Optimization of a pinhole collimator in a SPECT scintillating fiber detector system: A Monte Carlo analysis

Radiar. Phys. Chem. Vol. 43, No. 4, pp. 383-392, 1994 Printed in Great Britain. All rights reserved

0969-806X/94 S6.00 + 0.00 Copyright 0 1994 Pergamon Press Ltd

OPTIMIZATION OF A PINHOLE COLLIMATOR IN A SPECT SCINTILLATING FIBER DETECTOR SYSTEM: A MONTE CARLO ANALYSIS GEORGEJ. HADEMENOS* Department of Physics, University of Texas at Dallas, P.O. Box 830688, Mail Stop F023, Richardson, TX 75083, U.S.A. (Received 1 May 1992; accepted 1 December 1992) Abstract-Monte Carlo simulations were used to optimize the dimensions of a lead pinhole collimator in a SPECT system consisting of a line of equally spaced Tc-99m point sources and a plastic scintillating fiber detector. The optimization was performed by evaluating the spatial resolution and scanner sensitivity for each source distribution location and collimator parameter variation. An optimal spatial resolution of 0.43 cm FWHM was observed for a source distribution positioned 2.Ocm from the collimated scintillating fiber detection system with a pinhole radius of 1.0 mm and a collimator thickness of 3.0 cm for a 10,000 emission photon simulation. The optimal sensitivity of 257 cps/MBq (7704 cps/mCi) occurred for a source distance of 2.0cm, a radius of 3.0mm and a thickness of 3.0cm.

1. INTRODUCTION The collimator is a vital component in a Single Photon Emission Computed Tomography (SPECT) imaging system and is highly specific to radionuclide energy, location and size of the organ of interest, and desired imaging parameter enhancement. It is designed to reduce scatter and allows the photon detector to localize the radionuclide in the patient by

absorbing and stopping a large fraction of the emitted radiation except for that arriving almost perpendicular to the detector surface (Mettler and Guiberteau, 1983). Collimator optimization (Busemann-Sokole, 1987; Chang et al. 1988; Gantet et al., 1990; Gullberg et al., 1990; Wilson, 1988) is a process which considers the aforementioned factors in designing an “ideal” collimator for a particular imaging situation and, in effect, must be performed before further time and financial resources are invested in the experimental phases of construction and testing as well as clinical implementation. The most efficient and effective approach toward collimator optimization is the application of Monte Carlo theory for accurately characterizing the variation of collimator parameters and subsequent source distribution images by simulating the imaging characteristics of the detector system. In a Monte Carlo program, the probability distribution function governing a particular aspect of the overall process is randomly sampled in order to realistically simulate the system environment.

*Address all correspondence

and reprint requests to: Dr George J. Hademenos, Division of Medical ImagingI72 115, Department of Radiological Sciences, University of California, Los Angeles, CA 90025, U.S.A.

In testing the clinical feasibility of a detection method for SPECT, the primary objective is to assess the capabilities of a y-ray detector by observing its response to radionuclide distributions with welldefined energies. The problems arise in tracking the photons as they penetrate a lead collimator and deposit their energy via several possible physical interactions within the detector. It is an almost impossible task to determine, both qualitatively and quantitatively, the behavior of the photons and generated electrons analytically due to the random nature of the particle interactions. The random nature of the particle interactions makes this type of problem a candidate for Monte Carlo applications. In this paper, an experimentally verified Monte Carlo simulation program is presented in which photons from a line of Tc-99m point sources impinge upon a lead pinhold collimator and interact with a plastic scintillating fiber detector. The structure of the Monte Carlo code and the various components of the SPECT detector system are described in detail, followed by results from the optimization procedure and a discussion. 2. EXPERIMENTAL

Many researchers in medical physics have employed Monte Carlo simulations toward the development and improvement of all aspects of SPECT imaging (Beck el al., 1982; Floyd et al., 1984, 1985; Hademenos et al., 1993; Ljungberg and Strand, 1989; Persliden, 1984). In this investigation, a Monte Carlo simulation is used to accurately characterize the response of a scintillating fiber detector while optimizing the parameters of a pinhole collimator toward a highly efficient yet clinically effective detector system for applications in SPECT. Two factors which

