Development of a SPECT-Based Three-Dimensional ...

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Development of a SPECT-Based Three-Dimensional Treatment Planning System for Radioimmunotherapy Huan B. Giap, Daniel J. Macey and Donald A. Podoloff Departments ofRadiation

Physics and Nuclear Medicine, The Universizy of Texas M.D. Anderson Cancer Center,

Houston, Texas

and distant metastases (1). Over the past decade, there Two major obstacles in the development of improved methods for more accurate dose estimatesfor radioimmunotherapyhave been the difficultyin obtaining an accurate patient-specific three

have been more than 100 Rif and radioimmunodiagnosis

clinical trials (2). A variety of malignancies have been treated including: hepatoma, neuroblastoma, melanoma,

dimensional activitymap in vivo and calculatingthe resulting breast, kidney, lung, colon, lymphoma and ovarian malig absorbed dose. We propose a method for three-dimensional nancies. Theoretically, Rif has the potential to deliver internaldosimetrythat integrates the three-dimenaionalactivity radiation more selectively to a tumor than external beam map from SPECT with a dose-pond kernel convolution tech radiation. Even though Rif is still in the developmental nique to provide the three-dimensional distribution of absorbed dose. Methods: Accurate activityquantitation was achieved w@i stage, results from preliminary RIT clinical trials are prom ising. The most common radionuclide used in Rif is 1311. appropnate methods. The count density map from SPECT im has been successfully used to treat thyroid ages was converted into an activityconcentration map with a Although ‘@‘I calibration phantom approach. This map was then convolved with an 1311dose-point kernel and three-dimensional fast Fourier transform to yield three-dimensional distribution of absorbed dose, which was then processed to provide the absorbed dose distribution in regions of interest. Results: The accuracy of quantitative SPECT was validated to be within 16%. The calcu lated penetrating radiation absorbed dose was verified with ther moluminescent dosimeter measurements to be within 8%. With standard organs and configuration, the method calculated ab sorbed dose in good agreement w@ithe MIRDformalism (less than 14%). Conclusion: This method overcomes the limitations

cancer and other malignancies for more than four decades,

of planar imagingtechniques and the current routineimplemen

the therapist in optimizing a treatment planning strategy for each Rif patient (i.e., to specify the minimum adminis tered activity required to elicit significant therapeutic re

tation of the MIRD formalism. The results can be processed to provide the absorbed dose distribution in regions of interest and parameters for treatment optimization. Absorbed dose distribu ton from any plane can be graphically displayed in various ways.

KeyWords:single-photon emissioncomputed tomography; ra dioimmunotherapy; three-dimensional internal dosimetry J NucI Med 1995; 36:1885-1894

he principle of radioimmunotherapy (RIT) is to use radiolabeled antibodies as carriers of radionuclides for the selective destruction of a tumor. RIT should be particularly

beneficial for the treatment of tumors not easily amenable to surgical control and for the treatment of early recurrence @

Nov.4, 1994;rev@ion accepteiMay15,1995. For correspondence or reprints conta@ Huan B. Giap, MD, PhD, Department

ofRadsation Medicine, LomaUndaUnivers@y Med@Center,I l234Anderson St LomaLindaCA92354.

dose estimates for 1311therapies are still uncertain. These dose estimates require calculation of the activity in a spec ified volume of interest and accurate methods for calcula

tion of the absorbed dose. Accurate determination of ab sorbed dose for organs and tumors in patients treated with RIT is essential to establish fundamental dose-response relationships for efficacy and toxicity, which cannot be extrapolated from the dose-response data derived for ex ternal beam therapy. These estimates are required to guide

sponses and, more importantly,

to specify the maximum

administered activity that does not exceed normal tissue toxicity limits). The complicated movement of circulating monoclonal antibodies (MAbs) through and within tumors, which re sults in nonuniform distribution, can be explained by at

least two factors. First, the irregular tumor vasculature produces a complex pattern of diffusion gradients through which the MAb is transported throughout the tumor (3). Second, immunohistochemical studies with MAbs have shown that the antigens against which the MAbs have been

directed are not usually expressed uniformly in the whole tumor cell population. Several investigators have observed nonuniform

distributions

for a variety of antibodies

and

tumor types in animal and human studies (4—10).This heterogeneity was observed on both the macroscopic and microscopic

levels. Inspection

of SPECT transverse-sec

tion images with radiolabeled MAb demonstrates consid

SPECT-BasedThree-DimensionalTreatment PlanningSystem for Radiommurtotherapy• Giap at al.

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erable heterogeneity of activity distribution in organs and tumor at the macroscopic level. In some clinical trials of radiolabeled MAb, organ activ ities have been quantified by planar imaging methods and

average organ absorbed doses calculated by the MIRD formalism and standard man S factors (11 ). The planar imaging method, however, suffers from several deficien

cies. First, because planar images use only two views of a three-dimensional source distribution in the body, activi ties estimated from such two-dimensional images generally suffer from superposition of two or more source organs.

