rays is referred to as attenuation and is described by the exponentialequation

(27.1)

where I0

= initial intensityI(x) = intensity of rays after traveling a distance x through the material = linear attenuation coefficient of the material

27.1 INTRODUCTION

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27.2 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION

Over the range of gamma ray energies used in radionuclide imaging, the two primary interactionsthat contribute to the attenuation coefficient are photoelectric absorption and Compton scattering.Photoelectric absorption refers to the total absorption of the gamma ray by an inner shell atomicelectron and is the primary interaction in high-Z materials such as sodium iodide (the detectormaterial used in the scintillation camera) and lead. In low-Z materials such as body tissues, itscontribution to attenuation is relatively small. Compton scattering occurs when the incoming gammaray interacts with a loosely bound outer shell electron. A portion of the gamma ray energy is impartedto the electron and the remaining energy is left with the scattered photon. The amount of energy lostin the event depends on the angle between the gamma ray and scattered photon. Compton scatteringis the dominant interaction in body tissues.

High attenuation is desirable in detecting and shielding materials. Ideally materials used for thesepurposes would absorb every gamma ray. In the body, attenuation is very undesirable, butunfortunately, unavoidable. Attenuation reduces the intensity of gamma rays available for detectionand scattered radiation that reaches the detector causes a significant loss of contrast.

The scintillation camera is the primary imaging instrument used in nuclear medicine and is oftenreferred to as a gamma camera.2 The scintillation camera is a position-sensitive gamma ray imager.Although the entire field of view is available for detection, it processes one event at a time. Thespatial resolution is approximately 10 mm and it yields a count rate of 200 to 300 cpm/Ci in thefield of view (cpm = counts per minute). The field of view covers a large portion of the body and istypically 40 50 cm, although other sizes are available.

The rectilinear scanner was the first practical nuclear medicine imaging device and it was still inuse through the 1970s. It was invented by Benedict Cassen in 1950. The rectilinear scanner used thedetected count rate of a radiation detector to control the brightness of a small light bulb masked toexpose a small area of a film. The movement of the radiation detector and the bulb were linked sothat, as the detector moved in a raster pattern over the patient, the bulb tracked a corresponding pathover the film. The developed film revealed the internal distribution of the radiopharmaceuticalsampled by the scanning probe. While the invention of the rectilinear scanner was a major stepforward, it had several shortcomings that were inherent in its design. Because it sampled only onesmall area at a time, image acquisitions took a long time (10 to 20 minutes). In addition, only staticimaging was possible. Dynamic studies, such as those that followed the progression of a tracerthrough the body, required an imaging system with a large field of view where all areas are equallysampled, i.e., a scintillation camera.

The first scintillation camera was developed by Hal O. Anger in 1958.3 Although this system wasvery crude, it contained the fundamental components of all future designs: NaI(Tl) as the primarydetector and weighted signals from an array of photomultiplier tubes to determine the location ofdetected events. Table 27.1 gives typical performance values for a modern scintillation camera.

The gamma raysensitive element of the scintillation camera is a large, thin piece of NaI(Tl).Although the crystals originally had a circular cross section, most scintillation cameras now use arectangular crystal with dimensions as large as 40 50 cm. The thickness of NaI(Tl) in mostconventional cameras is 9.5 mm, but in systems that are used for coincidence detection, the crystalmay be twice as thick. NaI(Tl) is a scintillator; It converts gamma ray energy into visible light. Theamount of light generated is directly proportional to the absorbed energy. NaI(Tl) is very efficient atthis and the absorption of one 140-keV gamma ray will yield 5000 visible light photons. There area number of advantages associated with NaI(Tl) in addition to its high light output. It efficientlyabsorbs 140-keV gamma rays (with a photopeak efficiency of 85%) and it has a moderate energyresolution. Energy resolution is an important property since it provides the means to discriminateagainst scattered radiation. Gamma rays that undergo scattering within the patient degrade the qualityof images. However, scattered gamma rays necessarily have less energy than unscattered gamma raysand can be selectively eliminated on that basis. Another positive feature of NaI(Tl) is that it can be

27.2 SCINTILLATION CAMERAS



NUCLEAR MEDICINE IMAGING INSTRUMENTATION

NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.3

manufactured in many shapes and sizes. There are disadvantages though. NaI(Tl) actively absorbswater vapor from the air and loses its transparency. It must be hermetically sealed, and loss of this sealresults in irreparable damage. Another disadvantage is that the persistence of the scintillation is longenough that it limits the count rate that the crystal can accurately handle. Most nuclear medicineimaging is performed far below this limit. However, some first pass studies do result in significantcount rate losses from this limit. The biggest problem is encountered in coincidence imaging.

Converting the gamma ray energy to visible light is only part of the battle. In order for theinformation from the scintillation to be useful, it has to be converted into an electronic signal. This isaccomplished with a photomultiplier tube (Fig. 27.1). The photomultiplier tube is a vacuum tube witha photoemissive surface called the photocathode. Visible light hitting this surface knocks offelectrons. These electrons are accelerated to an electric terminal called a dynode. The first dynode has

TABLE 27.1 Scintillation Camera Specifications

FIGURE 27.1 Photomultiplier tube. The photomultiplier tube converts the scintil-lation into an electronic pulse preserving the linear relationship between the mag-nitude of the scintillation and the energy of the interaction. The location of a sourcecan be inferred from the magnitude of the signal change. The relationship betweenthe signal magnitude and source position is nonlinear, and positioning errors occurboth when the source is far from the PMT and when it is directly under it. The idealresponse can be approximated with the use of a light pipe.





a potential approximately 100 volts higher than the photocathode, and the electrons hit it withenough force to knock off about 4 more new electrons. The next dynode is another 100 volts higher,so the process is repeated. The same process occurs over a total series of nine dynodes resulting in asignal amplification of 1,000,000. Proportionality is maintained throughout this amplification so thatthe size of the electron pulse is directly proportional to the energy deposited by the gamma ray.

