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IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 53, NO. 5, OCTOBER 2006 2619

Cardiac Imaging Using a Four-SegmentSlant-Hole Collimator

G. Bal, E. V. R. DiBella, Member, IEEE, G. T. Gullberg, Fellow, IEEE, and G. L. Zeng, Senior Member, IEEE

Abstract—The main objective of this paper is to evaluate four-segment slant-hole (FSSH) SPECT for cardiac imaging. FSSH isa slant-hole collimator that is divided into four segments and ar-ranged such that the photons from the volume of interest (VOI)are projected four times for every location of the detector. Thesemultiple projections help to improve the sensitivity of the photonsfrom the VOI by a factor 4 cos3 , where is the slant angle ofthe collimator. Another advantage of FSSH SPECT is a reductionin the total scan time, since a gantry rotation of 2 is suffi-cient to satisfy Orlov’s condition. That means, for a slant angle of30 , a gantry rotation of 120 is sufficient to satisfy Orlov’s con-dition and obtain a complete dataset. In this paper, we evaluateand compare the reconstructed images obtained using an FSSHcollimator, for a gantry rotation of 180 and 120 , with those ob-tained from a parallel-hole (PH) SPECT system using a 180 ac-quisition. The reconstructed images from the three imaging geome-tries were compared in terms of average image noise, contrast, andpercentage error, for seven different clinical count levels and formultiple noise realizations in each case. The increase in sensitivityof the FSSH system was found to translate into a proportional de-crease in statistical noise for voxels in the VOI of the reconstructedimages. Finally, a physical phantom study was performed using aprototype FSSH collimator. Our findings show that FSSH collima-tors have the potential of being the collimator of choice for cardiacSPECT imaging. Though we explore the potential of FSSH collima-tors for cardiac imaging in this paper, the concept can be extendedfor imaging other organs such as the breasts, kidney, and brain.

Index Terms—Cardiac imaging, SPECT, three-dimensional(3-D).

I. INTRODUCTION

OUR goal in this paper is to introduce a promising newimaging geometry for cardiac SPECT and evaluate it with

respect to conventional parallel-hole (PH) SPECT using simu-lated as well as physical phantoms. The new geometry employsa four-segment slant-hole (FSSH) collimator [1] which is sim-ilar to the four-segment rotating slant-hole (RSH) collimatorsthat have been used since the 1980s for cardiac [2]–[10], brain[11]–[14], spleen [15], breast [16], [17], and renal imaging [18].Though the FSSH and RSH collimators are similar in terms of

Manuscript received August 2, 2004; revised February 25, 2005. This workwas supported in part by the Director, Office of Science, Office of Biological andEnvironmental Research, Medical Sciences Division of the U.S. Department ofEnergy under Contract DE-AC02-05CH11231.

G. Bal is with the GE Global Research, Niskayuna, NY 10239 USA (e-mail:[email protected]).

E. V. R. DiBella and G. L. Zeng are with the Department of Radiology, Uni-versity of Utah, UT 84108 USA (e-mail: [email protected]; [email protected]).

G. T. Gullberg is with the Department of Nuclear Medicine and FunctionalImaging, E. O. Lawrence Berkeley National Laboratory, Berkeley, CA 94720USA (e-mail: [email protected]).

Color versions of Figs. 6 and 8–10 are available at http://ieeexplore.ieee.org.Digital Object Identifier 10.1109/TNS.2006.877152

Fig. 1. (a) A schematic representation of an FSSH SPECT system for cardiacimaging. The arrows represent the direction of the photons that pass through theslant-holes. The projections, formed at each segment, are independent of eachother and are numbered from one to four for convenience. The axis of rotationof the gantry is parallel to the longitudinal axis of the patient. (b) The shadedportion shows the common volume (CV) while the sensitive volume is the regionfrom which projections are measured only by some segments.

field of view (FOV) and increase in sensitivity; a major dif-ference between the two lies in their imaging geometries. Forthe RSH geometry, the collimator rotates around its central axiswhile for the proposed FSSH geometry the collimator is sta-tionary and the gantry is rotated around the longitudinal axis ofthe patient, like a conventional SPECT scan.