Second, planar images generally cannot provide accurate estimates of depth-dependent three-dimensional distribu

tenuation correction, and correction for Compton scatter and septal penetration. The count density maps from SPED' images were converted to activity concentration maps with a calibration phantom approach. The activity concentration maps were convolved with an 1311dose-point kernel and a three-dimensional fast-Fourier transform ap proach to yield three-dimensional distribution of absorbed dose. The three-dimensional absorbed dose map was then processed to provide the absorbed dose distribution in regions of interest (ROIs). This method can overcome the limitation of planar imaging methods and MIRD S-factors

(i.e., it can provide heterogeneous distributions of ab

tion of activity within tumors or normal tissues. Accord

sorbed dose in source volumes of any size (down to the limitation set by spatial resolution of SPED') and shape

ingly, although it may be acceptable for uniform activity distributions within normal organs, planar activity quanti tation is inadequate for the nonuniform activity distnbu

with nonuniform distributions of activity). A three-dimensional dosimetry program can provide the unique information required to compare different radionu

tions of MAbs typically found in tumors. Third, the two dimensional planar images do not generally provide

clide therapy protocols, e.g., to match the emission of a radionuclide (type of particles and energy) with the re

accurate functional volumetric information necessary for calculation of dose distribution. Therapy radionuclides are selected to deliver most of their dose locally by nonpene mates can result if the functional volume does not match

quired dose distribution. In addition, the three-dimensional information on absorbed dose may be more valuable in the interpretation of the responses of heterogeneous antibody and dose distributions and would supplement information from the analysis of clinical response with dose-volume

the anatomic volume defined by CT or MRI images (12,13). These three deficiencies of planar imaging methods can

map of the activity distribution is desirable in clinical ap

trating radiation. Therefore, significant errors in dose esti

potentially be overcome by SPED'. The MIRD formalism, as currently implemented, is

histograms

or isodose distribution.

plications to demonstrate

A three-dimensional

qualitatively

and quantitatively

widely accepted for calculation of absorbed dose in RIT

the temporal and spatial distribution of the radionuclide in tumors and organs.

and has served as the framework to provide dose esti mates. The advantages in the use of MIRD S-factors are

MATERIALS AND METHODS

ease of application, availability of tabulated data, standard

ization of dose estimates among different investigators and extensive

validation.

There are several disadvantages

in

the use of the MIRD S-factors, however, including: (a) assumption of uniform distribution of activity in each source organ, (b) assumption of uniform deposition of ra diation energy in target regions, (c) lack of information about the distribution of absorbed dose in selected organs

of interest, (d) use of fixed idealized anatomic models (i.e., standard man, child or woman) for organ sizes and shapes and (e) lack of S-factors for tumor volumes. Because no S-factors are available for target volumes other than those defined in the MIRD approach for the standard man, dose estimates for tumors in particular have been a challenge (14). An improved method is required for

dose estimations when the underlying assumptions made by MIRD S-factors are not applicable. With planar imaging methods and MIRD S-factors, the errors of normal organ dose estimates may be as much as 200%, and most of this inaccuracy comes from the uncertainties

in the assessment

of the distribution of radioactivity in both tumors and or gans (15). This may account in part for the large range of

dose responses reported in recent RIT trials.

sion of count density into activity concentration.

The three-di

mensional activity distributions are then converted to absorbed dose with a dose-point kernel convolution technique. The pro

gram was written in FORTRANand compiledusing the Lahey 32-bit compiler (Lahey Computer System, Inc., Incline Village, NV), which contains graphic subroutines for image display. The

program was developed on a 486-DX2 (66 MHz) IBM-compatible personal computer with 16 MB RAM. The computer code con sists of two modules: SPECTRIT and DOSEFFT. SPED'RIT, which stands for “quantitative SPED' for RIT,―performs activ

ity quantitationwithprojectiondata acquiredfroman Angercam era system. DOSEFFT, which stands for “dose calculation by fast Fourier transform (FF1'),―provides the absorbed dose from the activity map furnished from SPECTRIT, which prompts the user to select options from menus, that includes methods of noise filtering, scatter correction, image reconstruction and attenuation correction. The user can bypass all these menus with a default method at the beginning of the program. The user can also use a pointing device (mouse) to select target volume and to obtain

This paper proposes a method for internal dosimetry of

‘@‘I based on SPED'. Accurate SPED' quantitation was achieved with appropriate methods for noise filtering, at

1886

An overview of the proposed three-dimensional treatment planning method is outlined in Figure 1, which includes two major components: activity quantitation by SPED' and absorbed dose calculation. The principal components of quantitative SPEC!' are noise filtering, subtraction of Compton scatter, boundary detec tion, image reconstruction, correction for attenuation and conver

information

such

as functional

volume,

descriptive

statistics

of

activity or dose or activity/dose profile plots about an ROl. DOSEFFT can provide important parameters for treatment opti

The Journal of Nuclear Medicine• Vol.36 • No. 10 • October 1995

Downloaded from jnm.snmjournals.org by on February 15, 2017. For personal use only.

