evaluation of targeting errors in ultrasound-assisted radiotherapy

Ultrasound in Med. & Biol., Vol. 34, No. 12, pp. 1944–1956, 2008Copyright © 2008 World Federation for Ultrasound in Medicine & Biology

Printed in the USA. All rights reserved0301-5629/08/$–see front matter

doi:10.1016/j.ultrasmedbio.2008.06.001

● Original Contribution

EVALUATION OF TARGETING ERRORS IN ULTRASOUND-ASSISTEDRADIOTHERAPY

MICHAEL WANG,* ROBERT ROHLING,* CHERYL DUZENLI,† BRENDA CLARK‡ and

NECULAI ARCHIP§

*Department of Electrical and Computer Engineering, University of British Columbia, Vancouver; †Department ofMedical Physics, British Columbia Cancer Agency, Vancouver; ‡Department of Medical Physics, Ottawa HospitalRegional Cancer Center, Ottawa, Canada; and §Surgical Planning Laboratory, Department of Radiology, Brigham

and Women’s Hospital, Harvard Medical School, Boston, MA, USA

(Received 26 January 2008; revised 14 May 2008; in final form 2 June 2008)

Abstract—A method for validating the start-to-end accuracy of a 3-D ultrasound (US)-based patient positioningsystem for radiotherapy is described. A radiosensitive polymer gel is used to record the actual dose delivered toa rigid phantom after being positioned using 3-D US guidance. Comparison of the delivered dose with thetreatment plan allows accuracy of the entire radiotherapy treatment process, from simulation to 3-D USguidance, and finally delivery of radiation, to be evaluated. The 3-D US patient positioning system has a numberof features for achieving high accuracy and reducing operator dependence. These include using tracked 3-D USscans of the target anatomy acquired using a dedicated 3-D ultrasound probe during both the simulation andtreatment sessions, automatic 3-D US-to-US registration and use of infrared LED (IRED) markers of the opticalposition-sensing system for registering simulation computed tomography to US data. The mean target localiza-tion accuracy of this system was 2.5 mm for four target locations inside the phantom, compared with 1.6 mmobtained using the conventional patient positioning method of laser alignment. Because the phantom is rigid, thisrepresents the best possible set-up accuracy of the system. Thus, these results suggest that 3-D US-based targetlocalization is practically feasible and potentially capable of increasing the accuracy of patient positioning forradiotherapy in sites where day-to-day organ shifts are greater than 1 mm in magnitude. (E-mail:[email protected]) © 2008 World Federation for Ultrasound in Medicine & Biology.

Key Words: Image-guided radiation therapy, Image to physical space registration, Localization, Ultrasound
guidance.
INTRODUCTION

The aim of radiotherapy is to deliver a lethal dose ofradiation to the tumor while minimizing the exposure ofsurrounding healthy tissues to radiation. With the adventof conformal treatment techniques such as intensity-modulated radiation therapy (IMRT), 3-D dose distribu-tions that accurately match the shape of the target vol-ume can now be achieved, facilitating increases in radi-ation dose to tumor sites and reductions in treatmentmargins. However, errors in patient positioning at thetime of treatment can still cause geometric miss of thetarget volume and irradiation of greater amounts ofhealthy tissue (Langen and Jones 2001). Thus, image-guidance methods for localizing the target inside the

Address correspondence to: Michael Wang, 5500-2332 MainMall, Vancouver BC, V6T 1Z4 Canada. E-mail: [email protected]

1944

treatment room have become the subject of considerableresearch efforts (Herman 2005; van Herk et al. 2004;Jaffray 2005; Mackie et al. 2003; Langen et al. 2005).

Ultrasound (US) imaging has a number of desirablefeatures that make its use in radiotherapy applicationsappealing. These include being relatively inexpensive,nonharmful to the patient, being able to visualize softtissue and acquire 3-D volumetric representations of theanatomy. Indeed, a number of commercial systems, in-cluding the BAT (Nomos, Chatsworth, CA), SonArray(Varian, Palo Alto, CA, USA) and Restitu (ResonantMedical, Montreal, Canada), which use US to verify theposition of the tumor before the commencement of dailytreatment, are now in clinical use (Lattanzi et al. 1999;Meeks et al. 2003; Cury et al. 2004). In addition, tech-niques for integrating B-mode US information in radia-tion treatment planning have been investigated (Wein et
al. 2007).
mailto:[email protected]

mailto:[email protected]

Targeting errors in US-assisted radiotherapy ● M. WANG et al. 1945

Although substantial improvements in patient posi-tioning accuracy have been reported for US-based targetlocalization systems, no attempts to characterize theiroverall start-to-end accuracy have been reported previ-ously in the literature. Tomé et al (2002) evaluated theaccuracy of the SonArray US patient positioning systemusing an array of passive infrared markers of the Polarisoptical sensor (NDI, Waterloo, Canada) attached to aphantom. Locations of targets within the phantom asmeasured by computed tomography (CT) in the treat-ment room were then compared with US-measured targetlocations. Drever and Hilts (2007) described a phantomfor quality assurance (QA) of the calibration accuracybetween treatment room lasers and US coordinates forthe Restitu system.