The main motivation to study FSSH SPECT is to increasethe sensitivity of photons detected from the volume of interest(VOI) compared to that obtained using a PH SPECT system.As shown in Fig. 1, FSSH collimators are built by dividinga slant-hole collimator into four segments and arranging themsuch that the photons from the VOI are simultaneously mea-sured (projected) four times on the detector. These multiple pro-jections for every location of the gantry (view) translate into aproportional increase in sensitivity along with tomographic in-formation from the VOI [10]. However, it has been shown thatthe sensitivity of a slant-hole collimator, with a slant angle of

is times that of a PH collimator built with an equiva-lent resolution [10], [19]. For example, for a slant angle of 30the sensitivity reduces by 35% for every segment. Hence, for anFSSH collimator the net increase in sensitivity is (about160% for a slant angle of 30 ) compared to a PH collimator witha similar resolution and same acquisition time.

Another motivation to study the FSSH imaging geometry isthe reduction in gantry motion required to obtain a completedataset. As will be explained later, a gantry rotation ofradians is sufficient to satisfy Orlov’s condition. This is animprovement over PH and fan-beam geometries where thegantry needs to be rotated by 180 [20] and 180 plus the fanangle [21], respectively, to obtain a complete dataset. Similarlyfor cone-beam SPECT, a complex imaging geometry satisfyingTuy’s condition is needed [22], while for the RSH geometry thecollimator needs to be positioned at multiple gantry locationsalong with a 360 rotation of the collimator [23].

0018-9499/$20.00 © 2006 IEEE

2620 IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 53, NO. 5, OCTOBER 2006

Fig. 2. (a) The imaging geometry for an FSSH collimator on the Orlov spherefor a gantry rotation of 180 . (b) The imaging geometry for a gantry rotationof 120 . The patient axis corresponds to the z axis. I and F represent the ini-tial and final positions of the gantry. The imaging geometry made by individualsegments of the FSSH collimator is denoted as , , , and . The over-sampled regions in the equatorial arc overlap and have been shown separatelyfor illustration.

As shown in Fig. 1(b), the four segments of the collimatorand the rotation of the gantry divide the sampled image spaceinto two volumes [24]. The volume that is projected simultane-ously onto the crystal, through all segments of the collimator, forevery location of the gantry is called the common volume (CV)[Fig. 1(b)], while the volume in the image space from whichprojections are obtained for only some locations of the gantry iscalled the sensitive volume. The CV is oversampled while thesensitive volume is undersampled and can contribute towardstruncation artifacts in the reconstructed image. Thus the advan-tage of increased sensitivity and reduction in acquisition timeis obtained at the cost of a reduced FOV and susceptibility tocontamination due to truncation of the background organs withhigh tracer uptake (such as the liver for cardiac imaging).

To validate FSSH SPECT for cardiac imaging, an ensemble ofsimulated projections using the NURBS-based CArdiac Torso(NCAT) phantom [25] were generated for different tracer up-takes and for multiple noise realizations. The reconstructed im-ages were visually and quantitatively evaluated for noise, uni-formity, contrast and percentage error, by analyzing a selectedregion of interest (ROI) in the myocardium. Finally, a phys-ical phantom study of an isolated heart phantom was performedusing a prototype FSSH collimator.

II. DESIGN AND METHODS

A. Imaging Geometry

The projection through every segment of the FSSH collimatorcorresponds to a point on the Orlov sphere [26]. As shown inFig. 2, every view or gantry location, generates four such points,two of which are located on the equatorial plane which is per-pendicular to the longitudinal axis of the patient and the othertwo points are located at an angle equal to the slant angle oneither side of the equatorial plane. The two points, on the equa-torial plane, are separated by an angle equal to twice the slantangle. Thus, for a slant angle of , a gantry rotation of issufficient to satisfy Orlov’s condition, which in turn determinesif all the frequencies of the image have been sampled for an ac-curate reconstruction. Since the collimator used here was builtwith a slant angle of 30 , a gantry rotation of 120 (FSSH120)on a single head camera is sufficient to satisfy Orlov’s condition.If we assume that the acquisition time per view and the angle

Fig. 3. (a) Transaxial slice through an FSSH collimator with a slant angle of30 and a complete torso NCAT phantom drawn to scale. The untruncated CVlies within the circle drawn around the heart. (b) Coronal slice through thephantom and FSSH collimator shows that the liver is partly in the CV and partlyin the sensitive volume.

of increment between views to be constant, then the FSSH120acquisition will result in a one-third reduction in the total ac-quisition time, compared to an FSSH or PH system with a 180acquisition (FSSH180 and PH180, respectively).