@ @ @

@ @

(DOSEFF7)

De@othm@ar

Noise Filtering Scatter Correction I

@

nT@•atm•nt@

Optimization

Image Reconstruction

Attenuation corrctio@J

FIGURE 1. Overviewof SPECT-baSed three-dimensional treatment planning strat

egy for internaldoalmetry.Ot@ectiveis to @ @

I

:

Countdnity map

(counts/voxel)

determine three-dimensional distribution of A@nc•ntrado@ map (MBq/mI)

@

(cGy/h)

dose-pout

kernel

mization, such as differential and integral dose-volume histo grams, isodose distributions, dose statistics (range, average and s.d.), points of regret and so forth. Absorbed dose distribution from any plane can be graphically displayed as line isodose, gray scale, colorwash, line profile and three-dimensional surface dis

absorbed dose rate. SPECT method @as devebped to providethree-dimensionalac tivitydistribution,whichwas then convolved wrnidose-pointkerneltechniqueto provide three-dimensional map of absorbed dose rate. Components of the SPECTRfl and DOSEFFT computer codes are outlinedwIth dashed lines.

Giap (18). For the studies in this paper, the projection data were filtered with a Metz ifiter, which was optimized specifically for ‘@‘I

andthe TrionixBIADsystem(19).The Metzfilterhas been found to improve image quality and activity quantitationsignificantly compared with conventional

low-pass filters (20—24).The Metz

play from transverse, coronalor sagittalviews. filter is the default noise filter for SPED'RIT, but the user can The diagramin Figure1onlyshowsthe calculationof absorbed select any conventional low-pass filters such as the ramp, Shepp dose rate from a single SPED' study. To calculate total absorbed

Logan, Hann, Hamming, Parzen and Butterworth.

dose for a RIT study, the time-dependentSPECT-basedactivity The subtractionof Comptonscatter and septalpenetrationwas measurementmustbe convertedto cumulatedactivityby numeric performed before image reconstruction with a modified convolu integration of a fitted time-activity function, which is described in

detail in references 11, 16 and 17.

DataAcquisItIon SPEC!' projection data were acquired with a Trionix dual headed BIAD gamma camera (Tnonix Research Laboratory, Inc., Twinsburg,OH). The acquisitionparameterswere 64 x 64 matrix, 45 projection views for each detector, 20 sec/view, circa tar orbit with radius of rotation of 27.3 cm and a high-energy parallel-hole collimator. Three energy windows were used to ac quire projection data for ‘@‘i. The photopeak window of 15% width was centered at 364 keV. The lower scatter window (15%

width) was located directly below the photopeak window. The upper scatter window (15% width) was located directly above the photopeak window. The total projection counts acquired were 9.3 million (5 million from the photopeak window and 2.6 million and 1.7 million from the lower and upper scatter windows). The pro jection data (binary format) were transferred by Ethernet link to the personal computer for all SPED' processing. The section images were reconstructed in a 128 x 128 matrix with 5.3 x 5.3-mm pixel size and 10.6-mm slice thickness.

tion-subtraction method with a dual-energy window (MCS-DEW) (25). This method can be considered a combination of the

Jaszczak et al. (26) dual-energywindow (DEW)approachwith constant multiplier and the Axelsson et at. (27) convolution sub

traction (CS) method. The major disadvantageof DEW is the assumptionof a constant scatter fraction in the photopeakwin dow at all projectionbins and views. This assumptionmay be valid for a source located centrally in a symmetric scattering

medium(such as a water cylinder)but may be invalidwhen the source moves off center or when the scattering medium is non uniform. The @S method provides scatter fractions that are de

pendent on source locations, but its major disadvantageis the assumptionof the spatial symmetry of the scatter distribution (i.e., the left and right wing ofthe scatter distribution curve are the

same and can be describedby a monoexponentialfunction).This mayexplainwhy it does not predictscatter distributionswhenthe source is off center or near the periphery of a transverse section.

The @S methoddoes not utilizeany informationfromthe scatter window,which can provide informationabout the symmetryof the scatter distribution.The MCS-DEWattempts to correct for these deficiencies in both methods. Knowledge of the spatial

SPECT Quantitatlon Methods for SPED' quantitation include noise filtering, cor

distributionand magnitudeof data collectedfrom a scatter win dow is used to predictthe scatter fractionof photopeakevents at

rectionfor Comptonscatter and septalpenetration,imagerecon

each location (projection bins and angles). The method relies on

struction, boundary detection and attenuation correction. General

estimatesof the scatter fractionin the photopeakprojectionwin dows(Pr) fromevents detectedin an overlappingscatter window

discussion of SPED' quantitation can be found in Macey and

SPECT-Based Three-Dimensional Treatment Planning System for Radiommunotherapy • Giap at al.

I 887

Downloaded from jnm.snmjournals.org by on February 15, 2017. For personal use only. (Ps) before reconstruction. An estimate of the true photopeak

projectionevents (PT)is givenby PT=PP—cPSØQ,

@

formly attenuating medium, the equation for PRIGM can be sim plified to

Eq.1

where 0 denotesa convolutionoperation,c is an optimizedcon

C(i,j)=

stant and 0 is the scatter transferfunction, which convolves with the projection data from a scatter window to yield an estimate of

canselectotherscattercorrection methodsinSPECTRIT suchas convolutionsubtraction,dual-energywindowwith constant scat ter fraction,DEWwith conditionalscatter fraction(29)or triple

2 nNf2 @:e@@―@'0@ L(ij,O@ +

.

