to move in synchrony with the opposing fan-beam collimator. Data from transmission and emission sources at different energies were acquired in one detector, ...
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Grant T. Gullberg, Senior Member, ZEEE, Hugh T. Morgan, Member, ZEEE, Gengsheng L. Zeng, Member, ZEEE, Paul E. Christian, Edward V. R. Di Bella, Member, IEEE, Chi-Hua Tung, Member, ZEEE, Piotr J. Maniawski, Yu-Lung Hsieh, and Frederick L. Datz
Abstract-A commercial three-detector single-photon emission computed tomography (SPECT) system that enables simultaneous acquisition of transmission and emission data without increasing patient scanning time has been designed and manufactured. This system produces a reconstructed attenuation coefficient distribution that can be used to correct for photon attenuation in the emission reconstruction. The three detectors with fan-beam collimators are mounted to the gantry in a triangular arrangement. A transmission line source assembly was mounted at the focal line of one of the detectors and controlled to move in synchrony with the opposing fan-beam collimator. Data from transmission and emission sources at different energies were acquired in one detector, while the other two simultaneously acquired emission data. A transmission source of 153 Gd was used with 99m Tc-labeled radiopharmaceuticals, and 57 Co was used with '"T1. Algorithms were developed to subtract crosstalk between transmission and emission energy windows in all three detectors. A transmission maximum-likelihood iterative algorithm was used to reconstruct the attenuation distribution, which was used in combination with an iterative maximumlikelihood expectation-maximization algorithm to compensate for the attenuation of the projection of the emission distribution. The results in phantom studies displayed greater uniformity of activity with attenuation-corrected reconstruction. This was demonstrated visually and quantitatively by using anterior-toinferior ratios close to one and low spatial %rms error as a measure of improved uniformity. Index Terms- Anger gamma camera, attenuation correction, brain SPECT, cardiac SPECT, clinical imaging hardware, clinical imaging software, emission computed tomography, fan-beam tomography, iterative reconstruction, simultaneous transmission and emission imaging, single-photon emission computed tomography, SPECT, transmission computed tomography.
I. INTRODUCTION HOTON attenuation in cardiac single-photon emission computed tomography (SPECT) is the major factor that Manuscript received June 3, 1996; revised January 20, 1997 and January 5, 1998. This work was supported in part by the NIH under Grant R 0 1 HL 39792 and Picker International. G. T. Gullberg, G. L. Zeng, P. E. Christian, E. V. R. Di Bella, and F. L. Datz are with the Department of Radiology, University of Utah, Salt Lake City, UT 84132 USA. H. T. Morgan and P. J. Maniawski are with Picker Intemational, Inc., Cleveland, OH 44143 USA. Y.-L. Hsieh was with the Department of Radiology, University of Utah, Salt Lake City, UT 84132 USA. He is now with G E Medical Systems, Milwaukee, WI 53201 USA. C.-H. Tung was with the Department of Radiology, University of Utah, Salt Lake City, UT 84132 USA. He is now with Picker Intemational, Inc., Nuclear Engineering, Cleveland, OH 44143 USA. Publisher Item Identifier S 0018-9499(98)04322-6.
contributes to the quantitative inaccuracy that results when SPECT is used to measure in vivo distribution of radioactivity. A commercial three-detector SPECT system has been designed and manufactured to enable simultaneous acquisition of transmission and emission data so that the reconstructed transmission map can be used to correct for photon attenuation in the emission reconstruction. The three detectors with fan-beam collimators are arranged in a triangle set with a transmission line source located at the focal line of one of the detectors. Transmission and emission data are acquired by one detector, while the other two detectors acquire emission data. The triangular arrangement presents a unique geometric configuration that utilizes a fan-beam transmission geometry with detectors positioned carefully in order to avoid truncation of the incident transmission flux by the detectors adjacent to the line source. However, due to the limited field of view of the fan-beam collimated detector, it is difficult to acquire fan-beam transmission data through the thorax without truncating the projections in most views. Still, even with the transmission data truncated it is hypothesized that, through the use of iterative maximum-likelihood (ML) algorithms to reconstruct both transmission and emission data, more uniform attenuation-corrected reconstructions are obtained with a precision determined by the emission statistics for those levels of statistics commonly used in clinical applications. A variety of transmission computed tomography (TCT) design concepts have been proposed for commercial SPECT systems. The designers of these systems have struggled with numerous decisions including: whether to use sequential [l]-[6] or simultaneous transmission and emission 171-11 11 imaging; whether to use parallel 131, [5], fan-beam [6], [SI, [11]-[13], or cone-beam 1141-[16] collimation; whether to use single detector [3], [5] or multiple detector [6], [SI systems; whether to employ flood [21, [31, El, scanning line [71, 1101, fixed point [141, [151, or fixed line 161, 181, [111-[131 sources; and whether to use an X-ray tube [17] or gamma-ray transmission sources [1]-[16]. In the first studies, the transmission study was performed before the emission study using a flood source mounted opposite a single parallel-collimated detector [3], [5]; unfortunately, this arrangement increased patient scan time. Efforts were then made to reduce scanning time by simultaneously acquiring transmission and emission data [7], [9], [lo]. In one attempt, flood transmission source and emission data were collected simultaneously from a single parallelcollimated detector [9]. The disadvantages of this approach
0018-9499/98$10.00 G) 1998 IEEE