Fig. 2 shows the FSSH180 and FSSH120 imaging geome-tries on the Orlov sphere. If we draw a great circle throughthese spheres, we observe that it intersects the FSSH imaginggeometries multiple times at certain regions. These multiple in-tersections mean that the FSSH geometry oversamples certainfrequencies in the VOI multiple times, thereby increasing thesensitivity of the system [1]. The effect of this oversamplingis studied in this paper by comparing the reconstructed imagesobtained using the FSSH geometries (FSSH120 and FSSH180)with those obtained using the PH180 geometry.

B. Simulated Phantom Study

A NURBS-based CArdiac Torso (NCAT) phantom [25] wasgenerated using the features of an average male torso lying in thesupine position with a raised diaphragm. The raised diaphragmcaused an increase in the attenuation of photons emitted fromthe inferior part of the heart. The concentration of Tc labeledMIBI in the heart and lungs were simulated based on the clinicaldata observed in [27]. The activity concentration in the liver wasmodeled same as that in the heart to study the worst case sce-nario of having a hot liver very close to the heart. The activity inthe blood and remaining body tissue was set at that of thelungs resulting in a relative ratio of the radionuclide concentra-tion for the heart, liver, lungs, blood and body of 26:26:1.5:1:1,respectively (Fig. 3). Note that part of the liver and lower part ofthe body lie in the sensitive volume which is not fully sampledand cannot be accurately reconstructed. Further, the detectionof counts from these undersampled regions can result in prop-agation of truncation artifacts into the CV and thereby degradethe quality of the reconstructed image.

Fig. 3 shows a transaxial and a coronal slice, drawn to scale,through the NCAT phantom for a 54 cm wide FSSH collimator.Using simple trigonometry, we observe that for a collimator sizeof 54 cm and radius of rotation of 21 cm the diameter of theCV is about 23 cm. The gantry was rotated around the phantomwith the center of the heart placed at the center of rotation of thegantry.

BAL et al.: FOUR-SEGMENT SLANT-HOLE COLLIMATOR 2621

Fig. 4. Simulated projections of an NCAT phantom after modeling attenuationand GRF for (a) PH collimator. (b) FSSH collimator.

To mimic myocardial ischemia a lesion with tracer concen-tration equal to 0.6 times that of the myocardium was addedon the lateral side of the myocardial left ventricular wall. TheNCAT phantom was discretized using a 256*256*256 imagematrix with a voxel size of 0.21 cm. The attenuation map forthe above phantom was generated using the attenuation valuesfor 140 keV photons [25].

A ray-driven projector was used to simulate the projectionof the voxelized NCAT phantom. Attenuation and the depthdependent geometric response function (GRF) were modeledin the PH as well as in the FSSH projection data [28] (Fig. 4).The collimator holes for both PH and FSSH corresponded toa LEHR collimator with similar resolution. Scatter responsewas not modeled in these studies. A 256 256 projectionmatrix with a bin size of 0.21 0.21 cm was used for eachsimulated view. The projected matrix was then resampledonto a 128 128 matrix with a pixel size of 0.42 cm. Theprojection data for the FSSH180 and PH180 imaging geome-tries were measured from 45 right anterior oblique (RAO)to 45 left posterior oblique (LPO) so as to avoid the highlyattenuated posterior projections. Sixty views (240 projectionsfor FSSH180) were collected for a gantry rotation of 180 with3 increments. The simulated projections were modeled for thereduction in sensitivity due to the slant angle by multiplyingeach pixel with the scaling factor of 0.65 [10]. Similarly, forthe FSSH120 imaging geometry, the projection data weremeasured from 15 RAO to 15 LPO. Here again 60 views(240 projections) were collected, this time with an increment of2 between views. The total acquisition time for the FSSH120projections was assumed to be two-thirds that of the FSSH180(or PH180) projections. This reduction in the acquisition timein the simulated projections was modeled by reducing the pixelvalues in the projection data to two-thirds of the original value,along with compensation for the slant angle. This resulted in ascaling of every pixel in the projection to 0.43 times the orig-inal value for the FSSH120 geometry. In spite of these scalingfactors, the FSSH180 and FSSH120 imaging geometries resultin an increase in sensitivity of 160% and 73%, respectively forthe voxels in the CV compared to a PH180 acquisition.