Eq.4

n=1

the scatter fraction in the photopeak window. The function 0, which is expressed in terms of spatial frequency, is the ratio of two modulation transfer functions (MTF): one MTF characterizes the scatter data under the photopeak window; the other MTF represents the scatter data in the scatter window (28). The user

1

The user can select other methods for attenuation correction in SPECTRIT such as the Chang or iterative least-square intrinsic method. The object boundary and attenuation map are required for all attenuation correction methods. In SPED', the term boundary is

usually reserved to describe the body or object contours. The

difficulty in boundary detection is due to the low spatial resolu energy window methods (30). tion, Compton scatter and high noise levels characteristic of The default image reconstruction method for SPECTRIT is the SPEC!' images. The problem becomes more significant as the size filtered backprojection algorithm, but the user can select other of an object decreases and the contrast between the object and methods from the RECLBL Library [provided by the Lawrence background activity decreases. A small error in defining the Berkeley Laboratory ofThe University of California (31)] such as boundary can lead to significant errors in attenuation correction. the iterative least-squarestechnique, backprojectionof filtered The authorsused a simple methodofboundary detection thatuses projections and configuration-spaceconvolution algorithm. an adaptive thresholding technique with a gray-level histogram Photon attenuationmay be the single most significantfactor in applied to data in a scatter image (34—36). Evaluation of the quantitative SPED'. Attenuation is a more significant problem in performanceof the methodwith a 22-cmdiametercylinderphan SPED' than in PET and other conventionalimagingsystems(CT tom filled with uniform activity concentrations indicated good and MRI) because the detected photons are at much lower count boundary definition to ±1 cm. Advantages of this method include densities. To correct for attenuation in SPECT', the following no modification to the hardware, no separate transmission scan, parameters are required: (1) an accurate distribution of counts in simple implementation and quick execution. This method, how the object, (2) an accurate boundary of the transverse section and ever, does not provide the attenuation map required for correction (3)the distribution of attenuating tissues in the object expressed as of nonuniform attenuating media. This shortcoming can be re an attenuation map appropriatefor the ED' photon energy. A solved with a specially designed detector system that could per form simultaneous transmission CT (TO') and SPED' to provide new postreconstruction attenuation method, called postrecon struction inverse of the sum of geometric means (PRIGM), was the map of attenuation coefficients and anatomy of the patient used in this study (32). The PRIGMmethodcorrects each pixel (i, (37,38). The new generation of detector systems, such as the

j)inthereconstructed imageforattenuation bymultiplication by Trionix TRIAD and ADAC Genesis VERTEX, offer simulta neous TCT-ECT capability.

an average attenuation correction factor C(i, j), defined as 1 C(i, j) = @

SPECT CalIbration MethOd .,

2 nN/2

Eq. 2

methods

for SPED'

depend

on a determina

the bodysection,and the use ofthe sensitivitymeasuredfromthis

n= 1

where N is the total numberof projectionangles, n is the index for

@

Most quantitation

tion of sensitivity from a phantom of the same size and shape as

- :@: [e@ @i,jO@) e@LJ(j,j@+ ir)jlf2

phantom to convert the counts in the section image to activity. Other methods that have been used to convert SPED' counts to activity include calibration of SPED' counts with the conjugate

projectionangle0, L(i,j, O@) is the distancefrompixel(i,j) to the bodyoutlinealongthe projectionviewn (or angleO@) alongwhich view planar method (39), retrospective calibration by activity the summation takes place and is the effective attenuation concentrationin urine (40) or measurementof the activity con coefficient for the photons used for imaging. Note that the atten uation map is used in the calculation of the attenuation correction factor qi, j), not the emission projection data. This method is similarto the approachintroducedby Chang (33). 1 C'(i,j)=

@

Eq.3

centration in peripheral blood (41). Conversion of the counts in a

transverse section image to activity requires a sensitivity factor, which should be based on a multitude of acquisition and image processing parameters (42). Detector sensitivity can be consid ered a conversion factor (or calibration factor) to convert the count density (numberof counts per voxel) in the reconstructed images to activity concentration at the corresponding location in

the object. For SPED', this conversionfactor dependson many

- :@: eL L(i,j,O@)

variables, such as source-to-collimator distance, activity levels, image acquisition parameters, methods of attenuation correction,

The Chang's correction factor C'(i, j) is the inverse of the arithmetic average of all attenuation path length; PRIGM takes the geometric mean of conjugate attenuation path lengths for the correctionmatrix.The PRIGMcorrectionmatrixis less perturbed near the edges of organs of nonuniformattenuationbecause the exponential terms are weighted by the conjugate terms. For uni

1888

scatter correction, noise filtering and image reconstruction algo rithm. In this study, a new method of SPED' calibrationwas used to

convertcount densityto activityconcentrationwith a calibration phantom to be used with the Alderson abdominalphantom (Fig. 2). The schematic of calibration process is shown in Figure 3. It

The Journal of NuclearMedicine• Vol.36 • No. 10 • October 1995

Downloaded from jnm.snmjournals.org by on February 15, 2017. For personal use only. count density (in counts/voxel) and I@is the intercept of the calibration curve (in megabecquerels per milliliter). Count density in any voxel or ROI in the image can be converted to activity concentration with this calibration curve. Dose-PoInt

Kernel Convolution

TechnIque

The total absorbeddoserate 13(r)at locationr consistsof two components:penetrating[D@(r)] and nonpenetrating[D@@(r)].