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were the contamination of transmission andor emission data from crosstalk of transmission and emission sources and the increased bulk and weight that resulted due to the plane source. An additional disadvantage of parallel collimation was poor transmission statistics, which were improved by placing the transmission source at the focus of a converging collimator, such as a cone-beam or fan-beam collimator. Another study [ 151 showed that improved image quality could be obtained by using a cone-beam collimator and a transmission point source with activity as low as 3 mCi. However, the disadvantage of converging collimation is truncation of transmission data with typical 40-cm-wide gamma cameras. A scanning line source with a parallel collimator reduces the truncation problem experienced with converging collimation [16]. By electronically blanking the detector and zeroing in on the transmission energy, these systems can reconstruct distributions of narrow beam attenuation coefficients. Some manufacturers have proposed the use of two scanning line sources opposite two parallel collimated detectors orientated at 90" to each other. These systems have approximately the same sensitivity as a three-detector system with fan-beam collimators. For small patients, a 90" rotation of each detector may be all that is required, whereas 180" rotation may need to be used for large patients in order to avoid truncation. However, the implementation by a motor-driven scanning line source and electronic blanking of camera electronics can be more complicated, as well as more expensive, than using a fixed line source. The parallel collimator also has lower sensitivity, thus requiring increased source strengths. The three-detector SPECT system offers the advantage of simultaneous transmission and emission imaging using a light weight, low-dose line source positioned at the focal line of one of the fan-beam collimated detectors. The system also reduces crosstalk between emission and transmission data energy windows. However, the triad arrangement of detectors dictates that fan-beam collimators of a short focal length (65 cm) must be used so that the incident transmission flux is not truncated by the detectors adjacent to the line source as the detectors move close to the patient during the scanning of body contouring orbits. The use of a fanbeam geometry with 40-cm field-of-view detectors truncates transmission projections. This is especially the case when imaging projections of the long axis of the body. To reduce the truncation of the transmission projections, Jaszczak et al. [6] have used longer focal length fan-beam collimators (110 cm). However, it is necessary to perform sequential transmission and emission scans with this technique so that during the transmission scan the detectors adjacent to the line source can be retracted to avoid truncating the incident flux. Ideally, simultaneous imaging requires greater emission energy than transmission energy, but this is not always possible. In the work presented here, a 153Gd (95.1-103.2 keV, 51%; 40.9-54.7 keV, 147%) transmission source was used with 99mTc (140 keV) to mimic gg"Tc-labeled radiopharmaceuticals. Additionally, a 57C0 (122 keV, 86%; 136.5 keV, 10%) transmission source was used with "'Tl (68.9-80.3 keV, 94.4%; 135 keV, 2.7%; 167 keV, 10%). Other groups [11] have investigated the use of the 59.5-keV energy of 241Amas
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a transmission source in order to reach below the 68.9-80.3 keV energy of '"T1. Transmission and emission data were simultaneously acquired in one detector, while at the same time the other two acquired emission data. All three detectors acquired data from an acquisition window centered on both the transmission and the emission energy. Transmission and emission data from all three detectors were corrected for crosstalk before reconstruction. The corrected transmission projections were reconstructed by using a transmission iterative ML algorithm to provide the reconstructed attenuation distribution. The reconstructed attenuation map was scaled to the correct emission energy. The reconstructed attenuation distribution was used to calculate attenualion factors that can be used to model the attenuation of the emission projection data as a large system of linear equations. This system was solved by using an iterative maximum-likelihood expectation maximization (ML-EM) algorithm to produce the attenuationcorrected radiopharmaceutical distribution. This paper includes descriptions of the essential hardware and algorithms delivered with the clinical three-detector simultaneous transmission and emission SPECT system and evaluates its performance using data obtained from phantom studies. Details are presented of both preprocessing algorithms used to eliminate the data contamination between transmission and emission data acquisition windows arid transmission and emission reconstruction algorithms. Both profile and bull'seye plots are used to evaluate the results qualitatively in terms of uniformity of uptake between inferior and anterior superior walls and between lateral and !septal walls, much like a physician does in diagnosing defects in cardiac patient images. The average anterior-to-inferior ratio and the spatial %rms error over the region of the lateral or septal wall are also recorded before and after attenuation correction to quantify improvement in distribution uniformity. 11. SYSTEM DESCRIPTIOIV A. STEP Imaging System A PRISM 3000XP SPECT system' was modified as shown in Fig. 1 to perform the simultaneous transmission and emission data acquisition. The primary purpose of the Simultaneous Transmission Emission Protocol (STEP) system is to provide improved SPECT image quality through accurate body attenuation correction. It should be noted that all three detectors collect emission data, and detector 3 (not seen but identified with three colored circular discs on the side of the camera) also collects primary transmission data. Two configurations of STEP have been incorporated into the design of the PRISM 3000XP system: a head configuration for brain imaging and a torso configuration intended primarily for heart imaging. These two configurations consist of different colliimators and slightly different software. The STEP hardware consists of a source drive assembly, a source holder assembly, and a set of three fan-beam collimators. The STEP software consists of the special acquisition and image processing software necessary to acquire and process a Picker International, Cleveland, OH.
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Fig. 1. The PRISM 3000XP STEP system. The system has three fan-beam collimated detectors with line source and source drive assembly shown at the top positioned between two of the detectors. The source drive assembly keeps the line source at the focal line of the fan-beam collimator mounted to the opposite detector.
IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 45, NO. 3, JUNE 1998
Fig. 2. Transmission source holder assembly with collimator and external attenuator. This locks into the source mounting block shown in Fig. 1. The small shaft extending from the back controls the shutter position and fits into the circular slit shown in Fig. 1. The open, closed, and calibration positions are indicated on the front face. The collimator and attenuators are made of tin instead of lead to avoid downscatter of lead X rays into the 8U-keV 201T1 energy window. Thinner tin attenuators are inserted as the source decays over time in order to maintain a constant transmission flux.