A scintillation camera needs to be able to record gamma ray events over a large area. This requiresuniform sampling by an array of photomultiplier tubes (PMTs). The PMTs are arranged in a closepacked array that covers the entire surface of the NaI(Tl) crystal (Fig. 27.2). The PMTs used inscintillation cameras are usually 2 or 3 in across, so that as many as 120 PMTs may be used. PMTshave been manufactured in variety of different cross sections in order to maximize their arealcoverage. Circular, hexagonal, and square tubes have all been used. The signals obtained from thePMTs will be used to determine two important properties about the gamma ray interaction: where didit occur and how much energy was deposited? At first blush, it may seen that even 2-in PMTs are toocoarse to determine the event location. However, we will see that magnitude of the PMT output isfairly sensitive to the location of a source.

If a PMT is mounted to an NaI(Tl) crystal and the signal output is plotted as a source is movedfrom left to right, the result shown by the solid line (Fig. 27.1) will be obtained. When the source ispositioned far from the PMT, the signal is weak and the location of the source is not certain. Whenthe source is directly under the PMT, the signal is strong, but the dependence on the position ismodest. However, when the source is just to the left or right of the PMT, the signal change withsource location is large. Over this region, the location of the source can be accurately tracked. Withonly a single PMT, we could not tell if the source was on the left or right (or front or back), but byconsidering the signals from other surrounding PMTs, that can be determined. The main problem iswhat to do about the poor response near the center of PMT. What can be done to transform themeasured PMT output into the ideal signal shown in Fig. 27.1? The early solution to this was todisplace the PMTs from the NaI(Tl) crystal with a light pipe.

A light pipe is a transparent material such as lucite or quartz that is interspersed between the crystaland PMT array. The displacement of the PMTs causes the light from the scintillations to spread out,yielding a more favorable signal output. To achieve even better results, some manufacturers haveused special mask patterns on the front surface of the light pipe along with sculptured grooves in theback. These efforts pay off in more accurate positioning of the event locations; however, there is aprice to pay. The effort to scatter the light also results in higher light losses, leading to increases instatistical fluctuations and ultimately degradations in spatial resolutions. In the early gamma cameras,there was no way around this dilemma. However, digital electronics provides the capability of

FIGURE 27.2 Photomultiplier tube array. The photomultiplier tubes arearranged in a close-packed array to cover the back surface of the NaI(Tl)crystal.





nonlinear mapping, which greatly reduces the demands put on the light pipe. This has allowed lightpipes to be made very thin and even eliminated.

Figure 27.3 shows a schematic of an analog scintillation camera. The scintillation light from anabsorbed gamma ray is transmitted through to the photomultiplier tube array. The energy of the eventis determined by summing all the photomultiplier tube signals. As will be seen, the energy signal isused both for scatter discrimination and for normalizing the position signals. The position of theevent is determined by summing weighted outputs from each photomultiplier tube. These weightingfactors are determined by the location of the photomultiplier tube in the array. Separate weightingfactors are used for determining the x and y signals. This process is referred to as Anger logic, sinceit is the scheme developed by Hal Anger in the first scintillation camera. In the initial designs, literallyall the photomultiplier tubes participated in the energy and position signal summations. It wassubsequently found that the signals from photomultiplier tubes located far from the event contributedmostly noise. In modern designs, the photomultiplier tube signal must exceed a threshold before it isincluded in the sum. Another point that should be made is that all the processing is performed on eachdetected event. The decision to include the event as a valid count is not made until the end of theprocessing when the pulse height analysis is done. If the event falls within the selected energywindow, the normalized x and y signals are available for either an analog display or digital storage.

The position signals determined from summing the weighted PMT signals vary with the brightnessof the scintillation, which itself depends on the energy absorbed in the crystal. This means that anobject imaged with a high-energy gamma ray like I-131 (364 keV) will be magnified when comparedto the same object imaged with Tc-99m (140 keV). This magnification is a concern even when onlyone gamma ray energy is imaged because of the finite energy resolution of the scintillation camerasystem. The pulse heights from the absorptions of identical gamma rays vary enough to cause slightminifications and magnifications, ultimately degrading spatial resolution. The solution to this problemis to normalize the position signals with the measured energy signal. This removes the image sizedependence with energy, thereby improving spatial resolution and allowing the simultaneous imagingof more than one radionuclide without distortion. This feature is the primary component forguaranteeing good multiwindow spatial registration.

FIGURE 27.3 Analog scintillation cameras. The signals from each photomulti-plier tube are sampled to determine both the position and energy of the detectedevent. Separate weighting factors are used for the x and y signal determinations. Theenergy signal is used to normalize the position signals and discriminate againstscattered radiation.





As has been previously noted, gamma rays that are scattered within the patient have distortedspatial information and degrade image contrast. Because scattered gamma rays necessarily loseenergy, they can be selectively avoided by accepting only events that have pulse heightscorresponding to the primary gamma ray energy. The pulse height analyzer provides this capability.A window is centered to cover 15 to 20 percent of the photopeak. All energy pulses that meet thiscriterion generate a logic pulse that indicates to the system that a valid event has occurred. This outputenables the recording of the x and y position information in a computer or display.