In this paper, a method is presented for evaluatingthe complete start-to-end accuracy of an US patient po-sitioning system. Because of the additional coordinatetransformations and calibrations required when addingUS-assisted positioning, a small addition to the overallerror can be expected compared with the current estab-lished method of aligning skin markers placed on thepatient with room lasers. The goal of this study is tocharacterize the magnitude of this error under ideal con-ditions for a new 3-D US-based scheme for target local-ization. Unlike previous studies, which measured a sub-set of errors in the patient positioning process (Tomé etal. 2002; Drever and Hilts. 2007), the sources of error atall steps are taken into account herein by measuringexplicitly where the radiation was delivered in a dose-sensitive phantom, and compared with the planned targetlocation. This novel validation approach allows the loopto be closed on the entire treatment process. Because thisphantom is rigid, a lower bound on the target localizationerror for the system can be established and comparedwith expected interfractional organ motion to evaluatewhether US guidance is beneficial. In addition, US po-sitioning results are compared with using room lasersalone and optical tracking as gold standards.

SYSTEM DESCRIPTION

The US guidance method used in this study is mostsimilar to the Restitu system, as intramodality US-to-USregistration is performed to compare the patient anatomyduring treatment and CT simulation. In contrast to theBAT and SonArray systems, which attempt to performCT-to-US registration based on image content, intramo-dality registration is a more tractable problem and doesnot suffer from the variable appearance of anatomicalfeatures in CT and US (Narayana et al. 1997; Hoffelt etal. 2003). The tradeoff is the need to perform an addi-tional US scan at the time of CT simulation. In addition,
to minimize reliance on operator input, and hence reduce
interoperator variability, which can significantly affectexisting systems (Langen et al. 2003), automatic pixelintensity–based image registration is used. Unlike exist-ing systems, which all use freehand US (Rohling et al.1999) to acquire 3-D patient data, a dedicated 3-D probeis used for this study. These offer a number of advan-tages, including minimizing the number of image acqui-sitions needed to obtain a 3-D perspective of the anatomy(one compared to multiple 2-D images), ease of use(removes the need to manually sweep a 2-D probe) andpotentially improved accuracy in the reconstructed 3-Dimage (Treece et al. 2002).

The 3-D US-patient positioning system consists of theGE Voluson 730 Expert US machine (GE Healthcare,Waukesha, WI, USA) interfaced with the OPTOTRAK3020 position sensor (NDI, Waterloo, Canada). The convexarray RAB4-8P transducer was used to acquire US data.This transducer has an internal mechanism that can physi-cally pivot the transducer elements, allowing an entire seriesof 2-D image slices in a fanlike pattern to be acquired whilethe probe remains stationary. This equipment is used forboth the “simulation” session, where a CT scan of thepatient is taken to facilitate treatment planning, and the“treatment” session, where radiation is delivered accordingto the plan. During the simulation procedure, infrared LED(IRED) markers of the OPTOTRAK are placed on thepatient’s skin before commencement of the CT scan so thatthey appear in the CT image as fiducial markers. Immedi-ately after CT imaging, a US volume of the same anatomyat the treatment site is acquired. During US imaging, boththe US probe and patient positions are tracked by theOPTOTRAK using IREDs attached to the probe and pa-tient. After the simulation procedure, the IRED markers areremoved from the patient. The acquired CT and US imagesare registered off-line using OPTOTRAK measurementsand the IRED fiducial locations in CT. This key step en-ables treatment planning to proceed conventionally usingCT data, while allowing the prescribed dose plan to betransferred to the US image to facilitate target localizationduring treatment by comparison of simulation and treatmentUS data.

In the treatment session, the patient is set up asusual using skin marks and alignment lasers. Because ofpossible errors, such as internal day-to-day organ shiftswith respect to the skin markers, the target may not beexactly at the expected location. Thus, another US vol-ume of the target anatomy is taken at approximately thesame viewpoint as the simulation US, while the probeposition inside the treatment room is measured byOPTOTRAK. Finally, this US is automatically rigidlyregistered to the US taken during simulation. This imageregistration step allows the location of the target, which
is known relative to the simulation US, to be found inside

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wn in

1946 Ultrasound in Medicine and Biology Volume 34, Number 12, 2008

the treatment room with respect to the linear accelerator(LINAC).

The series of coordinate transformations that areused by the proposed system to localize the treatmenttarget with respect to the LINAC are shown in Fig. 1 andsummarized below.

PatTCT: The CT-to-patient transformation. This iscalculated using point-based registration using corre-sponding IRED locations on the patient and in the CTscan (Horn 1987).

OTPat: The patient-to-OPTOTRAK transformationgiven by OPTOTRAK measured IRED locations on thepatient.

PTO: The OPTOTRAK-to-US probe transformationgiven by OPTOTRAK measured IRED locations on theprobe.

UTP: The US probe-to-US image transformation.This is constant for a given probe and can be predeter-mined in an off-line calibration procedure (Poon andRohling 2005).

U�TU: The transformation between simulation andtreatment US images of the target anatomy. This is to becalculated using an automatic image registration algo-rithm.

P�TU�: The US image to probe transformation for thetreatment session. If the same probe is used as during

Fig. 1. The series of coordinate transformations from tproposed system for target localization. Calibration paraif the same probe is used for US imaging, UTP � (P�TU

before each treatment fraction are sho

simulation, this is the inverse of UTP listed previously.