when the shutter is in the closed position. The cylindrical shutter rides on needle bearings. The means to open and close it is a solenoid-actuated rack and pinion mechanism. The STEP study. The acquisition software simultaneously acquires calibrate position of the shutter may be manually selected. two energy windows per detector. The processing software Selection places a precision thickness tin filter in the beam receives and preprocesses these data to eliminate spectral path. The calibrate position is used to acquire incident flux data crosstalk; it then reconstructs both transmission and emission that enables calibration of the transmission scan. A dial on the data. I ) Source Drive Assembly: The primary purpose of the front face of the source assembly indicates the position of the source drive assembly is to move the source assembly so that source shutter; it can also be used to manually move the shutter it remains at a fixed distance from the opposed detector head to any of its three positions (open, closed, or calibrate). During during a noncircular SPECT orbit. This makes it possible normal system operation, this shutter is under software control for the line source to remain fixed at the focal line of the and is opened only when the camera is acquiring transmission fan-beam collimated detector as the detector traverses a body- data. Another important component of the transmission source ascontouring orbit. This assembly is permanently mounted to the gantry rotation plate as shown in Fig. 1. Included in sembly is the source collimator, which is shown in Fig. 2. This this assembly is a drive motor and associated hardware that collimator confines the transmission beam to the size and shape can drive a source mounting block through 20 cm of radial of the opposed detector and improves transmission resolution travel and a solenoid-actuated shutter opening and closing in the axial direction. This collimator also reduces scatter of mechanism. The source drive assembly is interfaced through a the transmission beam. The V-shaped (or fan-shaped) body of motor control board to the gantry microprocessor board where the collimator is cast in lead; the 0.5-mm-thick 44-mm-long tin alloy vanes are spaced 2.5 mm apart. Sandwiched between its actions are controlled by software. 2 ) Transmission Source Holder Assembly: The STEP the collimator face and the electromechanical collision sensor, transmission source assembly is an electromechanical device shown on the bottom face in Fig. 3, is an interchangeable that safely houses a radionuclide line source. A photograph tin filter. Three thicknesses (1, 0.5, and 0.25 mm) have been of the complete STEP source assembly is shown in Fig. 2. designed for use with the system. The tin filter removes most of This assembly measures 1.5 in wide by 4.5 in high by 13.5 the lower energy emission of 153Gdand also undesirable lead in long and weighs 10.2 pounds. It is manually attached (and K-X rays emanating from the source holder and collimator unattached) to the drive assembly via a rotating “lock on” body. Replacing the thicker tin filter with a thinner one increases the transmission flux, thus extending the usable life ring mechanism. When installed in the source assembly, the line source is of the decaying line source. The judicious use of tin in the surrounded by the source assembly, source holder, and shutter source assembly design reduces lead K-X-rays contamination as shown in Fig. 3. The line source is shielded by the lead in the transmission beam to a negligible amount. In patient source holder and the lead beam stop of the source shutter studies the gross count rate for detector 3 remains under 10
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AIR WINDOW
Radiation on Imaging Position
Radiation on Calibration Position LEAD WINDOW
Radiation Off Position Fig. 3. Schematic of the source holder assembly. At the top it is shown in cross section in the closed position. The inner lead shield rotates 90' to the right to expose the line source to the tin window for calibration and rotates 90" to the left to expose the line source to the air window for imaging purposes.
K cps, which produces minimal count losses. Corrections are not necessary for this small amount of count loss. 3) Selection of Transmission Sources: Three radionuclide line source options have been developed specifically for STEP and are included in the system design. These sources are 153Gd (95.1-103.2 keV, 51%; 40.9-54.7 keV, 147%; T 1 p = 240 days), 57C0 (122 keV, 86%; 136.5, 10%; T1p = 272 days), and Q g m T(140 ~ keV, T112 = 6 h). The active and physical length of the line source is 235 and 267 mm, respectively, and the active and physical diameter is 5.3 and 6.3 mm, respectively. Typically, the source uniformity has a variation of less than &3%. In our studies, both 153Gd and 57C0 line sources were used with activities of 60 and 8 mCi, respectively. When acquiring transmission and emission images simultaneously, the energies of the transmission and emission radionuclides should be sufficiently different so that their acquisition energy windows do not overlap. Ideally, it is better to choose a transmission source with a lower photopeak energy than that of the emission source so that the emission is not contaminated by downscatter from the transmission source. However, this may not be possible with 'OlTl unless a transmission source such as 241Am (59.5 keV) [11] can be used. In the STEP system, when emission imaging a zolTl agent at 73 keV, either 57C0 or g g m Tmay ~ be used as the transmission source. When emission imaging a 9 9 m T agent, ~ 153Gd at 99 keV is used as the transmission source. The three STEP transmission sources can easily be installed and removed from the source assembly by the operator with a specially designed tool that is provided by the manufacturer. The energy spectrum of "'Tl
is 94.4% 68.9-80.3 keV photons, 2.7% 135-keV gamma rays, and 10% 167-keV gamma rays per disintegration. Only the 68.9 to 80.3-keV photons are used in the STEP system. The other photons, which are registered in the transmission energy window, are eliminated through calibration measurements by the emission-only acquisition detectors. 4) Fan-Beam Collimation: Fan-beam collimators are used on all three heads of the STEP system. For the STEP torso configuration, each collimator has a focal1 length of 65 cm which produces a 240 x 325-mm field of view at 10 cm from the collimator surface. The system resolution of these individual collimators is 3.5 mm at the surface and 8.0 mm at 10 cm distance, respectively. The system sensitivity is 285 cpm/pCi at 10 cm. These measurements were obtained by following the 1994 NEMA standard measurement procedures. A shorter focal length would limit the field of view too much; a longer length would result in interference between the transmission beam and the adjacent detectors for small body contour orbits. For the STEP head imaging configuration, 50 cm is the preferred focal length. Fan-beam collimation on the STEP system enables the use of a relatively low activity line transmission source placed near the focal length of the opposed collimator. The use of fan-beam collimators causes truncation of the transmission projections of the chest; the key to compensating for the truncation is in the design of the reconstruction software. 5) Acquisition Sofhvare: The STEP hardware design described above is interfaced into the PRISM system and controlled by the STEP acquisition software. The operator selects a STEP SPECT study from the acquisition menu, then the soft-
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I
Crosstalk Correction
I
iterative Reconstruction
Reconstruction
EMISSION IMAGES Fig. 4. Flow diagram for the STEP data processing software. First, the raw data is corrected for crosstalk between the various photon sources. Then, the transmission data is reconstructed to produce an attenuation distribution; using this reconstructed attenuation distribution, the emission data is reconstructed to give attenuation-corrected emission reconstructions.
ware checks for the presence of the STEP hardware. If found, the software prompts the operator to define a patient-specific orbit and to set up the simultaneous transmission-emission protocol. STEP acquisitions are performed by opening two energy windows on each of the three imaging heads. One window is set for transmission energy, the other for emission energy. The acquisition software controls the acquisition by moving and monitoring the system. The STEP source is moved so as to keep it at the opposed fan-beam collimator's focal distance within *l mm as this opposed head orbits around the body. The acquisition software also controls the opening and closing of the source shutter so that the patient is only exposed to transmission radiation while the opposed head acquires transmission data. The data from each of the six energy windows (two from each detector) are stored in separate data files and passed to the system image processor for processing and reconstruction. The header file contains orbit information used by the algorithm to define the support of the body attenuator as described in Section 11-B3.