The precision of locating gamma ray events by a scintillation camera is referred to as intrinsic spatialresolution. The original Anger camera had very poor intrinsic spatial resolution (~12 mm full-width-half-maximum). With the improvement in electronics and pulse processing methods, the spatialresolution of the scintillation camera has improved steadily, approaching 3 mm in modern systems.4

These improvements include: better-quality, low-noise PMTs; improvements in the Anger logicelectronics including signal thresholding; smaller PMTs, and improved PMT quality control. The mostrecent improvement has been the replacement of most of the analog processing with digital electronics.

With improvements in the speed of digitization electronics and decreases in component costs, thetrend in scintillation cameras has been to digitize the PMT signals (Fig. 27.4).5 The analog-to-digitalconverters assign a numeric magnitude to the PMT signals. All subsequent determinations of energyand positions can then be done by computer algorithms that can accurately model the nonlinearbehavior of the PMT signals with source position. The success of this approach has allowed thereduction and even total elimination of the light pipe. This moves the PMTs closer to the scintillation,thereby improving both the precision of the position determination (i.e., spatial resolution) and theenergy resolution.

Once there is an x and y coordinate that locates a valid event, this information has to be stored asimage data. Although it is possible on some scintillation camera systems to store the individualcoordinates sequentially (referred to as list mode acquisition), most systems store the informationdirectly in histogram or matrix mode. With this method, an array of computer memory, typically 128 128 or 256 256, is reserved for each image frame. The matrix elements or pixels are initially setto 0. The coordinates for each event point to one of the pixels and this pixel is incremented by 1.When the acquisition-stopping criteria are met, the image is complete. The information in the matrixis either gray-scale or color encoded to display the image data. The entire process is shownschematically in Fig. 27.5. A gamma ray originating in the patient is absorbed in the NaI(Tl) crystal.

FIGURE 27.4 Digital scintillation cameras. By digitizing the output of thePMTs, the analog weighting electronics can be replaced by a nonlinear po-sitioning algorithm. This allows a more accurate correction, culminatingwith the elimination of the light pipe.





The light from the scintillation is sampled by the PMT array, which determines both the x and ycoordinates of the event and its energy. If the energy signal falls within the window of the pulseheight analyzer, the x and y coordinates are used to increment the appropriate pixel. This process isrepeated for every detected event.

In order to form images with a scintillation camera, a collimator must be placed in front of theNaI(Tl) crystal. The collimator (Fig. 27.6) is the image-forming aperture of the camera system, andit is necessary for the imaging process. The collimator projects the gamma ray pattern originating inthe patient onto the NaI(Tl) crystal. It does this by selectively absorbing gamma rays. The collimatoris a close-packed array of holes in a lead plate. Most often the holes are parallel, but fanbeamconverging and diverging collimators are available. Gamma rays whose trajectory takes them througha hole get to interact with the NaI(Tl). All the others are absorbed. The design of collimators dependson the gamma ray energy and the ever-present trade-off between count sensitivity and spatialresolution.6,7 Collimators used for imaging Tc-99m typically have holes that are 1 to 1.5 mm acrossand are 20 to 40 mm thick. Typical collimator design parameters are given in Table 27.2.

Although a collimator is necessary for the formation of images, it represents the limiting factor inthe both count sensitivity and spatial resolution of the scintillation camera. Because of the bruteforceabsorption approach to forming images with collimators, they are very inefficient. Less than 1 in5000 gamma rays that hit the front surface of the collimator get through to the crystal. To improvethe count sensitivity, the collimator hole size could be increased and the hole length shortened.Unfortunately, these changes degrade the spatial resolution. The spatial resolution of the collimator isconstrained by the geometry of the holes and is typically in the range of 6 to 8 mm at 10 cm whenused with Tc-99m. This is the dominant factor in determining the overall system resolution, since theintrinsic spatial resolution is in the range of 3 to 4 mm.

One very important property to remember about collimators is that the spatial resolution getsworse as the source-to-collimator distance increases. This is illustrated in the set of phantom imagesthat were acquired from 5 to 30 cm from the collimator surface. To obtain the best-quality images,spatial resolution comes at the price of count sensitivity; therefore it is crucial to keep the collimatoras close to the patient as possible.

The modern scintillation camera has improved performance because of improvements in thecomponents and electronics. The availability of digital electronics has allowed the elimination of the

FIGURE 27.5 Scintillation camera. The scintillation camera processes eachdetected event to determine the x, y, and energy. If an event falls within theselected energy range, the memory location pointed to by the x and y coor-dinates is incremented. This process continues until the stopping criteria(number of counts or acquisition time) are met.





light pipe, which improves both energy and spatial resolution. However, this requires additionalcorrections because of the nonlinear response of the PMT array to the scintillations. If a collimatedpoint source were focused on a portion of the NaI(Tl) crystal located exactly on a photomultipliertube center, the energy spectrum would be distinctly different than one that was acquired from apoint between two tubes (Fig. 27.7). This difference reflects the efficiency for collecting all the

FIGURE 27.6 Collimation. The collimator is the image-forming aperture of the scintillationcamera. It projects an image of the radionuclide distribution onto the NaI(Tl) crystal by bruteforce absorption of all gamma rays except those whose trajectory takes them through the holes.The collimator is also the limiting factor of both spatial resolution and count sensitivity. Thespatial resolution significantly degrades with source-to-collimator distance.