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serve as optimization indices for a SPECT detector system are the spatial resolution and scanner sensitivity. 2.1. Spatial resolution and scanner sensitivity The primary objective of this Monte Carlo investigation was to assess the clinical feasibility of a plastic scintillating fiber detector in line with a pinhole collimator. This can be done on both a qualitative and quantitative scale through determination of the spatial resolution and scanner sensitivity of the detector system. The physical constraints placed upon such a system to yield the best possible image within a reasonable amount of time arises from the inverse relation between the two factors. In order to achieve a high spatial resolution of the detector system, the pinhole radius should be as small as possible while maintaining the ability to resolve the image of two distinct objects. However, as the radius is decreased, so too is the amount of detectable radiation, making it much more difficult for a sufficient number of photons to be detected in a reasonable amount of time. This results in a reduction in the scanner sensitivity. Thus, a compromise between the resolution and sensitivity must be reached. 2.1.1. Spatial resolution. Spatial resolution refers to the capability of an imaging system to record a distinct image of two or more closely spaced, high contrast objects and is dependent on the aperture size (Morgan, 1983). It also depends on the dimensions of the detector system, i.e. the source-to-collimator and source-to-detector distances. The detector response to a radioactive point source, plotted as count density versus pixel location, results in a Gaussian or bellshaped curve. The spatial resolution of a detector is assessed in terms of the full width at half maximum (FWHM) of the Gaussian curve. The spatial resolution of a detector system, R,, is determined from geometric considerations and is defined by two components: (1) intrinsic resolution, R,, obtained by the detector and the associated electronics; and (2) collimator resolution, R,, provided by the geometrical parameters of the collimator. The system resolution is thus given by: Rs = (Rf + Rf)“* (cm FWHM)

(1)

The intrinsic resolution in equation (1) was obtained by acquiring images of the linear source distribution at distances of 2.0, 4.0, 6.0, 8.0 and 10.0 cm from the detector system without the pinhole collimator. The images were then compressed into line spread functions (LSFs) from which the intrinsic resolution or the FWHM of the LSFs was determined by Ri = 2.3% where d is the standard deviation of the LSF. 2.1.2. Scanner sensitivity. As the spatial resolution is a qualitative index of measurement, the scanner sensitivity is more of a quantitative one. The sensitivity denotes the number of counts detected per unit

time per source activity. It is defined as (Brownell, 1959): S = 3.7 x 10’ nGf(cps/mCi) (2) where S is the scanner sensitivity, n is the fraction of disintegrations emitting the photon to be detected, G is the geometrical efficiency and f is the absolute detector efficiency. The parameter, TV,is, in effect, the fractional abundance of the given photon emission and is unique with respect to the radionuclide under consideration. For Tc-99m, this value is equal to 0.88. The geometrical efficiency, G, accounts for the photons physically intercepted by the detector. It is given mathematically as (Sorenson and Phelps, 1987): G = A/4&

= d2/16r2

(3)

where A is the surface area of the detector, d is the diameter of the greatest circle inscribed within the surface area of the detector and r is the distance between the source and detector. The absolute detector efficiency, L is defined as the fraction of events emitted by the source which are actually registered by the detector (Feynves and Haiman, 1969). It is described by: f=

1 _,-timxx

where p(E) is the total linear attenuation and x is the detector thickness.