II

I@(r)= l3@(r)+ l3@@(r).

Eq.6

The dose rate from nonpenetratingradiations [I3@@(r)] is pro duced by low-energy gamma, x-ray and beta emissions that de

posit all their energyin the SPED' sourcevoxel, whicht@rpically has a dimensionof 5 mm or greater. The calculationof D@@(r) is simplebecause it is equal to the product of the activityconcen tration A(r, t) and an equilibrium dose constant cd(cd = 11.03 cGy-gfMBq-hr or 0.408 cGy-g/hr-@Cifor 131J).The dose rate from penetrating radiations [D@(r)]is contributed by high-energy FiGURE2. Alderson abdominal phantom andSPED'calibrationgamma, x-ray and beta emissions, which can deposit energy be phantom (wIthoutbohis). yond the source voxel. The calculationof D@(r)is more involved

andgivenby the convolutionof activityconcentrationmatrixwith consisted of four cylindrical tubes that contained measured activ ity concentrations over a range that covered the anticipated max imum activity concentration in any organ in the body. The tubes, contained in a Lucite housing, were 42 cm long and 1.87 cm in diameter. The tube dimensions were chosen with consideration of the spatial resolution of the detector system, contours of the patient's body and minimal radiation hazard. From the recon structed images, the count densities and the known activity con centrations of the four tubes were used to construct a calibration curve for every individual section. This calibration curve was a straight line through the four data points fitted by the least-squares technique.

a penetrating dose-point kernel (k.,,).

Ô@(r) = J@A(r') 0 k,,,(r—r') . dV',

Eq.7

whereI@@(r) is the absorbeddoserate (incentigraysper hour)from penetrating radiations, A(r') is the activity concentration (in mega becquerels per milliliter)in the voxel at location r' and k.,,(r—r') is the penetrating dose-point kernel (in centigrays per megabec querel per hour) in tissue. The dose-point kernel represents the spherically symmetric dose rate distributioncaused by a point source of activity at the center of a sphere. The dose rate at a given distance from a point source of activity in an attenuating and

Eq. 5

scattering medium has been extensively studied. The dose-point kernel in spheric coordinates can be derived from MIRD Pam

where A is the activity concentration (in megabecquerelsper

phiet No. 2 for photonemissions(43)and MIRDPamphletNo. 7

@

A = m@ C + I@,

milliliter), m@is the slopeof calibrationcurve (in megabecquerels for beta emissions(44).A morefundamentalapproachis to use a per milliliter per counts per voxel) of transverse section z, C is the MonteCarlocalculation(45)witha computercodesuchas EGS-4

CalIbrationCurve Ac@* Co@

SPECTCelibration phantom (0.03, 0.07, 0.14, 0.28)

FIGURE 3. SPECTcalibrationphantom.

Liver

Tumor (0.18)

(0.14) Cou@ densfty

I

2

3

4

ActMtyconcentrations of the organs, tumor and calibration phantom are shown In megabecquerelspermillilIter. Thefourtubes of SPECT calibrationphantom contain in creasing and known actMty concentrations. Tubes are contained In Lucite housing. For each reconstructed transverse Image, count

densitiesand knownactivityconcentrations of the four tubes are used to constructa calibrationcurve, which Is a straight line fit

ted throughthe fourdata points.Countden SPECTcalibrationphantom

sity at any voxel or AOl Inthe image can be

convertedintoactivityconcentrationwiththe calibrationcurve.

SPECT-Based Three-Dimensional Treatment Planning System for Radioimmunotherapy • Giap et al.

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(4447) andETRAN (48).Preliminarystudieshaveshownthatthe

fl.a@id@amtsrc@

EGS-4 code and MIRD equations provide similar dose-point ker nels. In this study, dose-point kernels were calculated by both EGS-4 code and MIRD pamphlets, and this redundancy was used for verification. The dose-point kernel in spherical coordinates @

was converted into rectangularcoordinatesto use the three-di mensionalFEToperation.Thisthree-dimensionalrectangularma

hilledwith 1-131

@cm damtv

N@

TLD in

@r

coldrod

:::i::

///

ltD

@i [email protected]

trix had to be the same size as the activity matrix, which was 128 x 128x 32in this study. The conversionof a spheric,kernelinto a rectangularkernel was done with an interpolationalgorithm.

The convolutionwas carried out with a Fourier convolution technique. The procedure applied a three-dimensional fast Fou riertransformationof the activity distribution,F3{A(r)},andof the penetratingdose-point kernel, F3{k,,(r)}.The subscript3 indicates that this is a three-dimensional Fourier transform. The complex three-dimensionalarrayswere multipliedtogether, and the prod uct was inverse-Fourier transformed to yield the distribution of absorbed dose rate caused by penetrating radiation. The calcula tion is described by the following equation. t@@(r) = F@ ‘[F3{A(r)} . F3{k,,,(r)}].