TABLE 27.2 Collimator Specifications





scintillation light. When the source is located directly on a photomultiplier tube, more of thescintillation is sampled by the photomultiplier tubes, and the pulses are therefore somewhat larger onthe average than at other points. This position-dependent shift in the energy spectrum causes anoverall loss in energy resolution. It also means that portions of the crystal will be acceptingproportionately more scattered radiation. The solution to this problem is to locally sample the energyspectra and regionally adjust the energy window for each area. Typically the camera field of view isdivided into a 64 64 matrix and energy window adjustments are made for each of the 4096 regions.

Figure 27.7 shows the effect of energy correction when the scintillation camera is exposed to auniform flux of gamma rays. First, it should be noted that both the corrected and uncorrected imagesare highly nonuniform and are not adequate for imaging. The energy correction simply makes surethat each region of the crystal is contributing valid photopeak events to the image. This results in onlya subtle improvement in uniformity at this stage. However, it makes the subsequent corrections morerobust, since there will be much less dependence on the effects of scattered radiation, which can varyover a large range, depending on the imaging situation.

Because of the nonlinear response of the photomultiplier tubes, detected events are not correctlypositioned using Anger logic alone. This mispositioning of events has a profound effect on fielduniformity.8 The parameter that quantifies how well-detected events are positioned is called spatiallinearity. The optimization of spatial linearity requires the acquisition of an image from a welldefineddistribution (Fig. 27.8). Typically this is accomplished with a highly precise rectangular hole patternthat is placed directly on the NaI(Tl) crystal. A distant point source of radioactivity is used to projectan image of the hole pattern onto the scintillation camera. The image of this pattern appears similarto the image on the left with distortions caused by the mispositioning of events. Because the actualand measured location of the holes is known, regional displacements to the x and y coordinates canbe calculated for each hole. Displacements for regions in between holes that are not directly sampledare interpolated at a very high sampling frequency (1024 1024). This information is stored and isavailable as a lookup table. This measurement is usually done by the vendor at the factory and maybe repeated several times a year. When a valid event is detected, the initial x and y coordinates aremodified by the appropriate displacements that are read from the lookup table. Using this approach,events can be accurately positioned to better than 0.5 mm. The improvement in spatial linearity hasa profound effect on field uniformity. Both images show the response of the scintillation camera to a

FIGURE 27.7 Energy correction. Because of the spatial arrangement of thePMTs, the magnitude of the energy signal varies with location, degrading theoverall energy resolution. This problem is overcome by setting multiple localenergy windows across the field of view. This energy correction does not im-prove uniformity, but it does remove the dependence of the scintillation cam-era on scatter conditions.





uniform flux of gamma rays. With spatial linearity correction, the field becomes uniform to within10 percent of the mean image counts. This level of uniformity is adequate for most conventionalimaging.

There are still some residual nonuniformities that exist in the scintillation camera even after energyand spatial linearity correction have been applied. These can be further reduced by applyinguniformity correction (Fig. 27.9). Typically, a high count flood is acquired and a map of thenonuniformities is stored in a memory buffer. During acquisition, the number of valid events that isacquired is modulated by this map to ensure uniformity. With this additional correction, the fielduniformity can be reduced to within 3 percent of the mean image counts. It should be noted thatfield uniformity can be degraded by a number of factors, including the energy of the gamma ray.The most crucial factor for a system that is operating properly is the setting of the energy window.Figure 27.9 illustrates the dependence of uniformity with the energy window setting. Somescintillation cameras are more forgiving than others, but all show more nonuniformity when the energywindow is displaced from the center of the photopeak.9 There is often a gamma ray energy dependenceas well. Most scintillation cameras are optimized for the best performance for the 140-keV gamma raysof Tc-99m. In some systems, uniformity significantly degrades at other gamma ray energies.

Photomultiplier tubes are relatively unstable components. Their performance changes as they ageand is also sensitive to variations in temperature and humidity. In order for the energy, spatiallinearity, and uniformity corrections to remain valid, there must be some way of maintaining thephotomultiplier tubes at a constant operating point. Most scintillation camera systems havephotomultiplier tube stabilization firmware that dynamically adjusts the photomultiplier tubes inresponse to a known reference signal. Some vendors use a constant-output light-emitting diode insidethe electronics housing that flashes 10 times per second. The individual photomultiplier tube signalsfrom these calibration flashes are sensed by electronics that can compare the measured output to thedesired value, and then make appropriate adjustments to maintain the operating point. Anotherapproach uses the ratio between the count rates in a photopeak and scatter window to maintainconstant photomultiplier tube response. Photomultiplier tubes are also very sensitive to magnetic

FIGURE 27.8 Spatial linearity correction. Residual positioning errorsare corrected by imaging a precision hole phantom. A correction factortable is generated with the appropriate x and y offsets to reposition eventsto their correct location. The application of spatial linearity correctionhas a profound effect on image uniformity.





fields, and changes in the orientation of the camera with respect to the earths magnetic field areenough to cause measurable changes in field uniformity.10 To reduce this effect, each photomultipliertube is individually surrounded with mu-metal shielding.