To evaluate the effects of different noise levels for the threeimaging geometries, the projection data, with Poisson statis-tics, were generated for seven different count levels to repre-sent varying radioactive doses administered to the patient. Theseven different count levels were simulated such that the numberof detected counts for a slice of thickness 0.42 cm through theheart varied from 60 000 to 700 000 counts for the PH180 case.

Finally, an ensemble of projection data consisting of 25 noiserealizations, were generated for each of the seven count levels.

C. Physical Phantom Study

As shown in Fig. 5, an LEGP RSH collimator was modifiedso that it can be used as a prototype FSSH collimator. The RSHcollimator was rotated such that the septa dividing the segmentsare at 45 to the longitudinal axis of the patient. The prototypeFSSH collimator was then mounted on the IRIX (Philips Med-ical Systems, Cleveland, OH) scanner. Since the crystal is rect-angular, the region surrounding the circular RSH collimator wasshielded with lead to prevent any stray photons from paralyzingthe crystal. The RSH collimator used here had a diameter of 40cm and a slant angle of 30 . However, due to the central septaseparating the segments and the large dead space (of about 3cm) between the collimator and the crystal, the effective area ofthe crystal was further reduced.

A cardiac insert (Data Spectrum Corp, Hillsborough, NC)was used to avoid any truncation artifacts in the preliminarystudy. The myocardial wall of the cardiac insert was filled with2.58 kBq/mL of Tc while the central blood compartmentwas filled with 1.18 kBq/mL of Tc. The central blood com-partment activity concentration was set to half that of the my-ocardial wall to create a bright central compartment with uni-form values and a low-contrast myocardial wall in the recon-structed images. The long axis of the cardiac insert was alignedwith the axis of rotation of the gantry. A line source was alsoplaced parallel to the long axis of the cardiac phantom. Thoughthe line source was truncated in most of the views, it served asa means to compare the resolution of the projection data. TheLEGP FSSH collimator had a resolution of 8.5 mm (FWHM)for a point source located at a distance of 10 cm from the colli-mator (compared to a FWHM of 7.1 mm for the LEHR PH colli-mator). The cardiac insert was placed in air and sixty views withan increment angle of 3 were measured for a gantry rotation of180 , resulting in 240 projections. The radius of rotation (ROR)was adjusted such that the projection from the phantom was nottruncated in any view. A circular orbit was selected as varing theROR, reduces the effective CV. A 128 128 acquisition matrixwith a pixel size of 0.46 cm was used during the acquisition.For the FSSH120 imaging geometry, the first 40 views from theabove data were used to obtain 160 projections.

Sixty views for the same phantom setup were also measured,using a LEHR PH collimator for a gantry rotation of 180 .The LEHR collimator was used as an LEGP PH collimatorequivalent to the FSSH collimator was not available. The LEGPFSSH collimator was found to be four times more sensitivecompared to the LEHR PH collimator for the same acquisitiontime. Hence, 50 000 counts were acquired for each LEHR PHview while 200 000 counts were acquired for each FSSH view.

D. Reconstruction Algorithm

An ordered-subset expectation maximization (OSEM) basedalgorithm was used in the reconstruction [29]. For the simulateddata 15 subsets with 4 views per subset were used. The effects ofattenuation were compensated for, in the reconstruction of thesimulated data. Twenty iterations of OSEM were performed for


Fig. 5. (a) FSSH image acquisition setup attached to head 1 of IRIX (Philips Medical Systems, Cleveland, OH). The frame is mounted on the gantry before thecollimator is attached. The gantry is then rotated around the longitudinal axis of the patient like a conventional SPECT system to obtain a complete dataset. (b) Aprototype FSSH collimator on the mounting frame.

each imaging geometry and the “average noise” as well as the“percentage error” in the reconstructed images were calculatedafter each OSEM iteration. GRF were not compensated for inthe simulation studies.