(4) coefficient

2.2. The Monte Carlo simulation A typical Monte Carlo simulation begins with the initialization of all detector system parameters. Upon initialization, the simulation directs a 140 keV photon from a randomly chosen point source within the maximum solid angle subtended by the detector. As the photon approaches the detector, it will either pass through the collimator hole or strike the collimator and be absorbed. Once the photon successfully penetrates the collimator, it enters the detector where it interacts with the plastic scintillating fiber until its energy is deposited. From the nature of the interactions and the distances traveled between, the coordinates of the interactions within the detector are easily determined. The photon and electron interactions are followed until the particle energy reaches a predetermined threshold of 50 keV. At 50 keV, the tracks of both the photon and electron are terminated. The simulations performed in this research were 10,000 emission count planar acquisitions. Figure 1 displays the geometrical design of the detector system and the scintillating fiber. 2.2.1. Source distribution. The source distribution consisted of a line of 11 equally spaced (1 .Ocm) point sources of Tc-99m centered perpendicular to the collimator and detector. Calculations were made while varying the distance between the source distribution and the collimator hole, d, from 2.0 to 10.0 cm in increments of 2.0 cm. Photons of 140 keV, equivalent to the primary photon emission of Tc-99m, are emitted from a randomly chosen point source within

Optimization source Dlstrlbutlon

of a scintillating

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d

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Plnhole Collimator

Plastic

Sclntlllatlng Fiber Detector

a

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b Fig. 1, (a) Schematic diagram of the pinhole-collimated plastic scintillating fiber SPECT detector system employed in the Monte Carlo investigation. (b) The plastic scintillating fiber detector consists of linear arrays of fibers along the X- and y-axes, arranged perpendicular to the detector system or z-axis. In the detector configuration, the fibers are spaced 1.O mm between each other and 1.O mm between each array.

the solid angle subtended by the detector. The photons were transported through the detector system using direction cosines: R, = (1 - Lq)“’ x cos(f$)

Pa)

Ll, = (1 - Qz)“* x sin(d)

(5b)

R,=l-2xR

(5c)

where 4 = 2 x I[ x R and R is a randomly generated number within the range [O,l). The path of the emitted photon is determined using the direction cosines. An important point is that the photon travels in a straight line until it undergoes an interaction. The changes in position coordinates are described by: x=x,+R,xS @a)

RFC43,~F

y=_v,+R,xS

(6b)

z=z,+QxS

(k)

where x0, y,, z,, are the respective coordinates of the previous interaction and S is the mean free path length between interactions. The mean free path length of the photon is determined using the crosssections of all probable photon interactions. 2.2.2. Pinhole collimator. The pinhole collimator was constructed from lead and positioned along the detector system (z_) axis directly in front of the detector. Two collimator parameters were allowed to vary during the optimization procedure: the radius of the collimator hole, r, and the collimator thickness, t. The most important parameter in the collimator optimization procedure was the pinhole radius. A small aperture radius selectively restricts the number of photons reaching the detector resulting in a good spatial resolution but a very poor sensitivity. However, if the pinhole radius is increased, the sensitivity improves at the expense of the resolution. While