O�TP�: The US probe position inside the treatmentroom measured by the OPTOTRAK.

LTO�: The OPTOTRAK to LINAC transformation.This can be determined by measuring alignment laserpositions in the treatment room using the OPTOTRAK.

The matrix T represents a 6-degree-of-freedomrigid-body coordinate transformation. Thus, to local-ize a target location CTxiso defined in the simulationCT coordinates with respect to the LINAC at the timeof treatment, the following chain of coordinate trans-forms are used by our system:

XLiSO � LTO�O�

TP�P�

TU�U�

TUUTP

PTOOTPAT

PATTCTCTXiSO.

(1)

In summary, the main differences between the tar-get localization method used in this study and currentexisting systems, including the Restitu, are the use ofdedicated 3-D US probes for image acquisition, auto-matic pixel-intensity-based rigid registration of simula-tion and treatment US images and optical tracking ofboth patient and US probes during the simulation proce-dure.

METHODS

Two validation studies were performed to ascertain

ulation CT to the LINAC coordinates when using theknown a priori are shown by dashed arrows. Note thathe data collected during the simulation procedure and

the left and right panels, respectively.

he simmeters

)�1. T

the accuracy of the proposed 3-D US positioning system.


First, the accuracy of the method for registering simula-tion CT and US images (left panel of Fig. 1) was tested.This step is a key component of the target localizationprocess because it allows transfer of the dose plan de-fined in CT to US after the simulation and treatmentplanning processes. Then, the entire system accuracy isevaluated, including errors at the time of dose delivery.An idealized, rigid phantom is used in both experimentsto establish the lower bound accuracy of the system.

Simulation CT and US registration accuracyCT and US images of the phantom shown in Fig. 2

were registered using the series of transformationsshown in the left panel of Fig. 1. This phantom containswire crosses to facilitate US probe calibration and speed-of-sound verification, and eight IREDs attached outsideas fiducial markers. It is filled with an US tissue–equiva-

Fig. 2. The phantom used in the CT-to-US registrationexperiment.

lent material consisting of (by weight) 3% high gel–

strength agar (Sigma-Aldrich, St. Louis, MO, USA),88.4% distilled water (Anachemia, Montreal, Canada)and 8.6% n-propanol (Fisher Scientific, Hampton, NH,USA) to regulate its acoustic velocity to 1540 ms�1

(Burlew et al. 1980). Embedded within the gel were 43aluminum ball bearings (BBs) of 1 mm diameter to act aspoint features. These ball bearings were pushed into thegel with a needle just before it coalesced and wererandomly distributed throughout the phantom.

The speed-of-sound in the phantom was verified tobe within 0.5% of 1540 ms�1 by comparing the physicaldistance between the center of two vertical wire crosseswith their observed distance in an US image. Therefore,geometric distortions in US images of the phantom be-cause of speed-of-sound error are expected to be insig-nificant. The calibration matrix UTP of our US probe wasdetermined using the methods described by Poon andRohling (2005). Briefly, the IXI-shaped wire crosses inthe phantom were scanned while tracking both the probeand phantom positions with the OPTOTRAK. The re-sulting US image was rigidly registered to a virtualmodel of the wires generated using their locations, withrespect to the IRED markers on the phantom, which hadbeen measured precisely during its construction. Thisregistration step was performed using the Amira 3Dvisualization toolkit (TGS, San Diego, CA, USA) bymaximizing the image correlation coefficient (Fitzpatrickand Sonka 2000). The result of this registration (UTPhan)and the OPTOTRAK measurements of the phantom andprobe positions (PhanTO and PTO, respectively) were thenused to calculate UTP:

UTP � UTPhanPhanTO�PTO��1 (2)

CT imaging was performed at the British ColumbiaCancer Agency (BCCA) with a PQ5000 CT simulator(Picker International, Cleveland, OH, USA). On the sameday, the phantom was imaged using 3-D US, as shown inFig. 2. Nine US volumes were required to image the entirephantom. The required registration parameters, OTPat andPTO, were measured using OPTOTRAK for each volume.The remaining unknown PatTCT was calculated by least-squares minimization of the distance between the measuredIRED locations on the phantom and their correspondinglocations in the CT scan (Horn 1987). IRED locations in CTwere determined by finding their image centroid and apply-ing a calibrated offset (1.47 mm) along their axis of sym-metry to account for the offset of the IRED inside itshousing. The BB locations in CT were identified by calcu-lating the image centroid in a surrounding region-of-inter-est, as shown in Fig. 3. The BB locations in US weresimilarly identified after thresholding was used to removethe comet-tail artifact from US. Because the BB occupies
many voxels in both the CT and US images, their target


localization error (TLE) was considerably smaller than thedimension of a voxel (0.68 � 0.68 � 1.5 mm for CT, 0.55

Fig. 3. The phantom BB locations in CT and US.

mm for US). The difference in corresponding BB locations

in registered CT and US images was defined as the targetregistration error (TRE) and used to evaluate the accuracyof the registration. Because US data was acquired with acurvilinear array probe, for which beam width, line densityand lateral resolution decreases below the image center, BBlocations near the center of the volume were selected astarget locations for comparison.