The emission of gamma rays from a radioactive source has an exponential distribution. Thismeans that for any particular event rate, short intervals between events occur much more often thanlong intervals. Because the scintillation light persists for a finite time, there will eventually be lightemitted from more than one event as the count rate increases. In a scintillation camera, pulse pileupbecomes evident at count rates as low as 20,000 counts per second (cps), and it gets increasinglyworse as the count rate increases. Since information about pulse height becomes compromised, theperformance of the scintillation camera degrades at high count rates. In most conventional andSPECT imaging, the count rate is low enough that pulse pileup is not a major concern. However, highcount rates are encountered in some first-pass studies, and it is the primary problem in coincidenceimaging. Corrections can be made for pulse pileup, since the physics of scintillations is well known(Fig. 27.10). Pileup can be detected on the basis of pulse height analysis. If the pileup is the result ofonly two events, the contribution to the second pulse from the first can be accurately estimated andsubtracted, thus preserving both events. Multiple pileups can be identified and discarded. The countrate performance of scintillation cameras has improved dramatically in recent years because of thedemands of coincidence imaging. When a conventional scintillation camera is recording count data at120,000 cps, it is losing about 20 percent of the valid events. This loss increases with increasing countrate. In addition to the loss of sensitivity, both uniformity and spatial resolution get progressivelyworse as the count rate increases. Finally, it should be recognized that Anger logic will produceartifacts at very high count rates. Virtual sources will appear midway between real sources becauseof the signal averaging.11 This problem has been addressed in coincidence systems by using maximumlikelihood estimation instead of the conventional Anger logic.

FIGURE 27.9 Uniformity correction. Nonuniformities in the field that remain after spatial linearitycorrection are corrected from the acquisition of a high count reference flood image. In many cameras, thefield uniformity degrades when the energy window is not centered on the photopeak.





The performance of scintillation cameras is often specified in terms of the spatial resolution.Spatial resolution is a measure of image blur and it is often specified by the width of the pointspreadfunction. On a perfect imaging system, the image of a point has no dispersion so that all the countsfall at the same point. On a real imaging system the point is blurred and a count profile through it hasa gaussian shape. The width of this count profile at the half-maximum level (FWHM) is a commonlyused method for specifying the spatial resolution. The imaging system will more or less blur everypoint in a similar way, leading to a loss of contrast. There are two spatial resolution parameters thatare of interest with scintillation cameras. The intrinsic spatial resolution describes how precisely theevent location is determined when a gamma ray interacts with the crystal. The extrinsic or systemspatial resolution combines the effect of the collimation with the intrinsic resolution. The intrinsicspatial resolution varies from about 3 mm FWHM on systems with thin crystals and 2-in PMTs toabout 4.5 mm on systems with thicker crystals and 3-in PMTs.

The system spatial resolution depends on a variety of factors including the gamma ray energy,collimation, the source to collimator distance, and the intrinsic spatial resolution. Higher-energygamma rays require thicker septal walls, which limits the resolving power of the collimator. Asindicated above, there is a strong dependence on the distance that the source is from the collimator.The collimator and intrinsic resolution combine in quadrature like the sides of a right-angle triangleto yield the system resolution. The collimator resolution is generally more than 50 percent larger thanthe intrinsic resolution and therefore is the dominant factor. The system resolution at 10 cm rangesfrom 6.5 mm for an ultrahigh-resolution collimator to about 9.5 mm for a general-purposecollimator. It should be noted that there is nearly a factor of 3 loss of count sensitivity with theultrahigh-resolution collimator compared to the general-purpose collimator.

Scintillation cameras have the capability of simultaneously imaging different-energy gamma rays.Most scintillation cameras handle at least three and many can handle six or more energy windows. Itis important that there is no significant distortion of the images obtained at the different energies. Theparameter that monitors the correspondence between images acquired at different energies is referred

FIGURE 27.10 Pulse pileup. The persistence of the scintillation causes pulse pileup when the event rateexceeds 50,000 counts/s. Techniques can be applied to recover some of the lost information. This is espe-cially important for coincidence imaging. Note also that the Anger positioning algorithm causes ghostimages when the count rate becomes high.





to as multiwindow spatial registration. The multiwindow spatial registration should be less than 3 mmon a modern scintillation system.

In the future there will be further improvements in scintillation camera technology.12 Theseinclude photomultiplier tube replacements, new scintillators, and solid-state alternatives.Photomultiplier tubes are expensive, bulky, and susceptible to drifting. Avalanche photodiodes are apossible replacement for photomultiplier tubes (Fig. 27.11). The photodiodes use solid-statecomponents that are very compact and rugged. Photomultiplier tubes still have superior performancespecifications and are better matched for NaI(Tl). But considerable progress has been made with thephotodiode approach and special-purpose devices that use this technology are beginning to appear.Figure 27.11 shows a schematic for a miniature camera module that uses avalanche photodiodes. Thescintillator used in this application is CsI(Tl). The scintillation light emitted for CsI(Tl) is bettermatched to the properties of the photodiode. Because of the compactness offered by such designs,scintillation cameras can be designed in novel ways for specific applications. One potentialapplication is breast cancer imaging, where several of these devices could be arranged around thebreast to collect SPECT data more efficiently than bulky conventional scintillation cameras.

Solid-state detectors directly convert the absorbed gamma ray energy into collection of electriccharge and do not need photomultiplier tubes. Since the PMT is a bulky and expensive component,this represents a significant breakthrough. Cadmium zinc telluride is an attractive solid-state detector13

(Fig. 27.12). It can be manufactured in a pixelated array and has comparable gamma ray detection

FIGURE 27.11 PMT replacements. PMTs are an expensive and bulky component of thescintillation camera. Future designs may incorporate avalanche photodiodes.





efficiency to NaI(Tl) at 140 keV. Another advantage is that the energy resolution with CdZnTe isnearly a factor of 2 better than that of NaI(Tl). However, several production problems need to besolved before CdZnTe becomes a viable replacement for NaI(Tl).