B. Image Processing Sojhare
The basic software module flow diagram for STEP image processing is presented in Fig. 4. The processing software first preprocesses the acquired data to eliminate spectral crosstalk, then the reconstruction software reconstructs the transmission data. Transmission reconstruction is used to calculate attenuation factors that are used in the reconstruction of the emission data to correct for photon attenuation.
1 ) Data Processing Software (Cross-Talk Correction for 57C0/L01Tl): Three kinds of crosstalk occur in sihultaneous 57Co/201Tl transmissionlemission imaging. Transmission data acquired from the 122-keV window of detector 3 are contaminated by the 135- and 167-keV photons from 'O'Tl emissions, and emission data are contaminated in two ways by transmission photons (122 keV) that downscatter in photon energy. As shown in Fig. 5, the geometric positioning of the three detectors results in backscatter of transmission photons from the patient detected in the 73-keV energy windows of detectors 1 and 2, adjacent to the line source. Also, as shown in Fig. 6, some transmission photons will forwardscatter and be detected in the 73-keV energy window of detector 3 . The primary forwardscatter in detector 3 is caused by K-X rays from the lead collimator. The contamination of transmission data by the high energy photopeaks of 201T1 is removed based upon measurements from emission-only acquisition detectors. An estimate of the 135 and 167-keV photons of 201T1detected in the 122 keV transmission data window of detector 3 is obtained from data acquired in the 122 keV windows of detectors 1 and 2. An estimate is obtained by taking the average of the 122keV data acquired from detectors 1 and 2 at the same view. The estimated "'Tl photon contamination is subtracted from the transmission data of detector 3 before reconstructing the distribution of attenuation coefficients Ti(corrected) = Ti(measured) - [Ti(measured)
+ T i (measured)]/ 2
(1)
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Fig. 6. Forwardscatter. The contamination of emission projection data acquired from the simultaneous detector due to downscatter of 99mTc photons and Pb X rays (interaction of 140-keV photons and lead collimator) into the 201T1 energy window is illustrated.
Detector #2
Detector # l
Tianv”ion Source Holder A\\emhiy Y
(b) Fie. 5. Backscatter. (a) Illustration of the contamination of emission uro. jezion data due to the singly or multiply scattered 99mTc photons into-the ZolTl energy window when the transmission is at position A, (b) Backscattered photons when transmission source is at position B.
where T,” is the transmission window data for detector n at angle 0. This corrected transmission projection data, along with corrected incident flux data, are processed by the transmission reconstruction algorithm. The incident flux is measured through a calibrated filter, which is corrected for filter attenuation, decay, and detector dead time as needed. The correction for backscatter is based on the fact that the intensity of the detected backscatter decreases rapidly moving in a direction perpendicularly away from the line source, as illustrated in Fig. 7. When the line source is positioned at A [Fig. 5(a)] and B [Fig. 5(b)], the emission data sampled from detector 1 for position A, and detector 2 for position B, are equal. However, the detected transmission backscatter decreases in intensity from opposite ends of the data array. It is observed that in most of the projection data the profile
Fig 7 Measure backscatter in a transmission-only phantom study The intensity of the crosstalk decreases rapidly in the direction away from the transmission line source
of the backscatter photons decreases rapidly in the direction away from the line source so that the scatter profile in the data array at position A does not overlap with the profile in the data array at position B. Therefore, a simple technique can be used to remove the detected backscatter transmission photons. Each step of the correction procedure is illustrated in Fig. 8. First, the emission projection measurements from the two detectors are added, then the emission data from one of the detectors (for example, detector 3 at position B) are subtracted from the emission data from the other detector. The corrected emission data are obtained by determining the difference between the summed image and the absolute value of the difference image. This method requires obtaining two projt:ction samples for each projection angle with the position of the line source at opposite ends of the detector. As implemented, STEP requires 360” of data acquisition per detector. For another proposed method, see [18]. The outlined procedure corrects for the backscatter found in the detectors adjacent to the line source. The forwardscatter
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subtracted from transmission data in detector 3
T i (corrected) = T i (measured) - [T' (measured) + T i (measured)]/2. (3)
STEP 1: PA+PB 1 2
STEP 2: IPA-PB 1 21
STEP 3: PA+PB - PA PB 1 2 1- 21
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This simple correction has the potential of introducing less variance, thus making it more desirable to use a transmission source of a lower energy than that of the emission when executing simultaneous transmission and emission acquisition. However, the key consideration is whether the transmission and emission energy combination eventually result in less variance and bias of the corrected emission distribution. 3) Reconstruction Sofoivare Transmission (Truncation Correction): The magnification of the fan-beam geometry produces truncated transmission projections of the chest (Fig. lo). A 65-cm focal length is the optimum distance for minimizing projection truncation without truncating the incident flux during simultaneous transmission and emission acquisition. The field of view is approximately 25 cm at the center of rotation for a radius of rotation of 25 cm. When the patient is placed in the center, this amounts to approximately 25% o f the patient area being truncated for a 38 by 26-cm elliptical patient. If the patient is positioned so that their heart is in the center, approximately 50% truncation results. The transmission reconstruction is formulated such that the distribution of the attenuation coefficients is determined from the solution to the system of linear equations for only measured projections. This means that the truncated projections are not extrapolated or set to zero beyond the truncated region. The algorithm is a variation of the one proposed by Lange et al. [191. For each iteration n 1,the attenuation coefficient for pixel i is calculated using
I
Fig. 8. Backscatter compensation procedure. The three steps used to compensate for crosstalk are illustrated. The emission data for detectors 1 and 2 are measured at the same angular position.