For the physical phantom experiments, the cardiac insert wasplaced in air and, hence, the measured data were reconstructedwithout attenuation compensation. For the FSSH180 and PH180imaging geometries, 15 subsets with 4 views per subset wereused, while for the FSSH120 imaging geometry 10 subsets with4 views per subset were used. Two sets of reconstructed imageswere created; one with and the other without GRF compensationto study the effect of resolution recovery.

E. Figures of Merit

For a thorough comparison of the reconstructed images, threefigures of merit described below were calculated after each itera-tion [30], [31]. The activity concentration in the th voxel for the

th noise realization in region is defined as . The regionin this context, is either the spherical CV containing about

23 000 voxels or a group of 450 voxels located at the center ofthe myocardial wall depending on the study. A large number ofvoxels were used in the analysis so as to obtain a statisticallysignificant comparison from twenty-five noise realizations.

The mean activity concentration in the region for the thnoise realization is defined as . Then the apparent imagenoise or roughness [30] of the image is given by

(1)

where is the number of noise realizations and is the totalnumber of voxels in the region .

To compare the different reconstructed images obtainedunder similar conditions but for different noise realizations,the mean and standard deviation images for each voxel wasdetermined. Let be the uptake concentration of the th

Fig. 6. The percentage error for all the voxels in the CV versus average imagenoise in the myocardial wall (hSi ) for the three imaging geometries. Thecurve symbols represent the iteration number at which the values were mea-sured, increasing from left to right.

voxel for the th noise realization and be the averagevalue of the voxels over all noise realizations. The standarddeviation of each voxel in the reconstructed images over therespective ensemble, was calculated for the different imaginggeometries. The value of the standard deviation of each voxelwas divided by to obtain the normalized standard deviationfor each voxel (NSDV). The NSDV can also be considered asthe uncorrelated noise of the image denoted by

(2)

Note that the apparent image noise and NSDVare different and do not measure the same noise properties of theimage. Finally, the percentage error in the mean reconstructedimage, for the different imaging geometries, is calculated as

% (3)

where is the activity concentration of the th voxel in thesimulated phantom used in the study.


Fig. 7. (a) Coronal, transaxial, and sagittal sections through the simulated phantom. Corresponding slices through the reconstructed image for the PH180,FSSH120, and FSSH180 case are shown in (b), (c), and (d), respectively. Truncation artifacts (shown by arrows) are at a distance from the VOI [indicated by thecircles in (a)], and were found to have little effect on the myocardial regions.

III. RESULTS

A. Simulation Study: Clinical Equivalent Count Levels

The plot of the ‘percentage error’ (3) versus “apparent imagenoise” for the different imaging geometries is shown in Fig. 6.It was observed that the apparent image noise is lowest for theFSSH180 imaging geometry and highest for the PH180 imaginggeometry. The reduction in noise values for the FSSH imaginggeometries is expected as they had a higher sensitivity comparedto the PH180 case.

The percentage error within the CV for FSSH180 images isthe least for a given noise level while it is the largest for PH180images. The percentage error in the CV for the FSSH120 imageis similar to that of the PH180 image. However, the noise (Fig. 6)for the FSSH120 image is less than that for PH180 image.

Since reconstructed images from FSSH SPECT are suscep-tible to truncation artifacts, we show slices in all three planespassing through the center of the heart. The truncation effectdue to the reduced FOV of the FSSH collimator can be seenin the coronal and sagittal slices, compared to the transaxialslice (Fig. 7). Reconstructed images with the lowest percentageerror were selected for this comparison. These images corre-spond to iteration numbers 5, 4, and 6 for PH180, FSSH120 andFSSH180, respectively. For this paper, the time taken per iter-ation for the PH180 case was about 1.7 min while that for theFSSH case was about 3.6 min using a dual 2 GHz power PC G5

with 2 GB of SDRAM. Hence the total reconstruction time forthe PH180, FSSH120, and FSSH180 imaging geometries wereabout 8.5, 14.4, and 21.6 min, respectively.