Target localization accuracy of entire systemTo determine the overall accuracy of the proposed

system, the entire process of radiation therapy, fromsimulation to 3-D US guidance and, finally, irradiation,was performed on the phantom shown in Fig. 4. Thisphantom consists of a large box containing the same UStissue equivalent material as described previously (Bur-lew et al. 1980) and three embedded balloons filled withthe same material plus a high concentration of celluloseto act as hyperechogenic features in US. The lid of thebox was opened to allow access by the US probe. EightIREDs were taped to the side wall of the box as CTfiducials. In addition, four polypropylene bottles contain-ing a radiosensitive polymer gel were rigidly secured tothe wooden base with cable ties. The polymer gel ismade up of a gelatin matrix doped with a monomer thatpolymerize and cross-link when exposed to ionizing ra-diation, thereby inducing a localized change in the gel’sdensity. Details of its properties and manufacture havebeen previously described by Jirasek et al. (2006). Byusing the polymer gel to record the 3-D dose distributiondelivered during treatment, the actual dose received bythe phantom was compared with the planned dose, clos-ing the loop on the entire treatment process using 3-D USguidance.

Four target isocenter locations were chosen duringtreatment planning on the Eclipse system (Varian, PaloAlto, CA, USA), each near the center of one of the fourpolymer gel bottles on the phantom. Each isocenter wasgiven the three-field dose plan shown in Fig. 4. The6-MV beam from the Varian 21EX LINAC was chosento deliver the radiation with square fields 30 mm wide. Adose of 15Gy was assigned to each isocenter.

On the day of treatment delivery, the OPTOTRAKwas positioned inside the treatment unit against thewall at the inferior end of the couch, and the Voluson730 was stationed laterally beside the couch. To cal-ibrate OPTOTRAK to the LINAC machine isocenter(the transformation LTO� in Fig. 1), the patient align-ment lasers were used. The lateral and overhead lasercross-hair locations on the sides of a box positioned onthe couch were touched by a calibrated stylus whileOPTOTRAK measured its tip location using IREDsattached to the stylus base. The measured stylus tipposition corresponds to a point lying on one of the
orthogonal axes of the LINAC coordinate system. By

ions o


shifting the couch, and thereby the points of interceptbetween the surface of the box and lasers, 10 differentpoints along the left–right and anterior–posterior axisof the LINAC coordinate system were sampled usingthe stylus. These points were projected onto the least-squares fit plane, and the best fit orthogonal axis, alongwith the origin (isocenter location), were solved usingconstrained least squares.

In addition to using 3-D US localization to positionthe phantom for irradiation of the four targets, three otherindependent methods were also used to localize the tar-get in the treatment room before irradiation to allowcomparison of these approaches. For each target, theprocedures performed before administering the planneddose are described next.

Laser positioning. The conventional method ofaligning set-up marks placed during the simulation pro-cedure to the guidance lasers inside the treatment roomwas first used to position the phantom. Couch shifts wereused to implement the displacement from the referencepoints drawn on the phantom during simulation to theindividual target locations given by the treatment plan-ning system. After laser positioning, an artificial set-uperror was introduced for each target by shifting the couchusing the displacements shown in Table 1. These couchshifts were chosen to investigate errors along each of thecoordinate axes separately as well as combined. The5-mm level corresponds to the magnitude of interfrac-tional organ motion typically encountered in the prostate(Langen and Jones 2001).

3-D US localization. To reposition the phantom inthe correct treatment position, the 3-D US guidancesystem was used. A 3-D US scan with views of theballoons inside the box on the phantom was acquiredwhile tracking the probe position with OPTOTRAK(O�TP). This image was rigidly registered to a similar

Fig. 4. Phantom used to test 3-D US target localizatarrangement and isodose distribution surrounding one o

orthogonal cross-sect

image taken previously during the simulation procedure

(U�TU) by maximizing the image correlation coefficientusing Amira (TGS, San Diego, CA, USA). The simula-tion 3-D US had been registered before treatment to thesimulation CT using methods described previously.These results, along with calibration parameters LTO� andP�TU� were substituted into eqn (1) to find the shiftsrequired to bring the target into alignment with theLINAC isocenter. These shifts are displayed in Table 1.Note that it is possible to calculate the error from theother target localization methods by subtracting the shiftfrom 3-D US and adding the shift from another method.

Freehand ultrasound localization. Before the cor-rection shifts in Table 1 were implemented using thecouch, volumetric US data using the same probe, butwith a freehand scanning technique, were collected byacquiring a series of 2-D image slices while manuallysweeping and tracking the probe position over the phan-tom surface. The purpose of this test is to compare acombination of freehand scanning and contour-basedUS-to-US registration, currently used by all existing US

curacy. The right-hand side shows the planned fieldtarget isocenters inside a gel bottle, overlaid on top off the simulation CT.