found in the 13-keV energy window of the detector (detector 3) opposite the line source is eliminated by using a method based upon work of Frey and Tsui [9]. The forwardscatter is composed primarily of lead K-X rays; the profile illustrated in Fig. 9 is similar to that of the transmission data. A fraction f~ of the crosstalk-corrected transmission photons Tf (corrected) in each projection bin (the 122-keV window) is subtracted from the measured projection data E; (measured) for the same bin measured from the 13-keV emission window
+
I
(4) / E m WimITn (1 - w ) where pt is the attenuation coefficient for pixel i , IO,^ is the E;(corrected) incident, and I , is the detected number of photons for projection bin m. Both Io,, and I , are Poisson-distributed random E;(measured) - f~ x Ti(corrected), if >O f~ x E;(measured), otherwise. variables. The factor wzm is the length of the intersection, with (2) pixel i of the ray extending from the focal point to the center of projection bin m. This weights the contribution of each pixel to the projection by the length of the line that intersects the The fraction f~ = max{Ei (measured)}/ max{T; pixel. The sums over all projection bins m that contribute to (corrected)}, where the maximum is taken over all projection bins over all projection angles. For another proposed method pixel i in both numerator and denominator are backprojection operations. The factor w is a relaxation factor that depends on of calculating f ~ see , [8]. 2) Data Processing Soforare (Crosstalk Correction for the total counts: 153GdpgmTc):The correction for crosstalk found in siw = min(tota1 counts x 4 x IO-', 0.9). (5) multaneous 153Gd/ggmT~ transmissionlemission imaging is The algorithm is constrained to reconstruct the attenuation much easier than that of simultaneous 57Co/201Tlimaging. In correcting crosstalk of simultaneous 153Gd/9gmT~ imaging, map within a finite support. An estimate of the support is only downscatter of the 140-keV photons from 99mTc obtained from the detector radius information provided by the emissions into the transmission data of detector 3 [Ti system acquisition software. Since the camera moves near the (measured)] must be considered. First, an average of emission body during noncircular orbit acquisition, the detector radius downscatter in detectors 1 and 2 at each angle 0 is determined information can be used to approximate the body contour as from T i (measured) and T i (measured). This average is follows. First the image array is set to unity. At each detector
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20 200
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Fig. 9. Measured primary photons and forwardscatter in a transmission-only phantom study: (a) primary transmission photons and (b) fonvardscatter in the simultaneous detector (detector 3 ) . Horizontal profiles are drawn through the measured projections as indicated by the solid lines.
(a)
(b)
Fig 10 Truncation of transmission data (a) The incident flux is truncated by the adjacent detectors (b) The transmission projecbons of the patient are truncated This is intended as schematic However, if we assume a 40-cm camera face, the focal length is approximately 45 cm, and the long axis of the patient is approximately 29 cm
position, a straight line is drawn on the image array 3.8 cm in front of, and parallel to, the detector imaging plane. All pixels behind this line (i.e., toward the detector) are set to zero. This is repeated for every detector position. The resultant image array has a support with a value of one within the convex boundary of the inscribed contour and pixels outside have a value of zero. This is used as the initial solution for the reconstruction of the attenuation distribution. Because the algorithm is multiplicative, the zero values constrain the reconstruction to the support.
In the STEP method there is no smoothing of the transmission reconstruction before the attenuatjon coefficients are calculated. For the emission quantities aldministered in the clinic, the variance of the attenuation-correctedemission reconstruction is not influenced by transmission statistics if the incident flux is kept above a minimum of 750 counts per projection, but instead is determined by the emission counts [20]. To add any filtering step to the transmission reconstruction would only add potential bias to the corrected emission reconstructions.
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Because the reconstructed attenuation coefficients are energy-dependent, it is necessary to transform the reconstructed attenuation coefficients from the transmission source energy measured to the emission source energy. The conversion used is based on the assumption that the coefficient varies linearly for different attenuating materials as follows
~51:
where ,LLE and ,LLTare the emission energy and the transmission energy linear attenuation coefficients, respectively. The conversion factors for the STEP transmission and emission radionuclide combinations were determined experimentally to be = 0.97 for 153Gd/gg"T~ and 1.04 for 57Co/'01Tl. This is not necessarily an adequate correction for bone, thus dual energy methods have been investigated to better calculate the attenuation coefficient of bone [21I. 4 ) Reconstruction Software Emission: The measured emission projection data are reconstructed using the iterative MLEM algorithm to obtain an attenuation-corrected emission image. For each iteration, an attenuated projection and backprojection is performed using attenuation factors that are formed as exponentials of the partial line integrals of the reconstructed attenuation distribution from the pixel of interest to the detector [22]. The factors are calculated only once and stored in memory in order to speed up the ML-EM algorithm. For each iteration n 1, the emission distribution, pa+' for pixel i , is updated using the following [23], [24]:
+
where p , is the ith pixel for the reconstructed emission image, pm is the measured data at the mth projection bin, and f i m is the attenuation factor. The attenuation factors f,, are calculated from the reconstructed attenuation coefficient distribution using
(b)
Fig. 11. Jaszczak Cardiac Torso Phantom (38 x 26-cm elliptical transaxial area). (a) The phantom setup without external breasts. (b) The phantom setup with external breasts.