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in nuclear medicine is the plastic scintillating fiber (Hademenos et al., 1990; McIntyre, 1980a,b). The plastic scintillating fiber used in this research was composed of polystyrene doped with butyl-PBD (A,, = 430 nm) and cladded with polymethyl methacrylate (PMMA) (Blumenfeld et al., 1987), each layer possessing a characteristic index of refraction. 8 = tan-‘(2d/[) (7) As the 140 keV photon passes through the collimator, it will enter the detector and strike a scintillating where I is the length of the linear source distribution. fiber. Depending on the cross-section and subsequent The parameter, LX,is the collimator acceptance angle probability of either of the two possible physical and has been shown to produce the best or optimal interactions in this Monte Carlo simulation, the image at 70” (Anger, 1967). photon will either be photoelectrically absorbed reIn this work, the spatial resolution and scanner sulting in the emission of a photoelectron or be sensitivity consisted of collimator components using Compton scattered producing a Compton electron as the geometric collimator resolution for the pinhole well as the scattered photon. In either case, a charged collimator given by (Sharp et al., 1985): particle (photoelectron or Compton electron) is the R,,, = [(r + 4ltPe (8) product of the physical interactions. The charged particle strikes the scintillating fiber, thereby inducing where re is the effective pinhole radius: a scintillation event. re = r + 1/(2@)cot (a)) (9) In a scintillation event, light is emitted isotropically from the point of interaction. When the emitted light In equation (9), c( is the collimator acceptance angle. strikes the inner surface of the fiber, total internal The geometric efficiency of the pinhole collimator, reflection occurs for incident angles greater than the Epin, is given by critical angle, 0,, given by: E,,, = [r’ sin3(6)]/4d2 (10) sin 8, = noIn (14) The parameter, G, was previously defined as the where n is the index of refraction of the medium in geometrical efficiency of the detector or the number which the light is traveling (Femow, 1986b). In order of photons which are not intercepted and thus registo gamer the greatest amount of scintillation light tered by the detector. In this case, the collimator through total internal reflection, the fiber surface “decides” the number of photons that reach the detector. Therefore, the collimator efficiency, .I&, is should have a high index of refraction, decreasing the critical angle and increasing the range of incident substituted for the geometrical efficiency, G, in the angles that undergo total internal reflection. The top relation given for the sensitivity in equation (2). The layer, PMMA, has an index of refraction, nPMMA, system resolution for a SPECT system with a pinhole equal to 1.48. The index of refraction of the inner collimator is: layer, polystyrene doped with butyl-PBD, is np+ R s,p,n= (R; + (R,/IW)~)“~ (cm FWHM) (11) PBD = 1.60. The scintillation light is channeled through the fiber to photomultipliers coupled to the where M is a magnification factor incorporated into end of the fiber for position registration of the the system resolution to account for the magnified interaction. Figure 2 depicts these series of events in image of the object produced by the pinhole collimaa fiber once stricken by a charged particle. tor. It is equal to t/d. Thus, equation (11) becomes This scintillating fiber exhibits an attenuation R,,pi, = (Rz f (Ri (d/t))2)“2 (cm FWHM) (12) length for the scintillating photon of approximately 200 cm, a temporal resolution of 5 ns, and a threeThe overall scanner sensitivity for a pinhole-collidimensional position resolution of less than 1.Omm. mated imaging system is Spatial resolution of this magnitude is obtainable Ss.rin= 3.7 x lO’+?&,f (cps/mCi) (13) primarily due to the 0.5 x 0.5 mm2 cross-sectional 2.2.3. Scintillating fiber detector. Scintillation de- area of the fiber. Several characteristics which make plastic scintillators an attractive candidate for meditectors, especially those constructed from NaI(T1) cal imaging detector systems include: (1) a fast reand BGO scintillation crystals, have been the subject sponse time, (2) excellent spatial resolution, (3) easy of many research endeavors investigating the improvement of photon detection in medical imaging handling and manageability, (4) adaptability to any techniques (Anger and Davis, 1964; Beck, 1983; detector size and (5) cost effectiveness. Several disadvantages of the fiber in a medical imaging capacity Derenzo et al., 1988; Murayama et al., 1982; Strand and Lamm, 1980). The multicrystal detector utilizes are: (1) the high susceptibility to ageing effects from radiation and (2) that they are easily frayed. the scintillation process to image the photons from the radioactive source. Another type of scintillator In the imaging system, the detector is centered along the z-axis or detector system axis with the x-y currently under experimentation as a photon detector varying the collimator thickness and pinhole radius, images of the source distribution were acquired at increasing distances from the collimator faces. Other parameters involved in the collimator optimization are 6 and CI.The angle, 0, is the source displacement angle and varies with the source-to-collimator distance, d. It is given by:

387

Optimization of a scintillating fiber detector TOTALINTERNAL

/-ONCONE

\

the ejection of an atomic electron, termed a photoelectron. The probability of a photoelectric event occurring in the K-shell per g/cm* is given by (Fenyves and Haiman, 1969): p%)(E) = a,4,/?a4NZ5/A(m,c2/E)7~2

t \

PMMA CLADDING n=1.48

K&Kg Fig. 2. Upon striking the fiber, the charged particle induces scintillation events, creating photons which are channeled through the fiber by total internal reflection to a coupled photomultiplier tube.