Table 1. The couch shifts used to simulate a set-up errorafter the phantom is positioned using patient alignment lasers

is given as the “simulated error”

Target

Simulated Error(mm) Correction Shift (mm)

x y z x y z

1 5 0 0 �5.7 0.4 �1.22 0 0 5 0.4 1.4 �7.23 0 5 0 �2.1 �6.8 �2.24 5 5 5 �5.3 �5.4 �7.6

The shifts required to move the phantom back to its correct positionas determined by 3-D US localization is shown as the “correctionshift.” All vectors are given in the LINAC coordinate system (y couch

ion acf the

axis of rotation, z gantry axis of rotation, and x lateral axis). Note thatthe couch moves in one millimeter increments.

nstruct


guidance systems, with a combination of mechanical-sweep scanning and automatic intensity-based US-to-USregistration, as was used in this study. This allows thelocalization accuracy of the two approaches to be di-rectly compared and provides further insight into sourcesof error.

The freehand data was not used for positioning thephantom but was stored for off-line analysis. The con-tours of the balloons in each slice of the dataset wereextracted by thresholding and reconstructed into surfacepoints as shown in Fig. 5 using OPTOTRAK measure-ments of the probe position during scanning and theprobe calibration matrix. The freehand US acquired dur-ing treatment was rigidly registered to a similar freehandvolume acquired during simulation (U�TU) by using theICP algorithm (Besl and McKay 1992) to align thesurface points extracted from each. No manual interven-tion was required to initialize the registration, and thealgorithm converged within 30 iterations in each case.The simulation freehand US was registered to the simu-lation CT using the same technique used for 3-D USimages. Finally, the target location was calculated usingthese results and eqn (1).

OPTOTRAK localization. As a final check of thetarget location before the phantom was irradiated, thephantom position (O�TPat) was directly measured byOPTOTRAK using the IREDs left intact from simula-tion. The following series of transformations was then

Fig. 5. Contours of the balloons extracted from slicesreco

used to localize the target:

LXiSO � LXO�O�

XPATPATXCT

CTXiSO. (3)

where LTO� and PatTCT are as defined previously andCTxiso is the isocenter of the radiation plan with respect tothe CT image. The target location determined using eqn(3) was recorded for offline comparison with 3-D USlocalization.

Polymer gel response readoutTo extract dose information from the irradiated gel

bottles, radiation-induced density change in the gel wasdetected by CT imaging as subtle changes in itsHounsfield units (HU). CT imaging was performed atCanada Diagnostic Centers (Vancouver, Canada) withthe Toshiba Aquilion 64 multislice CT (Toshiba, Tokyo,Japan). To optimize image quality, a CT imaging proto-col following the guidelines set by Hilts et al. (2005) wasused. To prepare for imaging, each gel bottle was re-moved one at a time from the phantom and placed on aStyrofoam cradle taped securely to the CT table. Thisensured that all the bottles were placed in the sameposition within the CT bore during CT scanning. Eachgel bottle was scanned individually and oriented so itsaxis of symmetry was in approximate alignment with theCT gantry axis of rotation to reduce the presence ofimaging artifacts. The CT imaging parameters were0.5-mm slice thickness, 0.335-mm pixel size, 135-kVtube voltage, 350-mA tube current and 0.5-s scan time.

eehand US volume of the phantom in Fig. 4 after 3-Dion.

of a fr

To improve image signal-to-noise ratio, each bottle was


scanned five times and the images were averaged. Inaddition to the four irradiated bottles, a bottle of nonir-radiated gel was also scanned using the same imagingparameters. The average image of the nonirradiated bot-tle was subtracted from the average image of each irra-diated bottle to perform background subtraction andform a difference image displaying the change in gel HUafter irradiation.

The difference image near one of the targets isshown in Fig. 6. Aside from the bright pixels around theedges of the bottle because of a slight misalignment ofthe irradiated and nonirradiated bottles used to calculatethe difference image, the bright cluster of pixels corre-sponding to the location where the gel received thehighest dose can clearly be seen.

Fig. 6. The difference image near one of the targets showingchanges in polymer gel HU after irradiation. The data is win-dowed from 0–10 HU. The white outlines around the edge ofthe bottle are the result of slight misregistration of the bottles

used to generate the image.

To evaluate the accuracy of the target localization

methods tested in this study, the difference image foreach gel bottle was registered with the simulation CTusing BB fiducial markers placed on the bottles. Thisallowed the planned dose geometry, defined in the sim-ulation CT coordinates, to be transferred to the differenceimage. Next, the actual irradiated isocenter location wasdetermined by registering the 3-D dose plan, given by thetreatment planning software as a 3-D matrix of dosevalues, with the difference image by maximizing theimage correlation coefficient in Amira as was done pre-viously for US images. Jirasek et al. (2006) have shownpreviously that the gel formulation used in this experi-ment has an approximate linear dose response of 0.32HUGy�1. In addition, the planned isocenter locationdetermined previously was used to initialize the algo-rithm. Because this should be close to the irradiatedisocenter location, the possibility of misregistration viaconvergence to the wrong local minimum is unlikely.

The 3-D Euclidean distance between the plannedand irradiated isocenter locations was taken to be thetarget localization error. Because the 3-D US guidancewas used to finalize the phantom position before it wasirradiated, the target localization errors determined fromthe gel directly correspond to the overall accuracy of 3-DUS assisted positioning. The accuracy of the other meth-ods tested were derived by comparing the correspondingtarget localization results with that obtained with 3-DUS. For example, the difference in the target locationdetermined by 3-D US and laser positioning can befound by adding the simulated error and the correctionshifts shown in Table 1. This difference can then beapplied to the irradiated target location to find the posi-tioning error that would have occurred from laser align-ment. The target localization error for the freehand USand OPTOTRAK localization methods were obtainedsimilarly.