where the projection ray m intersects the pixel i at the point b,, nearest the detector and the point azm farthest from the detector. The weighting factor w,, is the length of the projection ray m intersecting pixel i . The terms ~ ( a , , ) and a(b,,) are the exponential of the line integrals of the attenuation distribution from the detector to the corresponding
points intersecting the pixel. 111. PHANTOM STUDIES A. Data Acquisition and Processing Phantom studies were performed using the large Jaszczak torso phantom2 with lung, bone, liver, and cardiac inserts, both with and without external breasts. A photograph of the phantom is shown in Fig. 11. Nontruncated transmission slices obtained via the PRISM 2000XP system3 are shown in Fig. 12; these slices are compared against the PRISM 'Data Spectrum Corporation, Hillsborough, NC Picker Intemational, Cleveland, OH
3000XP STEP truncated results. The cardiac phantom studies were performed using both 153Gd/gg"T~ and 57Co/'01Tl simultaneous transmissiodemission acquisition. Another phantom study using the Hoffman brain phantom' was performed using 153Gd/gg"T~ simultaneous transmissiodemiwion acquisition 1) Is3GdPgmTcJaszczak Elliptical Cardiac Phantom Study (Lesion in Septal Wall): For the 153Gd/ggmT~ simultaneous transmissionlemission study, g 9 m Twas ~ injected into the Jaszczak phantom and a 60-mCi 153Gd line source with a 1-mm tin filter was inserted into the transmission line source holder. The bone insert was positioned inside the elliptical container of the Jaszczak phantom. The liver was filled with water (1225 ml) and injected with 117.3 MBq (3185 pCi; 2.6 pCi/cc) of ""Tc. The cardiac insert has two concentric chambers separated by 1.0 cm. A 2 x 1.5-cm defect was placed inside this chamber so that when the cardiac insert was positioned within the elliptical phantom the defect was on
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120 projections were obtained over 360" of rotation. For both energy windows the projection data were digitized into 128 x 128 arrays of 120 projection angles of 2 5 s duration (total acquisition time of 50 min). The activity level was approximately equal to a normal 8-mCi patient injection. The longer-thannormal clinical acquisition time was performed specifically to minimize noise fluctuations so that the ]performance of the attenuation correction could be better evaluated. Both the transmission and emission projections were reconstructed into 128 x 128 matrices with pixel sizes of 3.56 mm using 30 iterations of the transmission algorithm in (4) and the emission algorithm in (7). The reconstructed attenuation distribution was scaled to the correct energy using (6). A reconstruction without attenuation correction was also obtained by using 30 iterations of the emission algorithm with the attenuation distribution set to zero. After reconstruction, the uncorrected and the corrected emission reconstructions were filtered with a 3-D low-pass Butterworth filter of order nine and cutoff frequency of 0.4. The transmission reconstructions were 3-D low-pass filtered with Butterworlh filter of order 7.9 and cutoff frequency of 0.4. This was for display purposes only; the unfiltered transmission reconstruction was used in the calculation of the attenuation factors. On the reformatted short axis slices, the average anteriorto-inferior ratio was calculated for a series of seven short axis slices between the apex and base. Also, the %rms error was calculated for the lateral wall using %rms ierror = 100 x o / p , where 0 is the standard deviation and p is the mean, sampled over a region of 71 voxels in the lateral wall. 2) 57C0/L01 Tl Jaszczak Elliptical Cardiac Phantom Study (Lesion in Septal Wall): For the 57C~/"01Tlsimultaneous transmissiodemission study, 201T1 was injected into the Jaszczak phantom and a 8-mCi 57C0 line source was inserted into the transmission line source holder. The same configuration used in the 153Gd/ggmTcstudly was applied. The liver was injected with 110.4 MBq (3000 pCi; 2.45 pCi/cc) (b) of 'OlTl. The cardiac insert was injected with 10.7 MBq (292 pCi; 2.73 pCi/cc) of zOITl.Body background activity was Fig. 12. Nontruncated transmission results showing transaxial, sagittal, and coronal slices of the Jaszczak Cardiac Torso Phantom in Fig. 11: (a) without simulated by injecting 66.2 MBq (1800 pCi;O.2 pCi/cc) of breasts and (b) with breasts. The results were obtained by using the PRISM 2000XP scanning line source transmission system. The transaxial slices are 201T1in water that filled the spaces other than liver, cardiac, ordered from superior to inferior. The sagittal slices are from the left side lung, and bone. The two upper peaks of "'Tl are not used of the left lung toward the midline through the heart. Coronal slices are in the emission reconstruction. from anterior to posterior. The slice thickness is approximately 13 cm. The Cardiofan collimators (65-cm focal length) were used for transmission source was 153 Gd. the acquisition. Two pulse-height energy windows were set: the 20% window was centered at 122 IteV for the 57C0 the septal side of the heart. The remaining volume of 107 cc transmission data, and the 30% window was centered at 73 between the two chambers was filled with water and injected keV for the ''IT1 emission data. Two sets of 128 x 128 prowith 10.2 MBq (278 pCi; 2.6 pCi/cc) of 997nTc.The liver and jections of 3.56 mm pixels, one for each energy window, were cardiac inserts were positioned inside the elliptical container acquired for each detector. A total of 120 projections were and the remaining spaces (8,820 ml) were filled with water. acquired over 360" by binning the continuous acquisitions into The water was then injected with 92.8 MBq (2,522 pCi 0.29 durations of 25 s per projection (total acquisition time of 50 pCi/cc) of 99mTcto simulate tissue, Water was not injected min) Again, a longer than normal clinicd acquiqition time into the simulated lung regions that contained Styrofoam beads. was performed, specifically to better evaluate the performance Cardiofan collimators of 65-cm focal lengths were used for of the attenuation correction. the acquisition. Two pulse-height energy windows were set: Both the transmission and emission projections were reconthe 20% window was centered at 100 keV for the 153Gdtrans- structed as described above for the 153Gd/"'mT~study, except mission data and the 15% window was centered at 140 keV that the transmission reconstructions were not filtered before for the g 9 m Temission ~ data. Through continuous acquisition, being displayed.