surface of the detector facing the source distribution. The fibers in the detector were arranged parallel to the z-axis, as can be seen from Fig. l(a). The fibers are 1.Omm in diameter and spaced 1.Omm between each other. The dimensions of the detector are 6.0 x 6.0 x lO.Ocm. Two advantages of this type of detector are the acquisition of x and y information from each Compton interaction and the flexibility in the use of greater lengths of fiber. Thus, the primary disadvantage follows that depth information is not available from a photon interaction. 2.3. Physical interactions of the photon The photoelectric effect and the Compton effect are two ways in which photons less than 1.022 MeV energy interact with atoms and are the two processes which will be of importance in fiber scintillating detectors. Two other possible photon interactions include coherent scattering and the emission of fluorescence photons, both which have a negligible effect on the image simulations. These interactions have a dominant effect at low energies (-eV range) and since the energy threshold was defined at 50 keV, coherent scattering and fluorescence yield were not included in the Monte Carlo simulation. The differences between the photoelectric and Compton effects include the nature of the interaction and the initial energy of the photon. The photoelectric effect is an absorption process where the incident photon is absorbed, releasing an electron from an atom while the Compton effect is a scattering of the photon on a loosely bound electron of an atom. The photoelectric effect is the dominant process for photon energies ranging from the order of 1OOeV to 100 keV, depending on the atomic composition of the scattering medium. Compton scattering is the dominant process in the energy range of about 100 keV to MeV energies. 2.3.1. Thephotoelectric effect. In the photoelectric effect, the energy of an incident photon is totally absorbed by an atom. This energy transfer results in

(15)

where a,, the Thomson scattering cross-section, is (8x/W:; r,, the classical electron radius, is 2.82 x 10-‘5cm* a the fine structure constant, is Avogadro’s number, is 7.29 x ld-3; ’ N, 6.022 x 1023mol-1; Z, atomic number of detector media, is approximately 6; A, atomic mass of detector media, is 12; m,, mass of the electron, is 0.510 MeV/c*; and c, the speed of light, is 3 x 10’ m/s. The energy of the photoelectron can be determined by: E,=E,-Eb, (16) where E, is the energy of the photoelectron, Ep is the energy of the incident photon and Ek is the binding energy of the electron. For a plastic scintillator, Ebc is 64.7eV (Leo, 1987a). The differential cross-section for the photoelectron emission angle from the source for a given detector arrangement is (Barrett and Swindell, 1981): da@) -= dR

A sin*(e) (1 - /I COS(~))~

(17)

where A is a normalization constant, /? is v/c, v is the velocity of the emitted photoelectron and 0 is the angle between the photoelectron ejection direction and the axis of the acceptance cone of the detector. Since the initial energy of the photon is less than 1 MeV, a Born approximation can be applied to equation (17) (Condon, 1958): da(O) = A sin*(fI)(l + 48 cos(B)) dG

(18)

2.3.2. The Compton effect. The Compton effect is a scattering process where the incoming photon is scattered by a bound atomic electron. However, the photon energy and the energy transferred to the electron are much greater than the binding energy of the atomic electron. Thus, the atomic electron is essentially “free”. Upon scattering, the atomic electron is liberated and is scattered with an energy, E,, at an angle, 4,. dependent on the initial energy of the photon, Ep. The photon scatters at a reduced energy, E,, from the incident photon and at an angle, t+,, dependent on the initial conditions. The probability per electron for a Compton scattering event is given by: P,(E) = (NZIA)cc(E) (19) where u,(E) is determined from the differential crosssection for Compton scattering (Femow, 1986a): da/dR = (1/2)rz(@f/A*)(l + (afr*/@A))

(20)

where @ = (1 + cos*(e)), A = 1 + a,(1 - cos(@), f = (1 - cos(0)). Each of the previously defined Compton

GEQRGE J. HADEMENOS

388

scattering variables are related through the following relations (Leo, 1987b): E, = E,/[I + a,(1 - cos(O))J

one step advanced. They then correct the extrapolation using derivative information at the new point (Press et al., 1986). The predictor, in this case, is Euler’s method and the corrector is the trapezoidal method. The angular deflections of the electrons are described by the Rutherford differential scattering cross-section (Fernow, 1986a):