A reduction in noise within the CV is observed in the recon-structed images obtained using the FSSH imaging geometry,compared to those obtained using the PH180 imaging geometry(Fig. 7). However, there is no visually appreciable differencebetween the FSSH180 and the FSSH120 reconstructed images.The tradeoff of increase in sensitivity versus reduction in FOV isclear from Fig. 7. For FSSH SPECT, the increase in sensitivityresulted in a reduction in average image noise and percentageerror, for voxels within the FOV, at the expense of truncation ofthe liver and other organs lying in the sensitive volume.

To elucidate the above point, Fig. 8 shows the mean and Fig. 9shows the NSDV obtained by using the multiple noise realiza-tions. Here again, the images with the lowest percentage errorwere selected. The mean image clearly shows the truncated re-gions and the streak artifacts caused by it without the effects ofnoise. However, as these artifacts are distant from the CV, theireffects are hardly observed in the voxels within the myocardium.

It can be seen that FSSH images have a better contrast re-covery, compared to PH images in the region within the CVwhile in the region that lies outside the CV the contrast re-covery is not as good, due to the truncation artifacts. The con-trast-to-noise ratio of the myocardial wall to the lesion for the


Fig. 8. (a) The mean value of each voxel obtained using the ensemble of 25 noise realizations for the PH180 case. Corresponding slices for the FSSH120 andFSSH180 imaging geometries are shown in (b) and (c), respectively. The truncation artifacts are denoted using the arrows.

PH180, FSSH120, and FSSH180 geometries were found to beequal to 0.14, 0.34, and 0.31, respectively. The improved con-trast to noise values of the FSSH system translates to better le-sion detectability of small lesions with low contrast using theproposed method.

The right ventricular wall is not reconstructed accurately be-cause of blurring from the GRF of the system and the trilinearinterpolation used in the projector/backprojector. The effect ofreduced resolution recovery is more pronounced in the PH re-constructed images because of the relative increase in statisticalnoise in the projection data. Thus, we predict that FSSH SPECTsystems will be able to resolve smaller lesions better than PHSPECT systems.

The NSDV of the reconstructed images obtained usingPH180, FSSH120, and FSSH180 imaging geometries areshown in Fig. 9. Thus, Fig. 9 is a pictorial representation of theuncorrelated noise properties of the different imaging geome-tries in the three-dimensional (3-D) space. It is observed thatwithin the CV, FSSH120 and FSSH180 imaging geometrieshave a lower NSDV. However, for the FSSH images, the NSDVincreases drastically and sometimes has very high peaks forvoxels outside the CV. These high peaks outside the CV arecaused by undersampling of these voxels and due to the effectsof truncation artifacts. Thus, noise in the reconstructed image,to a large extent, depends on the sampling sensitivity of eachvoxel in the 3-D image. From Fig. 9 it can be seen that the 3-Dregion with a low NSDV (fully sampled), for the FSSH imaginggeometry, is relatively large and, hence, patient positioningfor cardiac imaging should be relatively simple. Further, itwas also observed that the size of the fully sampled regionfor the FSSH180 geometry is slightly larger than those of theFSSH120 geometry as shown by the arrows in Fig. 9(c). This

increase in size of the CV for the FSSH180 imaging geometrycould be due to the 240 angular sampling of the patient at theequatorial plane (Fig. 2).

B. Simulation Study: Varing Count Levels

The lowest percentage error for a given imaging geom-etry and for a given count level were selected and plottedin Fig. 10(a). The percentage errors were found to be slightlyreduced with an increase in counts in the projection data.Hence, for a given uptake value the percentage error was foundto be maximum for PH180 while it was least for the FSSH180imaging geometry.

The reconstructed image with the lowest percentage error wasselected to determine the average image noise in the myocar-dial wall [Fig. 10(b)]. The average image noise was least forthe FSSH imaging geometry while it was maximum for PH180.The average image noise was found to decrease in proportionto , where is the number of detected counts in the pro-jection data. Thus, FSSH has better noise properties than PHSPECT, especially for low count imaging.