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On the reformatted short axis slices, the average anterior-toinferior ratio was calculated for a series of 12 short axis slices (13 for the attenuation-corrected case) between the apex and base. For the %rms error calculation, the mean and standard deviation was calculated over a sample of 113 voxels for the uncorrected and 126 voxels for the attenuation-corrected reconstruction. 3) 153 G@'"Tc Jaszczak Elliptical Cardiac Phantom Study with Breasts (Lesion in Lateral Wall): For this particular 153Gd/ggmT~ simultaneous transmissiodemission study, synthetic breast appendages were secured to the outside of the Jaszczak elliptical cardiac phantom [Fig. ll(b)]. The rest of the phantom configuration was similar to that used in the previous studies, with the exception of the lesion being located in the lateral wall instead of the septal wall. The ggmTc was injected into the Jaszczak phantom and a 60-mCi 153Gd line source with a 0.25-mm tin filter was inserted into the transmission line source holder. The liver insert was injected with 117.3 MBq (3185 pCi; 2.6 pCi/cc) of 9 9 m T ~The . cardiac insert was injected with 10.2 MBq (278 pCi; 2.6 pCi/cc) of 9 g m T ~Body . background activity was simulated by injecting 92.8 MBq (2,522 pCi; 0.29 pCi/cc) of 9 9 m Tinto ~ the rest of the phantom torso. Cardiofan collimators (65-cm focal length) were used for the acquisition. Two pulse-height energy windows were set: the 20% window was centered at 100 keV for the 153Gdtransmission data and the 15% window was centered at 140 keV for the 9 9 m Temission ~ data. A continuous acquisition over 360" was performed. For both energy windows, the projection data were digitized into 64 x 64 arrays of 120 projection angles of 15-s duration (total acquisition time of 30 min). This study was performed with a normal clinical acquisition time, with a normal activity level (equivalent to an 8 mCi injected dose), and with the large-size breasts so that the performance of the attenuation correction could be better evaluated under simulated clinical conditions. Both transmission and emission projections were reconstructed into 64 x 64 matrices with pixel size of 7.12 mm using both 30 and 99 iterations of the transmission algorithm in (4) and the emission algorithm in (7). The reconstructed attenuation distribution was scaled to the correct energy using (6). A reconstruction without attenuation correction was also obtained using both 30 and 99 iterations of the emission algorithm with the attenuation distribution set to zero. After reconstruction, the uncorrected and the corrected emission reconstructions were filtered with a 3 D low pasa Butterworth filter of order nine and cutoff frequency of 0.4. The transmission reconstructions were not filtered. The average anterior-to-inferior ratio was calculated for a series of short axis slices between the apex and base as described before. Also, the %rms error was calculated as described above, except the mean and standard deviation was calculated over voxels sampled in the septal wall instead of the lateral wall because the lesion was positioned in the lateral wall for this study. 4 ) 153GdPgmTcJaszczak Elliptical Cardiac Phantom Studies with Breasts (No Lesion): Another study was performed with the phantom with breast. The phantom configuration was
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similar to that used in the previous study except no lesion was inserted in the walls of the cardiac insert and the liver emission intensity was a factor of two less than in the previous study. The liver insert was injected with 55.2 MBq (1500 pCi; 1.22 pCi/cc) of 9 9 m T ~The . cardiac insert was injected with 10.1 MBq (276 pCi; 2.6 pCi/cc) of g g m T ~Body . background activity was simulated by injecting 92.0 MBq (2500 pCi; 0.28 pCi/cc) of ""Tc into the rest of the phantom torso. The remainder of the data acquisition and data processing was the same as that described in the previous study. 5) 153GdpgmTcHofSman Brain Phantom Study with Alumina Bone Simulator: For the Hoffman brain phantom study an alumina sleeve was placed around the cylindrical phantom to simulate bone attenuation. The alumina sleeve had a thickness that varied from 114 to 1/16 in to approximate bone attenuation. The phantom was injected with 147.2 MBq (4 mCi) of g g m T ~A. 60 mCi 153Gdline source with a 1-mm tin filter was inserted into the transmission line source holder. The same 65-cm focal length cardiofan collimators used in the previous studies were used for this acquisition. In most cases, 50-cm fan-beam collimators are recommended for brain studies. The longer focal length was used in this case to avoid truncation of the phantom, which was somewhat larger than an actual human head. Two pulse-height energy windows were set as previously described. The acquisition was setup with a circular detector orbit-of-radius equal to 18.5 cm. A step-andshoot acquisition of 60 s per stop collected 120 projections over 360" for each energy window (total acquisition time of 120 min). Each projection was binned into 128 x 128 arrays of 3.56-mm pixels. The acquisition time was greater than a normal clinical acquisition time to minimize noise fluctuations so that the performance of the attenuation correction could be better evaluated. Both the transmission and emission data were reconstructed into 128 x 128 arrays using 99 iterations of both the transmission and emission algorithms. Ninety-nine iterations were used instead of 30 iterations to obtain better resolution in the attenuation-corrected reconstructions. The reconstructed attenuation distribution was scaled by using (6), which is expected to be in error of 7% for alumina. A reconstruction without attenuation was also obtained using 99 iterations. After reconstruction, the uncorrected and the corrected emission reconstructions were filtered with a 3-D low-pass Butterworth filter of order 12 and cutoff frequency of 0.5.
B. Results The first two cardiac phantom studies (without breast) were performed with normal activity levels with an acquisition time 513 times longer than what is usually performed in the clinic and were reconstructed with a pixel resolution (3.56 mm) of twice that normally used in the clinic. The third and fourth studies (with breast) were performed with approximately normal activity levels and a normal acquisition time in order to obtain clinical count levels that were reconstrncted using a normal clinical pixel resolution of 7.12 mm. In the third study a lesion was located in the lateral wall, whereas in the first and second studies a lesion was located in the septal
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Fig 13 Results of 153Gd/g”T~ simultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom (Study 1) The two columns of images on the right are attenuaaon-corrected short axis slices of the cardiac insert The two columns of short axis slices on the left are uncorrected The order of slices starts from the apex of the heart and proceeds to the base reading from left to right and down The lesion can be seen in the septum (on the left in the mid-ventricular to basal short axis slices)
wall. The fourth study was performed without any lesion. The brain phantom study was performed with an acquisition time approximately three times longer than normal clinical acquisition times. I ) ls3GdPgmTcJaszczak Elliptical Cardiac Phantom Study (Lesion in Septal Wall): Fig. 13 shows the results of the 153Gd/ggmT~ simultaneous transmissiodemission study using the Jaszczak cardiac phantom. Short axis slices, with and without attenuation correction, are shown along with horizontal and vertical profile plots through the center of six representative slices. Each image and its corresponding profile are scaled to the maximum count in the entire series of images. The two columns of short axis slices on the left are reconstructed without attenuation correction. In the images and profiles, inferior regions are observed with markedly reduced uptake of radionuclide concentration. Profile plots show an overall improved uniformity of activity from one wall to the other in the attenuation-corrected reconstructions. However, the vertical profiles do not show complete uniformity. Some of this may have been caused by partial volume effects, scatter, or lack of algorithm convergence. Because the group of profiles on the left, as well as the ones on the right, are scaled individually to the group maximum, they do not show the overall increase in intensity in the attenuation-corrected
reconstructions. It is also difficult to appreciate from the plots the significant increase in contrast of the imyocardium to the background in the attenuation-corrected reconstructions, which is more evident in the images themselves. Note that in the attenuation-corrected results there appears to be a hot region projecting upward from the left side of the liver. This is an artifact that appears to be caused by the truncation of the liver. It is probably more the result of the transmission map being truncated at the edge of the image array than the truncation of the emission data. One can see in the reconstructed transmission transaxial images in Fig. 15 that the phantom is shifted off-center. The right side of the phantom (left side of the images in Fig. 15) is truncated by the image array with some high intensity pixels along the edge. These larger attenuation coefficients can cause an overcorrection in emission intensities. Improved uniformity throughout the myocardium in the attenuation-corrected reconstructions is better appreciated in the bull’s-eye plots shown in Fig. 14 (note that each bull’s-eye is scaled to its own range). The defect is shown clearly in both the attenuation-corrected and uncorrected reconstructions. The apical slices can also be better appreciated using bull’s-eye plots. The inferior portion of the apex has somewhat higher intensity after attenuation correction, whereas the superior por-
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F1g 14 Bull's-eye plots of the 153Gd/ggmT~simultaneous transmission and emission imaging of the Jaszczak Cardmc Torso Phantom (Study 1) On the left are the uncorrected and on the nght are the attenuabon-corrected bull's-eye plots The lesion is seen at 180° (Note that the gray levels are scaled so that 50% of the grayscale is black)
Fig. 15. Reconstructed transmission maps in the Gd/99mTc simultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom. The transaxial slices are ordered from superior to inferior. The phantom was shifted of€-center so that cardiac insert is at the center of rotation. The body contour is well defined on the right of the phantom but is blurred due to truncation on the left.