Single-photon-emission computed tomography (SPECT) produces tomographic images of the inter-nal distribution of radiopharmaceuticals.14,15 It is most commonly used in the diagnosis of coronaryartery disease and in tumor detection. Projection images collected by one or more scintillationcameras are mathematically reconstructed to obtain the tomographic slices. Most clinical SPECTstudies are qualitative with simplistic corrections for attenuation and scattered radiation. QuantitativeSPECT requires corrections for attenuation, scatter, and spatial resolution, although these have notbeen routinely implemented in the past because of their computational load. SPECT instrumentationhas evolved to include coincidence imaging of positron-emitting radiopharmaceuticals, specifically18F fluorodeoxyglucose.

A SPECT system consists of one or more scintillation cameras mounted to a gantry that canrevolve about a fixed horizontal axis (the axis of rotation)1619 (Fig. 27.13). SPECT studies areusually acquired over a full 360 arc, although myocardial perfusion studies typically use only datafrom the 180 arc that minimizes tissue attenuation. SPECT acquisitions are performed with thescintillation camera located at preselected angular locations (step-and-shoot mode), or in acontinuous rotation mode. In the step-and-shoot mode, the detector rotates to each angular positionand collects data in a preselected frame duration while the detector is motionless. In the continuous-rotation mode, the study duration is selected and the rotation speed is adjusted to complete the orbitduring this time. Projections are collected as the detector rotates and are binned into 60 to 120frames over 360.

It is crucial to maintain close proximity to the body as the detector rotates about the patient toachieve the best possible spatial resolution. Although a number of different approaches have beenused to accomplish this, the most common method moves the detectors radially in and out as afunction of rotation angle. Myocardial perfusion studies are the most commonly performed SPECTprocedures. Because the heart is located in the left anterior portion of the thorax, gamma raysoriginating in the heart are highly attenuated for views collected from the right lateral and rightposterior portions of the arc. For this reason, SPECT studies of the heart are usually collected usingthe 180 arc that extends from the left posterior oblique to the right anterior oblique view.20 Thisresults in reconstructed images with the best contrast, although distortions are often somewhat morepronounced than when 360 data are used.21 Because of the widespread use of myocardial perfusionimaging, many SPECT systems have been optimized for 180 acquisition by using two detectors

FIGURE 27.12 Solid state imagers. Although NaI(Tl)-based scintillation cam-eras will be the choice for the near future, improvements in solid-state detectorssuch as CdZnTe may lead to competitive imaging systems.

27.3 SPECT SYSTEMS





arranged at ~90 (Fig. 27.14). This reduces the acquisition time by a factor of 2 over single detectorsand is approximately 30 percent more efficient than triple detector SPECT systems. Positioning thedetectors at 90 poses some challenges for maintaining close proximity. Most systems rely on themotion of both the detectors and the SPECT table to accomplish this.

The heart is continually moving during the SPECT acquisition, and this further compromisesspatial resolution. Because the heart beats many times per minute, it is impossible to directly acquire

FIGURE 27.13 SPECT acquisition. One or more scintillationcameras collect images at typically 60 to 120 angles around a360 orbit. The scintillation camera acquires projection imagesfrom a large volume simultaneously.

FIGURE 27.14 SPECT system configurations. Although a single scintillation camera can beused to acquire SPECT data, multiple detectors improve the overall sensitivity. Two detectorsarranged at either 180 or 90 are the most common configuration.





a stop-action SPECT study. However, since the heart motion is periodic, it is possible to obtain thisinformation by gating the SPECT acquisition.22 In a gated SPECT acquisition, the cardiac cycle issubdivided and a set of eight images spanning the ECG R-R interval is acquired for each angularview. These images are acquired into predetermined time bins based on the patients heart rate, whichis monitored by the ECG R wave interfaced to the SPECT system. As added benefits of gating, themotion of the heart walls can be observed and ventricular volumes and ejection fractions can bedetermined.22,23

Although most SPECT imaging samples more or less static distribution of radionuclides, someSPECT systems can perform rapid sequential studies to monitor tracer clearance. An example of thisis determination of regional cerebral blood from the clearance of 133Xe.24 Multiple 1-minute SPECTstudies are acquired over a 10-minute interval. When one acquisition sample is completed, the nextbegins automatically. In order to minimize time, SPECT systems that perform these studies canalternately reverse the acquisition direction, although at least one SPECT system utilizes slipringtechnology so that the detectors can rotate continuously in the same direction.

In order to produce accurate tomographic images, projection data representing the line integralsof activity in the internal distribution have to be acquired. This information is not directly availablebecause of tissue attenuation. Simple attenuation correction methods can be used in regions of thebody such as the abdomen or head where the tissue density is more or less uniform. However,compensation for attenuation in the thorax requires an accurate attenuation map for eachtomographic plane. This is especially important for myocardial perfusion studies, since the artifactsresulting from tissue attenuation mimic the patterns caused by coronary artery disease. In recentyears, all the SPECT manufacturers have offered systems that can perform transmission measurementsalong with the emission studies. These systems use the scintillation camera to detect the transmissionof gamma rays from an external source.17,25,26 Several different configurations are available. Most usea line source that is translated across the camera field of view at each angular stop in much the sameway as a first-generation CT scanner (Fig. 27.15). Typically a source such as Gd-153 or Ba-133 is