RESULTS AND DISCUSSION

Simulation CT and US registration accuracyThe target registration error (TRE) of 65 BBs ob-

served in registered CT and US volumes were �0.50 �0.53 mm, �0.12 � 0.70 mm and 0.23 � 0.33 mm (mean� standard deviation) in each orthogonal direction, in-dicating the small level of systematic errors in this por-tion of our system. The overall root-mean-square (RMS)3-D TRE was 1.09 mm. No noticeable detrimental ef-fects on the performance of the IREDs were found afterexposure to X-rays, supporting their continued use as CTfiducials.

The major sources of error in these registrationsteps include calculating the IRED locations in CT andthe US probe calibration matrix UTP. The effect of errors
in determining CT IRED locations on the accuracy of

tions arnits ar


PatTCT can be reduced if this error is nonsystematic anda sufficient number of IREDs are used to establish pointcorrespondences. The RMS fiducial registration error(FRE) for the 8 IREDs was found to be 0.88 mm.Although the OPTOTRAK IREDs were found to besuitable as fiducial markers in CT for this study, it maybe desirable to attach them to a larger base object that isclearly defined in CT to improve the accuracy of deter-mining their locations.

The accuracy of the probe calibration process usedhere has been evaluated previously by Poon and Rohling(2005) using the same US system and probe. The RMSpoint accuracy was found to be 2.15 mm. However, theRMS 3 D TRE for this experiment was 1.09 mm, includ-ing errors from PatTCT as well. This discrepancy may bebecause of the different materials used as targets and thebackground material by Poon and Rohling (2005).

In a similar study, Tomé et al. (2002) reported UStarget localization errors of 0.2 � 0.7 mm, 0.9 � 0.6 mmand 0.6 � 1.0 mm along each orthogonal axis for theSonArray system when compared with target locations inthe planning CT. These results are comparable to the CTand US TRE in this study.

Target localization accuracy of entire systemThe accuracy of all four methods tested for localizing

the four targets are summarized in Table 2. The positioningresults for one target is also shown in Fig. 7. The targetlocalization accuracy of the optical guidance methods tested(i.e., laser and OPTOTRAK) were better overall than USguidance with the proposed system using either 3 D orfreehand data. As mentioned in the Introduction, this is notsurprising, because the number of error sources for the laserand OPTOTRAK methods are significantly less than USguidance, as can be seen from Table 3. Nevertheless, theextra sources of inaccuracy involved with US target local-ization only amounted to approximately 1 mm of additionalerror in practice compared with optical guidance. In addi-tion, the positioning accuracy of the proposed system using

Table 2. Target localization error of the ph

Target

Laser positioning 3-D US

x y z 3-D x y z

1 0.0 0.9 0.4 1.0 �0.8 1.2 �0.82 1.0 �1.9 0.9 2.3 1.5 �0.4 �1.23 0.7 �0.7 �0.1 1.0 �1.4 �2.4 �2.24 1.2 �1.5 0.6 2.0 0.9 �1.9 �2.1Mean — — — 1.6 — — —

The error vectors between the planned and irradiated isocenter locaaxis of rotation, and x lateral axis), as well as their 3-D magnitudes. U

3 D ultrasound data was comparable to using freehand data

and contour-based registration. The major sources of errorand their estimated magnitudes for each target localizationtechnique tested are summarized in Table 3 and discussednext.

Laser positioning. The accuracy of the conventionallaser positioning method is limited by the accuracy of thealignment lasers, and couch movement used to imple-ment the displacement from the reference marks on thephantom to the actual target location. Quality assuranceprocedures at the BCCA ensure that the alignment lasersare accurate to within 2 mm in each axis in marking thetrue LINAC isocenter location. In addition, they are 0.5mm wide. Therefore, in the worst-case scenario, a max-imum error of �3 � 2.252 � 3.9 mm can occur with

the laser indicators. In practice, however, both the laserand couch errors were considerably smaller than theirmaximum possible values, as shown by the small overalllaser error magnitudes in Table 2

Couch error. The motor-driven couch is only ableto move in 1-mm increments. Thus, a maximum error of

�3 � 0.52 � 0.87 mm in the phantom position be-

cause of finite couch precision is possible.UTCT error. The method used for registering simu-

lation CT and US (UTCT) was evaluated in the previoussection on a similar phantom, where a RMS TRE of 1.09mm was obtained. This level of accuracy can be expectedin this case as well.

P�TU� error. The probe calibration error during sim-ulation UTP is already included in UTCT error. However,its effect during treatment (P�TU�) must also be taken intoaccount. Because the overall RMS accuracy of UTCT wasfound to be 1.09 mm, error because of probe calibration(both UTP and P�TU�) can be expected to be less than thisfigure. The same reasoning applies to PatTCT.