(21)

where a, = ErlO.511 and E, and Ep are in MeV. The energy of the Compton electron is simply the difference between the energy of the primary photon and the energy of the scattered photon. The angle of the Compton scattered photon, 0, was determined through random sampling. Once 0 was known, the emission angle of the Compton electron was determined by (Leo, 1987b): cot(&) = (1 + a,)tan(0/2)

do/dR = 4Z~Zjr~(m,c2/Bp)2(1/84)

(24)

where Z, , Z, are the atomic numbers of the incident particle (- 1) and target particle (N 6) respectively, p is the momentum of the incident particle and 0 is the scattering angle of the incident particle. The distance traveled by the electron before an interaction is also determined in a random fashion. The average distance traversed by the electron before undergoing an interaction is known as the mean free path. The electron mean free path, A.,, is defined by the equation (Leo, 1987~):

(22)

2.4. Physical interactions of the electron The generated electrons, i.e. Compton electrons and photoelectrons, undergo an energy loss and change in direction through collisions with the atomic electrons of the plastic scintillator (Jackson, 1975), leading to excitation and ionization processes. The large number of small angular deflections or multiple Coulomb scattering are a result of inelastic collisions with nuclei. The kinetic energy, dT, lost by a nonrelativistic electron as it traverses a path of length, dx, in matter is given by the Bethe formula (Williamson and Duncan, 1986):

1, = I/j@)

(25)

The electrons are exponentially attenuated as they penetrate the scintillating material and thus the distance, s, traversed by the electron between interactions is (Hademenos, 1992): s=

dT/dx = - 7.83( pZ/AT)ln( 174T/Z) (keV/pm) (23)

-1,ln[l-R]

where R is a random number.

where p, the density of the detector media, is approximately 1 g/cm3 and T, the electron kinetic energy, is measured in keV. Equation (23) is a first-order differential equation in terms of the electronic kinetic energy, T. In this simulation, the energy loss equation is handled by a predictor-corrector method. Predictorcorrector methods store the solution from each step and use those results to extrapolate the solution

2.5. Parameter optimization The optimization process of the pinhole collimator consisted of acquiring 10,000 emission count image acquisitions from a linear point source distribution while varying the source-to-collimator distance, the pinhole radius and the collimator thickness. The

200 t=lcm t=2cm t-3cm 6 d (cm)

Fig. 3(a)--legend opposite

Optimization 5000

389

of a scintillating fiber detector

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Fig. 3&c) Fig. 3. Scanner sensitivity analysis of the pinhole-collimated plastic scintillating fiber SPECT detect0 system expressed as a function of the source-&collimator distance, d, with varying collimator thicknesse of t = I-3cm and a constant pinhole radius, r, of (a) l.Omm, (b) 2.0mm and (c) 3.0mm.

source distribution was imaged at 2.0, 4.0, 6.0. 8.0 and lO.Ocm from the collimator face, while varying the collimator radius from 1.0 to 3.0 mm and the collimator thickness from 1.0 to 3.Ocm. 2.6. Imaging statistics The count distribution from each Monte Carlo simuIated image was subject to Poisson statistics.

From the properties of the Poisson distrib statistical error, 6, is expressed mathem; (Ingram and Bloch, 1984): f = loO?qJC where C is the number of counts in e Thus, for approximately 10,000 counts c each image, the error involved was 01 of 1%.

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GEQRGEJ. HADEMENOS

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1 0

2

4

6

6

10

6

10

d (cm)

a

b

6 d (cm)

C Fig. 4. Spatial resolution analysis of the pinhole collimated plastic scintillating fiber SPECT detector system expressed as a function of the source-to-collimator distance, d, with varying collimator thicknesses of f = 1-3 cm and a constant pinhole radius, r, of (a) 1.0 mm, (b) 2.0 mm and (c) 3.0 mm. 3. RESULTS

In each simulation performed by the Monte Carlo program, the detector response to the source distribution was determined numerically using the ex-

in equations (12) and (13) for the spatial resolution and the scanner sensitivity. In order to visualize eminent trends in the optimization procedure, the system spatial resolution and scanner pressions