FIGURE 27.15 Attenuation correction. To obtain accurate SPECT results, corrections must be made fortissue attenuation. In regions where there are large variations in tissue density such as the thorax, thisrequires an independent transmission study. This shows one possible configuration where line sourcesof Gd-153 are translated across the field of view to collect the transmission data. This information isreconstructed to obtain a crude CT image of the thorax to correct myocardial perfusion studies. (Cour-tesy of GE Medical Systems.)





used that has a different gamma ray energy than those used in the emission study so that the studiescan be acquired simultaneously (with appropriate correction for cross talk). Other approaches usemultiple stationary line sources to collect this information. At least one vendor has an x-ray tube andseparate detectors to obtain a moderate-quality CT scan.27 In each configuration, the transmission dataare collected and reconstructed to yield the attenuation maps. This information can be incorporatedinto an iterative algorithm to effect the correction.

In spite of the energy discrimination available on all SPECT systems, Compton scattered radiationstill accounts for about 30 to 40 percent of the acquired counts in SPECT imaging. Scattered radiationdecreases contrast and can impact other corrections. For example, when attenuation correction isapplied without also correcting for scattered radiation, the count density in the heart walls near theliver may be overenhanced. SPECT systems in the future may resort to other detectors that havesubstantially better energy resolution than that of NaI(Tl), but for now, scatter compensation routinesmust be employed. Scatter correction has been performed in several different ways.15,2833 The easiestto implement is the subtraction method, where information is simultaneously acquired into a secondenergy window centered below the photopeak in the Compton scatter region of the energy spectrum.After establishing an appropriate normalization factor, the counts from the scatter window aresubtracted from the photopeak window. The corrected projections are then used in the reconstructionalgorithm. The disadvantage of this approach is that it increases noise and it is difficult to establish anaccurate normalization factor. To accommodate this type of correction (and also to image different-energy gamma rays), SPECT systems allow the simultaneous acquisition from multiple energywindows. The number of energy windows available varies for each manufacturer, although everysystem is capable of imaging from at least four energy windows.

The details of SPECT image reconstruction are beyond the scope of this article. However, thedemands of image reconstruction do impact the features required by the computer. The CPU must befast enough and have enough memory to accommodate the entire SPECT data set. This is well withinthe capability of home PCs. A typical SPECT study is less than 5 Mbyte and the current processorspeeds approaching 1 GHz are fast enough to render reconstructions using either filter backprojectionor optimized iterative algorithms in an acceptable time (less than 10 minutes).

Typical performance specifications for SPECT imaging systems are summarized in Table 27.3. Aswith conventional planar imaging, the scintillation cameras, and the associated collimation are theprimary factors affecting the performance. SPECT spatial resolution is nearly isotropic with anFWHM of 8 to 10 mm for brain imaging where the detectors can get close to the radioactive source.

TABLE 27.3 SPECT System Performance

27.3.1 SPECT Image Reconstruction

27.3.2 SPECT System Performance





The spatial resolution degrades to 12 to 18 mm for body imaging because the detectors can not bepositioned as close. The components of SPECT spatial resolution and their relative importance can beidentified from the equation shown below:

As before, R int and Rcol represent the intrinsic and collimator resolution components. Rfilter is theFWHM of the smoothing kernel required to yield an acceptable reconstruction. The intrinsic spatialresolution is the least important factor in this calculation, since it is usually a factor of 2 or moresmaller than the other components. The trade-off between spatial resolution and count sensitivity isexplicit in this equation. Decreasing Rcol to improve spatial resolution will often require Rfilter tobecome larger to compensate for increased noise.

27.3.3 SPECT/PET Hybrid Systems

The primary reason for the success of nuclear medicine imaging is the availability ofradiopharmaceuticals that provide crucial diagnostic information. For cancer diagnosis and followup,18F flourodeoxyglucose (18F FDG) is an exquisite imaging agent for a wide variety of malignanciesincluding lung, colon, breast, melanoma, and lymphoma. Because 18F is a positron emitter that yieldsvery high energy x-rays (511 keV) when the positron combines with a free electron, it can not beimaged well on conventional SPECT systems. The thin NaI(Tl) crystals have a low efficiency fordetection at this energy (less than 10 percent). Also, the collimators designed for the 511-keV photonshave low count sensitivity and poor spatial resolution. The collimation can be eliminated if coinci-dence detection is used. Annihilation photons from positrons are always colinear. This feature can beexploited to count only events that are simultaneously detected by opposed detectors. Two opposedscintillation cameras with their collimators removed and the addition of coincidence electronics willturn a SPECT system into a PET tomograph (Fig. 27.16).34,35 The efficiency for coincidence detectionequals the product of the individual efficiencies, so that the coincidence efficiency is about 1 percentfor conventional scintillation cameras. This is actually much higher than the efficiency with collima-tors. Coincidence efficiency can be improved by using thicker NaI(Tl) crystals, and all the vendorshave done this. However, thicker crystals degrade the intrinsic spatial resolution when the scintillation

FIGURE 27.16 Two or more opposed scintillation cameras can be used as co-incidence detection systems for PET imaging. Valid events are establishedwhen the two detectors record events within the 10- to 15-ns timing window.Graded filters and lead septa are placed in front of the detectors to limit scat-tered radiation. (Courtesy of GE Medical Systems.)