Eq. 8

Both penetrating and nonpenetrating components can be cal culated by the Fourierconvolution method by substitutionof the penetrating dose-point kernel Ic,,by the total dose-point kernel k@0151 in the convolution operation, as follows.

in the phantom both at the times when the TLDs were inserted and removed. The assay was performed by measurement of the averageactivity in five aliquotsof solution fromthe phantomwith two different detectors (CAPINTEC CRC-7 Capintec, Inc., Ram

sey, NJ) and Nal spectrometer model 2600 (Ludlum Instruments,

Inc., Sweetwater,TX). SPED' projectionimagesof the phantom

from the two types of radiation to the total absorbed dose. In clinical applications, in which the total dose is ofinterest, the is used. There is no difference in computer time whether k@o@i or k@is used in the convolution calculation. Similarly, for a radionu clide that emits only beta particle, such as @°Y, the Ic,,in the

ing activity concentrations were 0.71, 1.4 and 1.8 MBq/ml (1.88, 3.69, 4.94 uCi/ml). The activity concentrations in the calibration tubes were 0.03, 0.07, 0.14 and 0.28 MBq/ml.

where ktotai@S defined as k@ = k@+ k@1,.

Eq. 10

Penetrating and nonpenetrating components were calculated separatelyin this study to examine the magnitudeof contributions @

stacked small TW capsules.

were acquired with a Trionix BIAD detector system immediately after the TLDs were removed. The second phantom used to validate the method was the Alderson abdominal phantom (Fig. 2), which contained various activity concentrations of ‘31I in the liver, spleen and a spherical tumor(diameter= 5.52 cm). This phantomwas scanned with the calibrationphantom. The transverse section of these two phan torus is shown in Figure 3. The volumes for the liver, spleen and

O(r) = F3 1[F3{A(r)}. F3{lç0@(r)}J,

Eq. 9

FiGURE 4. LucIterod (8-cmdiameter)centered in the 22-cm diameter cylinderfilledwith 1311solution.TW tube contains 36

convolutionoperation includes the contributionfrom the high energy beta radiation.The location of the beta particle has to be assumed either at the center of the voxel or distributed uniformly throughoutthe voxel.

The first phantom, which is shown in Figure 4, was used to validatethe dose calculationcausedby the penetratingradiations with the three-dimensional FFT approach. This phantom con sisted of an 8-cm diameter Lucite rod centered in a 22-cm diam eter cylinder filled with ‘@‘I at an initial concentration of 0.123 MBq/mI (3.3 pCi/ml) and two thermoluminescent dosimeter (TLD) tubes. Each TLD tube contained 36 small plastic cylinders (5 mm diameter and height) filled with Li3F powder (TLD-100). The wall thickness of the tissue-equivalentTLD tube was 6 mm,

tumorwere 1579,159and88ml,respectively,andthe correspond

RESULTS AND DISCUSSIONS The TLD measurements of absorbed dose along the diameter of the cylindric phantom are shown in Figure 5, which displays absorbed dose caused by penetrating radi

ation (Dr) compared with the TLD location. The TLD measurements (circled data points) are compared with the

calculated absorbed dose profile shown as a solid line. The results show good agreement (within 8%) between the cal culated and measured data inside the phantom (in tissue equivalent material). Outside the phantom, i.e., in air, the

ThD measurement followed the inverse-distance-square

whichwas sufficientto stop the beta radiationemittedby ‘@‘I. The pattern, although the calculated dose, which assumed the TLD tubes were located in defined positions in an 8-cm diameter medium as tissue equivalent, was much higher. This result Lucite rod that was inserted throughthe ‘31I-fflled center of the indicates one limitation of the dose-point kernel convolu

cylinder. The TLDs were exposed to the photons of 131!for one tion technique, which was noted by Meredith et al. (12). week and then removed from the phantom.After another2 days, The dose-point kernel convolution technique is strictly the TLDs were read with a MARK IV TLD reader model 1100 (Radiation Detection Co.). The TLD reading(in voltages per gram valid only for a homogeneous medium for which the ker ofpowder)were converted into absorbeddose (incentigrays)with nels are calculated (in this study, the medium was water). a TLD standard that had been irradiatedto a calibrated @°CoThis technique becomes less accurate in regions of the beam. The readings were also corrected for nonlinearity. The body where different tissue densities coexist, such as lung activity

1890

concentration

was determined

by an assay of the solution

or bone. Kwok et al. (49) reported that the dose to tissue

The Journal of Nudear Medicine• Vol.36 • No. 10 • October 1995

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Although these conditions may not be realistic in a clinical

@ene@a@ng)Abs?ibed I@se(cOy)

setting, the results are nevertheless valid for this experi ment. Volumetric calculations from SPED' images are fundamentally challenging because of the limited spatial

;:i‘‘:[email protected]—@——. resolution, Compton scatter and high noise levels charac teristic of SPED' images. The problem becomes more significant as the size of an object and the contrast between

the object and background activity decrease. Volumetric

—@.—

md—30

calculation

Olki

::— JIlTIffluhIIlII40..:—@

10

20

problem

compared

with activity

parison of the true activity concentrations with the values

calculated for the whole liver, spleen and tumor. The re sults are shown in Table 2. The percent differences were

.——@E1.131

U

is a separate

quantitation and dose calculation, and rigorous investiga tion of this issue is beyond the scope of this present study. The activity quantitation results were validated by com

30

40

50

60

70

80

Poelhion (ptxd I)

4.4%, — 14.3% and — 16.3% for the liver, spleen and tumor,

respectively.