C. Physical Phantom Study

As expected, the reconstructed images obtained usingthe LEHR PH collimator showed better resolution com-pared to those obtained using the LEGP FSSH collimator[Fig. 11(a)–(c)]. However, the resolution of the reconstructedimages, for the three imaging geometries, was comparable afterGRF compensation. The GRF for the FSSH collimator is asym-metric [28] and depends on the direction of the slant angle. Thereconstructed images with GRF compensation, correspondingto iteration number 5, 6, and 6 for PH180, FSSH120, and


Fig. 9. (a) The NSDV in the ensemble of 25 noise realizations for PH180 case. Corresponding slices for the FSSH120 and FSSH180 imaging geometries aregiven in (b) and (c), respectively. Comparing (b) and (c) we notice that the slice of the CV is slightly larger for the FSSH180 geometry compared to the FSSH120imaging geometry (shown using bars).

Fig. 10. (a) Plots the percentage error for seven different noise levels in the pro-jection (denoted in terms of counts per slice). The count levels per slice rangedfrom 60 000 to 700 000 counts per slice. (b) The average image noise in themyocardial wall versus counts per slice is shown.

FSSH180, respectively, are shown in Fig. 11(d)–(f). Attenua-tion compensation was not performed during reconstruction,resulting in the walls closer to the detector being slightly

Fig. 11. (a) Reconstructed image of the cardiac insert without GRF compensa-tion for the PH180 case. (b) The FSSH120 case. (c) The FSSH180 case. Theseimages corresponded to iteration number 4, 3, and 3. Reconstructed image withGRF compensation for (d) the PH180 case, (e) the FSSH120 case, and (f) theFSSH180 case.

brighter than the other walls. Further, flood corrections couldnot be performed on these data as the images were acquired in aspecial override mode resulting in artifacts in the reconstructedimage. This paper shows proof of concept for FSSH imaging,and further work needs to be done to optimize the imagingparameters.

IV. DISCUSSION

In this paper, we studied two novel FSSH imaging geometries(FSSH120 and FSSH180) for cardiac imaging and evaluatedthem with respect to PH180 SPECT using simulation studies aswell as a physical phantom experiment. The simulation studieswere performed for a wide range of clinically realistic countlevels.

It is well known that lesion detectability helps to evaluatethe presence as well as severity of the cardiovascular diseases.


Lesion detectability, in turn, depends on the size of the lesion,contrast ratio, percentage error and noise in the reconstructedimage. Noise gives the reconstructed image an undesirable ap-pearance. Thus, reducing noise in the reconstructed images canincrease the contrast sensitivity of the system thereby improvingits diagnostic capability. As the percentage error and noise in thereconstructed image for the FSSH imaging geometries are lessthan PH180 (Fig. 6) it becomes easier to resolve lesions that aresmaller in size, or to detect lesions with lower contrast.

It is known that lower the percentage error for a given noiselevel means the reconstructed image, obtained for a particularimaging geometry, is more accurate, since it is closer to the orig-inal phantom. As shown in Fig. 6, the lower error for FSSH180results from the uniformity of the myocardial walls (Fig. 7)and the reduced NSDV in the CV of the reconstructed images(Fig. 9). We postulate that the uniformity in the myocardial wallsis due to the fact that the FSSH180 projections were acquiredover a wide angular range, resulting in a larger CV, hence, re-ducing contamination from distant truncated regions. For ex-ample, though the gantry for the FSSH180 imaging geometryis rotated from 45 RAO to 45 LPO, the projections on theequatorial plane are actually measured for about 240 , rangingfrom 75 LPO to 75 RAO (Figs. 2 and 3). The upper and lowersegments of the FSSH collimator measure oblique projectionsfrom 45 LPO to 45 RAO. As a result, sufficient projectionsfrom the liver and other regions surrounding the heart are mea-sured, enabling an accurate reconstruction of a larger CV. It mayalso be noted that the four projections measured for every view,correspond to four different planes passing through the zero fre-quency in the Fourier domain [1]. These projections, measuredfrom multiple planes, result in a finer sampling of the CV inthe frequency domain, and an overall reduction in error for theFSSH180 imaging geometry.