U�TU error. The 3-D US-to-US image registrationalgorithm used in this study has been tested previously

positioning methods tested for all targets

Freehand US OPTOTRAK

x y z 3-D x y z 3-D

�0.8 �1.0 1.3 1.9 0.6 1.6 0.6 1.91.1 2.7 �2.5 3.8 1.1 �1.4 0.0 1.8

�0.6 �0.3 �3.5 3.5 0.3 �0.2 �1.8 1.8�0.7 �1.1 �1.6 2.1 0.7 �1.3 �1.0 1.8

— — — 2.8 — — — 1.8

e shown in the LINAC coordinates (y couch axis of rotation, z gantrye in mm.

antom

3-D

1.62.03.62.92.5

on a cross-wire phantom. The image registration re-

shown


sults were compared with a gold standard obtainedfrom OPTOTRAK measurements of the US probepositions during image acquisition and knowledge ofthe probe calibration parameter. A RMS registration

Fig. 7. Phantom positioning results for Target 2 are showirradiated isocenter locations are projected onto the slicplanned and irradiated 80% isodose contours are also sho3-D US-determined target locations are used to simulateshown in (b). This was also done to simulate the OPTO

results

accuracy within the order of the pixel size, 0.62 mm,

was achieved (Wang 2006). Because the phantomimaged in this experiment contains similar materialand idealized, well-defined features with high contrast,a comparable image registration accuracy can be

ne axial slice of the difference image. The planned andare marked using the cross and dot, respectively. The

ing white outlines. The difference between freehand andsitioning results for the freehand localization approachpositioning results shown in (c), and laser positioning

in (d).

n in oe andwn usthe po

TRAK

expected.


An additional benefit of our system is that the effectof systematic US imaging errors, such as inaccurateprobe calibration (UTP), speed-of-sound distortions of theimage geometry or even tissue deformation because ofapplied probe pressure, can be minimized if U�TU � I,where I is the identity matrix. This can be demonstratedby introducing a transformation matrix Terr representingthis error into eqn (1):

LXiSO

� LTO�O�

TP�P�

TU�(Terr)U�

TU(Terr)�1UTP

PTOOTPAT

PATTCTCTXiSO.

(4)

Even if U�TU � I, some error cancellation is possi-ble. In practice, treatment and simulation US images arelikely to be taken from similar viewpoints and containsimilar anatomical features. This, along with the use ofdedicated 3-D ultrasound probes, which removes theneed for the operator to manually sweep the probe acrossthe skin, promotes the reproducibility of US images usedfor registration and the likelihood of error cancellation.

Additional errors are introduced when registeringfreehand US data. These include inaccuracies in freehandreconstruction, US and OPTOTRAK data synchroniza-tion and contouring (Rohling and Gee 1996). Theseerrors cause variations in the reconstructed surface con-tours to occur from different freehand images (an exam-ple reconstruction was shown in Fig. 5). The variationsbetween contours reconstructed from the simulation andtreatment freehand images can be gauged by inspectingthe residual error between closest matching contourpoints after registration using the ICP algorithm. Forlocalization of the four targets in this experiment, themean RMS contour registration error was 1.29 mm. It isimportant to note that even with the presence of sucherrors, it is still possible for an accurate registration resultfor U�TU to be obtained, as long as the errors are suffi-ciently small and uniformly distributed in all directions

Table 3. Major sources of error for each target localizationmethod tested, along with their estimated magnitudes in

parentheses, if known

Sources of Error Laser 3-D US Freehand OPTOTRAK

Laser (�3.9 mm) ✓ ✓ ✓ ✓Couch (�0.87 mm) ✓ ✓ ✓ ✓UTCT (1.09 mm) ✓ ✓P�TU� (�1.09 mm) ✓ ✓PatTCT (�1.09 mm) ✓U�TU (� 0.62 mm) ✓ ✓

A tick indicates a contributing source of error. Note that PatTCT erroris included in the overall UTCT error for the 3-D and freehand USguidance methods.

over all the contour points. Thus, contour-based regis-

tration of freehand data was still capable of achieving asimilar accuracy as registration of 3-D US data in ourexperiment.

Phantom effects on experimental resultsWith the phantom used in this experiment, the effect

of any US-to-US image registration error on the accuracyof the target localization results may be amplified by thelever-arm effect. This is because the target locations arenot within the US image field-of-view, because the poly-mer gel bottles are adjacent to the box used for USscanning. The mean distance from the center of thesimulation US volume to the four targets within the gelbottles is 160 mm. Even with a registration error of 1°between the simulation and treatment US volumes, themisalignment at this distance away would be 2.8 mm. Inconventional use of our system, the target isocenterwould be directly imaged using US. However, this wasnot possible with the phantom in this experiment becauseit was impractical to incorporate the polymer gel into thebox used for US scanning. Nevertheless, because thetarget localization error was still relatively low for boththe 3-D and freehand US methods and well within theexpected upper bounds, it seems that this effect did nothave a significant impact on the overall results.

It is important to recognize that the phantom used inthis experiment represents an idealized patient, and thatthe positioning accuracy achieved by the various meth-ods tested represents lower bounds. The automatic pixelintensity–based 3-D US-to-US registration results ob-tained correspond to the best-case scenario for imageregistration, because there were many clear featurespresent in the US image. In addition, artifacts such asrefraction and speed-of-sound error were not present aswould likely be for a real patient. The presence of clearfeatures with high contrast also enables the extraction ofcontours from freehand US images of the phantom to beperformed automatically. In US images of real anatomy,contours may need to be drawn manually, which in-creases the likelihood of further errors.