It is well recognized

that the accuracy

de

FIGURE5. Absorbed dosecausedby penetrating radiations pends on source volume. The error in localization of the measured from TLD vs. those calculated by DOSEFFT along the center of 131l-fllled22-cm diameter cylindricphantom withthe 8-cm

edges of larger organs is less. Because the volume of the

20% to 40% because of backscatter of low-energy elec

liver is more than 10 times the volume of either the spleen or tumor, the observed percent difference in activity quan titation for the three organs was as expected. The s.d. indicates the precision of the measurement, which was found to be about 0.3 p@CVml. The coefficient of variation

trons. This limitation is significant for absorbed dose esti mates for the thorax and bone marrow. Three levels of quantitation were evaluated with the

relates to the relative precision of the calculation; the av erage COV was found to be about 11% for this study. The

diameterLuciterod. within 2 mm of a tissue-bone interface is underestimated by

(COV) is defined as the ratio of the s.d. over the mean and

Alderson phantom: functional volumes, activity concentra

magnitudes

tions and absorbed dose rates. The functional volumes are

variation of activity concentration within the organs of interest.

required for activity quantitation and absorbed dose calcu

lation. The approach used for volume quantitation was a gray-level histogram method (GLH) (50). For each trans verse section, sources from organs and tumor were iden tilled, and edges of organs were found with a threshold selected by the GLH method. The results which showed

less than 6% percent difference are shown in Table 1. The good agreement observed here resulted from the following

conditions: 1. The phantom sizes and activity concentrations were known in advance. 2. No background activity was present. 3. The tumor did not overlap organs. 4. Tumor volume was larger than the resolution volume

of the detector system. 5.

Activity

concentrations

were

sufficiently

high

to pro

vide goodSPEC!'images. TABLE 1 Accuracy of Volume Determined from SPECT for Source Organs wIth lodine-131 in the ,4jderson Abdominal Phantom CalculatedTrue @dume (ml)volume (ml)Difference (%)Liver157916202.6Spleen1591685.9Tumor8885—4.0

of the s.d. and COV are correlated

with the

The activity distributions of the Alderson abdominal

phantom were then used in the calculation of absorbed dose rate by the dose-point kernel convolution technique. The calculated absorbed dose rates were accumulated slice by slice for the whole liver, the spleen and the tumor. These accumulations were used to calculate the average absorbed dose rates for each source volume at the time of

the SPED' acquisition. These average absorbed dose rates were then compared with those calculated by MIRD S-fac tors. The S-factors from MIRD Pamphlet No. 11 (52)were corrected for the differences in organ mass in this phantom

from those used in the MIRD reference man model. The S-factor for the spheric tumor was derived from MIRD Pamphlet No. 8 (53). The results are shown in Table 3. The percentage differences of the mean absorbed dose rate calculated for the two methods were 9. 1%, 13.7% and 0.9% TABLE 2

Accuracyof ActivityQuantitationfrom SPECT for Source Organs wIth lodine-131 in the Aiderson Abdominal Phantom True activIty actMty concentration concentration of (JhCi/ml@Calculated (%)Liver1.881.964.416Spleen3.693.16—14.312Tumor4.944.14—16.36.0 variation (@CVmODifference (%)Coefficient

SPECT-Based Three-Dhnensional Treatment Planning System for Radionimunotherapy • Giap at al.

1891

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for the liver, spleen and tumor, respectively. MIRD 5-fac tors only provide the mean absorbed dose rate, whereas the convolution technique provides the three-dimensional

Differential Dose-volum.Histogram

distribution of dose rate. The results can be processed to

provide the absorbed dose distribution in ROIs and param eters for treatment

optimization

such as differential and

integral dose-volume histograms (Fig. 6), isodose distribu tion, dose statistics (range, average and s.d.), points of regret (cold spot in tumor or hot spot in normal organs) and so forth. Distributions of activity concentration and ab sorbed dose rate from any plane can be graphically dis played as line isodose, gray-scale, colorwash, line profile and three-dimensional surface display from transverse, coronal or sagittal view (Figs. 7, 8 and 9). Because there

Volumsfrc@on

Absorbid dos• raio@

are no assumptions with regard to the size, shape and

location of target volume, this method can be applied to any patient's anatomy, tumor location and size (limited

Cumulative

Dosi-volum.

Histogram

only by the spatial resolution of the SPED' camera). The three-dimensional FFT calculation for this matrix of 128 x 128 x 16 was completed in 45 sec on a 486-DX2 (66 MHz)

personal computer. This high speed was achieved through the efficiency of the FFT algorithm. The reproducibility of the method was evaluated by the repetition of phantom

Vokimi fraction

studies with different source configurations and activity

concentrations.