It can be seen from Fig. 9 that, the CV for the FSSH120 ge-ometry is smaller than that of the FSSH180 geometry and themyocardial region appears closer to the undersampled regionin the reconstructed image. However, the NSDV within the CVfor the FSSH120 geometry is much less than the NSDV for thePH180 case showing that reconstructing images obtained usingthe FSSH120 imaging geometry contained less noise but similarpercentage error as the PH180 case (Fig. 6). Similar percentageerror was observed in spite of the fact that the simulation studieswere modeled using a liver with a large tracer uptake placedvery close to the myocardium. As shown in Fig. 7–9, the recon-structed FSSH images did not show any streak artifacts or degra-dation in terms of contrast and noise for voxels inside the CV.These results put to rest one of our main concerns regarding thepropagation of the truncation artifacts into the CV for the FSSHgeometry.

Thus, we believe that as long as the cardiac region is con-tained within the CV the propagation of truncation artifacts isgoing to be insignificant. Hence, it is important to position thepatient such that the heart lies in the CV during clinical studies.This positioning of the heart should be relatively easy, comparedto other imaging modalities such as RSH [23], fan-beam [21]and cone-beam [22], as FSSH has a simple scanning protocol.

However, further studies need to be performed to determinethe effects of truncation artifacts on lesion detectability, espe-

cially when the heart is partly truncated in some of the pro-jections. Similar studies were performed for RSH SPECT [32],using physical phantoms as well as animal studies for variousdegrees of truncation of the heart. It was shown that when theprojections from the heart are truncated, then truncation artifactscan be significantly reduced by extrapolating the projection datafor the different segments.

Finally, as shown in Fig. 3, increasing the slant angle bringsthe CV closer to the detector, thereby improving the resolutionof the voxels in the CV. Increasing the slant angle also decreasesthe degree of the gantry rotation needed to satisfy Orlov’s condi-tion thereby reducing acquisition time. However, increasing theslant angle also reduces the size of the CV and decreases the sen-sitivity of the system [10]. Thus, based on the tradeoff betweenacquisition time, resolution, size of CV and sensitivity, an op-timal slant angle has to be determined, depending on the appli-cation. For example, in dynamic cardiac imaging, the tissue up-take of the radioisotopes needs to be measured rapidly resultingin high statistical noise in the reconstructed image. Hence, if weuse the FSSH120 imaging geometry then the increase in sensi-tivity, decrease in radius of rotation and the reduction in acquisi-tion time, along with enhanced tomographic information in eachview, will result in faster sampling of the VOI compared to theFSSH180 and PH180 geometry, thereby making it suitable fordynamic cardiac studies.

V. CONCLUSION

This paper shows that the main benefits of an FSSH system,compared to the PH180 system for cardiac imaging are: (1)an increase in sensitivity of 73% and 160% for the FSSH120and FSSH180 geometries, respectively, resulting in a reductionin the average image noise within the CV in the reconstructedimages; (2) reduced percentage error, even in the presence ofsignificant background activity in the sensitive volume; (3) en-hanced tomographic information from the VOI in every view;and (4) reduction in image acquisition time. On the other hand,drawbacks of the FSSH system include: (1) reduced FOV; (2)possible contamination from truncated background organs; and(3) an increase in computing time required for reconstructing allthe projections. However, as shown in the simulation studies, theFOV was found to be large enough for cardiac imaging and thetruncation artifacts had negligible effects on the voxels withinthe FOV. Thus, improved cardiac images can be obtained withthe addition of a relatively inexpensive FSSH collimator to anexisting SPECT system.

ACKNOWLEDGMENT

The authors would like to thank Dr. S. Dale of the KarolinskaInstitute, Stockholm, for loaning the FSSH collimator to us, Dr.R. Clackdoyle and P. Christian of the University of Utah, M.Smith of 3DX, and Dr. D. Gagnon of Philips Medical Systems,Cleveland, OH, for helping us with the mounting of the colli-mators on to the scanner, and the anonymous reviewers for theircomments in improving the paper.

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