cameras are used for conventional (i.e., collimated) studies. A solution to this problem, calledStarBrite, has been recently introduced. Thick NaI(Tl) crystals (25 mm) have slots machined in thePMT side of the detector to prevent the scintillation light from scattering throughout the crystal tomaintain good spatial resolution. There are other concerns, primarily the count rate capability. Withmost SPECT imaging studies, there are essentially no count rate losses resulting from the time it takesto process each event. The amount of radioactivity that can be safely administered and the lowsensitivity of the collimation keep the count rate in the range that the electronics can easily handle.However, when the collimators are removed for coincidence imaging, the NaI(Tl) crystals are ex-posed to count rates above the capacity of the conventional electronics. Much effort over the pastdecade has been devoted to increasing the count rate capability of the scintillation cameras. In theearly 1990s, the maximum observed counting rate for a scintillation camera was in the 200,000 to400,000 count/second range. This rate has been extended to over 1,000,000 counts/second byshortening the integration time on the pulses and implementing active baseline restoration. Becausethe light is proportional to the energy deposited in the crystal, one can shorten the pulse integrationtime without extreme degradation. Even with this substantial improvement in count rate, the maxi-mum activity the system can handle is about 3 mCi. Typical performance values for a coincidencescintillation camera system are given in Table 27.4.

In addition to specialized electronics, other measures have been taken to help reduce the countrate burden. One example of these is a graded absorber placed in front of each detector to helpreduce the scattered radiation component.36 Because scattered radiation has lower energy than theunscattered photons, low-energy scatter will be preferentially absorbed by lead, since thephotoelectric cross section is inversely proportional to the cube of the gamma ray energy. If only leadis used, the lead characteristic x-ray resulting from the absorption will be detected. Additional layersof tin, copper, and aluminum will absorb the respective characteristic x-rays. This graded filter causesa significant reduction in the low-energy scattered radiation and, since the scintillation camerasprocess every event, also reduces the overall count rate burden.

Even though the thin NaI(Tl) crystals have low intrinsic efficiency, the uncollimated detectorspresent a large solid angle to the annihilation photons. Maximum sensitivity is achieved when allcoincidences, even those at large angles, are accepted. This makes the camera sensitivity extremelysensitive to the source location. Sources located near the central axis of the detectors have a largesolid angle, while those at the periphery can interact only with a very small portion of the detectors.In addition, the scatter component increases to well over 50 percent when large coincidenceacceptance angles are used. This problem has been addressed by using lead slits aligned perpendicularto the axial direction of the system to restrict the angular extent of the coincidences. While thisreduces the sensitivity of the imaging system, it also reduces the scatter component to less than 30percent and limits the solid angle variability.

The intrinsic spatial resolution of the hybrid systems is comparable to that of the dedicated PETsystems with a FWHM of 4 to 5 mm. However, the count sensitivity is at least an order of magnitudelower. This, along with the maximum count rate constraint, guarantees that the coincidence cameradata will be very count poor and therefore require substantial low-pass filtering when reconstructed.As a result the quality of the reconstructed images is perceptibly worse than the dedicated PETimages. In head-to-head comparisons, it has been found that the hybrid systems perform well on

TABLE 27.4 SPECT/PET Hybrid System Performance





tumors greater than 2 cm in diameter located in the lung.3739 Tumors smaller than these and thoselocated in high-background areas are detected with a much lower sensitivity. These results areimportant since they provide a guide where the application of the coincidence camera will be useful.

In a conventional scintillation camera, Anger logic is used to determine the location of aninteraction. The tacit assumption in this approach is that only one event is occurring at a time. At thehigh count rate encountered in coincidence imaging, multiple interactions are likely, and when thisoccurs, the events are improperly located somewhere between the two true locations. Improvedalgorithms have been developed that can identify multiple hits and that use a maximum likelihoodcalculation to correctly determine event locations.

The projection data collected by the coincidence cameras require correction for randomcoincidences, scatter, and attenuation if accurate tomographic images are to be obtained. Random oraccidental coincidences occur when two unrelated photons are detected. These random coincidencesincrease rapidly with the count rate and give rise to a nondescript background that has to besubtracted from the projections. Typically randoms are either monitored in a separate time window orare calculated from the singles count rate and are subtracted. Scatter correction is sometimes ignoredin coincidence PET, or maybe monitored by a separate energy window and subtracted as discussedfor SPECT imaging.

Accurate reconstruction of PET data requires correction for attenuation since the degree ofattenuation for coincidence imaging is very high, approaching values of 100 or more. Attenuationcorrection requires information about the transmission of the annihilation radiation through the bodyat the coincidence lines of response. When attenuation correction is ignored, severe artifacts are seenin the reconstructed images. As with SPECT imaging, a separate transmission study with an externalsource, typically Cs-137, must be acquired to provide the attenuation map used in the correction.

SPECT imaging is expected to play a continuing important role in medical imaging. Future improve-ments in SPECT instrumentation are likely to include new detectors and collimation schemes. Thecoincidence scintillation cameras will also continue their evolution with the addition of more camerasand multidetector levels optimized for SPECT and coincidence imaging. Reconstruction algorithmswill evolve as new techniques are developed and as the performance of the computer expands. Inspite of the importance of instrumentation, the primary motivating factor in SPECT imaging willcontinue to be the creation and implementation of new radiopharmaceuticals. While SPECT willcontinue to be highly utilized for myocardial perfusion imaging, SPECT use in tumor imaging willprobably experience the largest growth. Applications will include treatment planning for internalradiation therapy as well as diagnostic studies.

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27.4 SUMMARY

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27 nuclear medicine imaging instrumentation

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