CONCLUSION 0.3

Accurate specification of radiation absorbed dose is cen tral to all treatment modalities in radiation oncology and to establish fundamental dose-response relationships in RIT.

The objective of this study was to develop a method that

0$

0.9

1.2

1.5

Absorb.d dm51rat, (cOylhr)

FIGURE6. Differential andcumulative dose-volume hlatogram of the spleen, liverand tumor in Alderson abdominal phantom.

could provide more accurate Rif dosimetry, which de pends on activity quantitation by SPED' and the dose

point kernel convolution technique. The objective of SPED' is to represent the three-dimensional distribution of activity in large tumors and organs accurately within the available spatial resolution. Achievement of this goal is constrained by the physical characteristics of the detector

I- 131 Spleen/Liver/Tumor S'ice

AC1T\'IJ

•

@

construction

algorithm, noise filter and scatter and attenu

0

At. (uCi'

@ @

@

treatment planning approach are di

TABLE 3 Com@on of Average Absorbed Dose Rate (Both Penetrating and Nonpenetrating Radiation) Calculated for Source Organs in the ,@JdersonAbdominal Phantom MIRDaverage average absorbed dose rate absorbed dose rate (cOy/br)CalcUlated (cGy/hr)Difference (%)[email protected]

@li':e

j

cn:. @).47 O/@4t.41 .@ 2.35 :82 @@l)

ation correction methods). Important limitations of this three-dimensional

‘

activity

COrIC ENTF@TIOt@ [email protected]

S 27

system (variation of spatial resolution and sensitivity across the field of view of the detector), the emission char

acteristics of radionuclides and the methods used to derive quantitative information from the acquired data (image re

phontom without background

—@1

(Tronsverse plane)

# 28

@Ii@e •29

•--@

0@@@@

4.24 4.71 @.I8 @.65 12 [email protected]ô7.53 9:

@s

•

•

•

•

•

.

.

I

ABSORBED DOSE RATE I1AP slice

:@

Slice S 2t

[Shce$27 _________ ..

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Rcit@

72

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182

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FiGURE 7. Gray-scaleimagesforfourtransversesectionsofthe Alderson abdominal phantom. Top row shows activItyconcentration maps, and bottom row shows absorbed dose rate maps of the

correspondingfourslices.Locationsof liver,spleen,tumorand cal ibrationphantomare similarto the schematic in Figure3. Gray scales that correlate differentshades of gray to the absorbed dose rate and activityconcentration are also shown.

1892

The Journal of Nuclear Medicine• Vol.36 • No. 10 • October 1995

@

@7J@ 9

Downloaded from jnm.snmjournals.org by on February 15, 2017. For personal use only.

I- 131 Spleen/Liver/Tumor (Coronal plane)

phantom without background

ACTIVIT

the limiting normal organ for RIT.

CONCENTRATION i.1@F' I@

@

H:.

[@_@:@:e

A.:t. Cr. ‘.47“-#4 141 (tiCi/ ‘@l1

ltd

2. Dose calculation for bone marrow, which is usually

activity 4

3. SPED' quantitation for heterogeneous media, which requires a detector system capable of simultaneous acquisition of transmission and emission images and

L

@___________

::o

2t2

35 --

3.77 4@4 4.71 114 0 1 I I

4E1@@11RBED DOSERATEMAF

also methods of image correlation of ED' and TD' images. I....

1@ @•4 ;e

4. Validationof the methodin Rif patients,whichre quires a method of measuring absorbed dose in vivo

such as implanted miniature-liDs

(51).

5. New methods for more accurate volumetric

calcula

tions from SPED' images.

@ L:b@

6. Extension of the method to other radionuclides.

7. A practicalmethodfor calculationof the cumulative I *111

activity and total absorbed dose.

FiGURE8. Gray-scaleimagesfor fourcoronalsectionsof

Alderson abdominal phantom. Top row shows activityconcentration ACKNOWLEDGMENTS maps, and the bottom row shows absorbed dose rate maps of the This project was funded in part by an MD/PhD fellowship grant corresponding four slices. The round AOl on the left is the tumor, of the University of Texas Health Science Center at Houston, the andthe oval ROlon the nghtlathe spleen. Grayscales thatcorrelate

differentshades of gray to the absorbed dose rate and activity University of Texas M.D. Anderson Cancer Center and National

concentration are also shown.

rectly related to the inherent limitations of SPED' imaging with an Anger camera. The minimal volume in which ab sorbed doses can be derived with this method depends on the spatial resolution of the SPED' images, typically 1 to 2 cm. This limits the specification

of dose distribution

to

volumes greater than 8 ml. Future studies should include the following areas: 1. Investigation of the dose-point kernel convolution technique for heterogeneous

tissues, i.e., bone mar

row and lung.

I- 1:31 Spleen/Liver/Tumor @:.Ii.— I

Zwelling.

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Development of a SPECT-Based Three-Dimensional Treatment Planning System for Radioimmunotherapy Huan B. Giap, Daniel J. Macey and Donald A. Podoloff J Nucl Med. 1995;36:1885-1894.

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