The phantom is also incompressible, unlike realtissue. Thus, nonrigid organ motion was not present toimpair the performance of the image registration algo-rithm, which assumes a rigid transformation betweenimages. Because accurate registration of the anatomy isthe vital link between simulation and treatment stages,the presence of nonrigid organ motion could potentiallypose serious problems for image guidance techniques forradiotherapy.

A potentially significant source of error in practicefor freehand US is caused by tissue deformation fromprobe pressure as it is manually swept over the area ofinterest to be imaged (Artignan et al. 2004; McGahan et
al. 2004). If this deformation is nonuniform in different


slices of the freehand data acquired, then the locations offeatures in the reconstructed volume will appear to be outof alignment. Because the phantom imaged in this ex-periment is incompressible, this aspect of freehand UStarget localization is not tested.

It should be noted that the additional positioningerror associated with US guidance in this study is not anindication that US is not valuable in target localizationfor human subjects. Indeed, internal day-to-day organshifts and skin mark migration in real patients will in-troduce additional errors when using external markers toposition the patient via alignment to lasers, or opticaltracking by cameras. In contrast, US imaging can enablethe internal anatomy surrounding the target to be visual-ized directly, whereas optical guidance techniques canonly localize external fiducials on the surface of thepatient at some distance from the treatment target. Be-cause a rigid phantom is positioned in this experiment,the full advantage of US over laser and OPTOTRAKlocalization methods is not fully exploited. Hence, al-though more accurate target localization results wereobtained using optical guidance than 3-D US guidance inthis experiment, the presence of internal organ motion inreal patients means that the former will not necessarilybe more accurate in clinical practice.

Although it is impossible to simulate the entirebehavior of a real patient under clinical conditions usinga phantom, the study presented herein still providespertinent information that would otherwise be impossibleto obtain from patients. Currently, accurate verificationof 3-D dose distribution delivered to tissue is not possi-ble. Hence, the true irradiated target location can only bedetermined accurately from phantom studies.

Comparison to previous studiesStudies of target localization accuracy and quality

assurance of US patient positioning systems for radio-therapy have not been reported widely in the literature,although there has been a proliferation of these systemsin clinical usage. We previously compared our CT withUS registration results with those obtained by Tomé et al.(2002) using the SonArray US target localization system.In the same manuscript, US localization accuracy wascompared with optical guidance and were found to agreeto within 1 mm along each orthogonal axis.

The overall target localization accuracy of our 3-DUS positioning system, as shown in Table 2, was mea-sured to be slightly larger than that of Tomé et al. (2002).However, the experiment described herein does not useoptical guidance as the “ground truth” for target local-ization accuracy. Instead, the phantom is positioned likea real patient in the treatment room, with respect to anabsolute frame of reference (that of the LINAC iso-
center). Furthermore, the phantom is irradiated, and
changes in its material properties are compared with thetreatment plan to close the loop on the entire treatmentprocess to verify the accuracy of 3-D US positioning.

Thus, the entire process of radiotherapy treatment,including simulation imaging, treatment planning, pa-tient positioning at the time of treatment and irradiation,is simulated in this experiment. Errors such as couchmovement inaccuracy, inaccurate calibration of the op-tical camera system to the LINAC and accuracy of theLINAC itself in delivering the prescribed dose are in-cluded in the target localization error, giving a morerealistic measure of the overall expected treatment accu-racy. No previous study on image guidance systems forradiotherapy has attempted to show this type of start-to-end validation. The general experimental methodologiesconcerning the use of polymer gel dosimetry describedherein can be adopted for quality assurance testing ofother image guided radiotherapy systems as well.

CONCLUSION

The accuracy of a 3-D US system in positioning arigid phantom for irradiation of four target locations wasvalidated using polymer gel dosimetry. The mean posi-tioning error for the four targets was 2.5 mm. Despiteintroducing a number of extra sources of error into thetarget localization process, the overall accuracy of theproposed 3-D US patient positioning system was onlyapproximately 1 mm worse than the optical phantompositioning methods tested, including the conventionalradiotherapy technique of using alignment lasers andset-up marks. In clinical practice, the magnitude of in-terfractional organ motion at most treatment sites islikely to exceed the additional error (approximately 1mm) introduced by 3-D US target localization (Langenand Jones 2001). Because US guidance is capable ofcompensating for internal organ motion, using it to po-sition patients for treatment should still present a signif-icant improvement over the current established methodsin radiotherapy. The experimental methodology pre-sented in this paper can be readily extended to evaluatethe accuracy of other types of patient positioning systemsand to include nonrigid errors.

Acknowledgments—This research was funded by the Natural Sciencesand Engineering Research Council of Canada. Generous support wasgiven to carry out the experimental aspects of this work by Dr. SusanZhang and Kurt Luchka of the Medical Physics department at theBritish Columbia Cancer Agency. Assistance in producing the polymergel was provided by Dr. Andrew Jirasek, Department of Physics andAstronomy, University of Victoria. Neculai Archip was supported inpart by the NIH grants U41 RR019703, P01 CA067165 and R03EB006515.

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