391

Optimization of a scintillating fiber detector sensitivity were plotted against the source-to-collimator distance for each set of trials in which one parameter was varied while all other parameters remained constant. As the distance between the source distribution and the collimator increases, the sensitivity worsens, as depicted in Fig. 3. However, as the pinhole radius is increased, the sensitivity substantially improves which implies that the best sensitivity is obtained for the source distribution closest to the collimator with the largest pinhole radius. The simulation with the best sensitivity, 257 cps/MBq (7704 cps/mCi), is obtained when t = 3.0 cm and r = 3.0 mm. The spatial resolution is fairly reasonable at 0.49cm FWHM. Similar to the sensitivity, the resolution also degrades with distance, as illustrated in Fig. 4. Thus, the best results are obtained for the source distribution closest to the collimator. The resolution determined for t = 3.0 cm and r = 1.Omm is maximal at 0.43 cm FWHM but at the expense of the sensitivity which is 25 cps/MBq (747 cps/mCi). Figure 5 shows the simulated images exhibiting the optimal spatial resolution [Fig. 5(a)] and scanner sensitivity [Fig. 5(b)]. The image plot of the detector response is a two-dimensional scatter plot with the xand y-axes representing the x- and y-dimensions of the detector. The points graphed in the scatter plot are the two-dimensional position coordinates of the detected photon. The image with the best system resolution [Fig. 5(a)] is visually, a sharp, welldefined image. This type of image would be ideal in a clinical environment. However, the lack of counts within the image leaves the physician with an inadequate amount of information concerning the organ of interest to render a sound diagnosis. The image with the best system sensitivity is displayed in Fig. 5(b). Ten times more counts are observed within the image but with a slight but noticeable blurring. Since there

10

are appropriate filtering and reconstruction tech. niques to compensate for such effects, this image compares slightly more favorable to the former one 4.

In order for a detection technique to be efficienl and effective in a clinical environment, it must, firs1 and foremost, possess an excellent intrinsic spatial resolution, a high detection or quantum efficiency and a good temporal resolution in photon detection. The plastic scintillating fiber meets all of these criteria and bears further investigation into applications 01 medical imaging. The fiber arrangement employed in this SPECT detector system exhibited an optimal system resolution of 0.43 cm FWHM and an intrinsic resolution of 0.23cm FWHM. In this arrangement! the fibers had a cross-sectional area of 0.5 x 0.5 cm; and were spaced 1.Omm apart. However, with the continual improvement in existing technology, the intrinsic resolution could easily be better than 1.Omm (Ri < 1.Omm). One technological advance involves novel polymerization techniques which can produce fibers with a 0.25 x 0.25 cm* cross-sectional area. Also, under experimentation are quadratic fibers with crosssectional areas of 0.5 x 0.5 cm* and smaller as well as square and hexagonal fibers. These geometrical fiber shapes offer an optimal fit between each fiber contact. A decrease in the spacing between the individual fibers in the detector configuration will dramatically increase the resolution of the detector. Another advancement which will greatly improve the scintillating fiber in a medical imaging capacity is the newly developed solid state photomultiplier tubes (SSPMs) (Petroff and Atac, 1989). The advantages of the SSPMs include the approximately 4 times larger quantum efficiency (_ 80%) compared to standard

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Fig. 5. Images of a linear source distribution exhibiting optimal (a) spatial resolution and (b) scanner sensitivity.

10

GEORGE

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J. HADEMENOS

position-sensitive photomultipliers and about one order of magnitude shorter pulse length (N 2-3 ns) which permits the increase of count rates to a range of lo’-108 cps. The overall assessment of the plastic scintillating fiber as a detector medium for SPECT imaging is extremely promising, as evidenced by the results presented herein. A prototype of this scintillating fiber detector has been constructed in our laboratory and is currently under experimentation. Future direction in the Monte Carlo investigation of the SPECT scintillating fiber detector system include: (1) the implementation of various multihole collimators, i.e. parallel-hole, converging and diverging collimators; (2) incorporation of primary y emissions from different radionuclides clinically relevant to SPECT, e.g. Tl-201, In-l 11 and I-123; and (3) other possible scintillating fiber detector configurations and their applications in nuclear medicine. REFERENCES

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