tion of the apex had higher intensity before correction. From the grayscale, this represents less than 10% overcorrection. (Note that the gray levels are scaled so that 50% of the grayscale is black.) Some of this increased intensity may be due to overcorrection as a result of scatter from the liver. In particular imaging situations, it has been shown that scatter from the liver can result in an overcorrection in the inferior myocardium of attenuation-corrected images [25]. To quantify the attenuation correction, the average anteriorto-inferior ratio was calculated for seven short axis slices, yielding 1.26 before correction and 0.97 after correction (Table I). Also, the %rms error was calculated for the lateral wall of the left ventricle. The %rms error was 11.3% in a 71voxel region sampled over seven 3.65-mm slices of the lateral wall before attenuation correction and 6.3% after attenuation correction. Note that these calculations were performed on the reformatted short axis slices. The measure of mean and %rms
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tin filter was utilized in the transmission Source assembly, The study ln a maximum transmission flux Of 352 counts/pixel/proJectionview. The gross count rate for detector 3 was approximately 20K cps, which produced approximately 5% count loss including - scatter. No correction was performed for this count loss. The outer boundary of the support region, which is the nonzero, nonblack region, can be seen. This region is calculated from the orbit and bed setup parameters. The sharp, straight vertical boundary on the left side is the edge of the image array. The phantom was offset, which is commonly done in the clinic with patients so that the heart is positioned in the center of the field of view. Because the phantom is positioned off-center, a support region that goes beyond the image array is calculated from the orbit setup parameters. The reconstruction algorithm truncates the calculated support to the image array. This does not necessarily mean that the contour of the phantom goes beyond the image array since the support is calculated to be larger than the body contour. The lung and outer boundary of the phantom are clearly delineated on the right in each transaxial slice. It is observed on the left of each transaxial image that the truncation has blurred the lung image and the outer body contour. The left edge of the square image array also shows an apparent increase in the attenuation coefficient. Simulations [26], [27] have shown that these apparently increased areas of attenuation do not adversely affect the attenuation-correctedimages in the heart region of interest, but may affect regions outside of the heart that are near the truncation edge as was noted above. This is a case where, even though the image is distorted, the line integrals contributing to the attenuation factors within the heart region have small errors (less than 10%) [27]. Notice that the simulated vertebral column in the bone insert is well resolved in the image. 2} 57Co/L01 TI Jaszczak Elliptical Cardiac Phantom Study (Lesion in Septal Wall): Fig. 16 shows the results of the 57Co/201Tlsimultaneous transmissiodemission study. Results in this study were very similar in terms of overall uniformity of activity to those results obtained in the 153Gd/99mT~ study (Fig. 13). The vertical profile in the lower right, however, is somewhat less uniform than in the previous 153Gd/99"T~ study. This could be due to partial volume effect, since these slices were located near the base of the cardiac insert, or it could be the result of lack of convergence. The attenuationcorrected bull's-eye plot in Fig. 17 shows a fairly uniform plot with some noise mottle. The lesion is better resolved in the attenuation-corrected bull's-eye. Notice also that, even with a worst scatter situation, the attenuation-correctedreconstructions do not seem to be overcorrected over a large region of
GULLBERG et al.: SIMULTANEOUS TRANSMISSION AND EMISSION TOMOGRAPHY TABLE I AVERAGE ANTERIOR~~NFERIOR RATIOSAND %RMS ERRORS IN THE
Average Anterior/Inferior Ratio
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JASZCZAK CARDIAC PHANTOM
EXPERIMENTS
%RMS Error (Lateral or Septal Wall)
No Breast, Lesion in the Septal Wall 153Gd/g%'c 1.26 k 0.08 (uncorrected, 30 iterations)
11.4%(147.2+16.7, 7slices, 7lvoxels)
153Gd/9*c 0.97 f 0.09 (corrected, 30 iterations)
6.3%(1680+106, 7slices, 7lvoxels)
No Breast, Lesion in the Septal Wall 57~~/201~1 1.26 f 0.09 (uncorrected, 30 iterations)
15.0%(41.9+6.3, 12slices, 113voxels)
1.06 f 0.11
9.8%(1065+104, 13slices, 126voxels)
57~~/201~1
(corrected, 30 iterations) Breast, Lesion in the Lateral Wall 153Gd/9%'c 1.24 f 0.08 (uncorrected, 30 iterations)
1 2 . 8 % * ( 6 8 8 + 8 8 , 9slices, 70voxe~ls)
153Gd/9-c 1.16 f 0.08 (corrected, 30 iterations)
12.5%*(6041+756, 9slices, 68volxels)
153Gd/Q*~ 1.22 f 0.05 (uncorrected, 99 iterations)
10.4%*(177+18.4, lOslices, 96voxels)
153Gd/g*c 1.07 k 0.05 (corrected, 99 iterations)
9.4%*(1577+148.5, gslices, 87voxels)
Breast, No Lesion 153Gd/g-c 1.30 f 0.09 (uncorrected, 30 iterations)
17.3%(1089+189, Plslices, 190voxels)
153Gd/9qc 1.03 k 0.09 (corrected, 30 iterations)
11.1%(9598+1061, Zlslices, 190voxels)
153Gd/9qc 1.29 k 0.11 (uncorrected, 99 iterations)
14.6%(1006+147, llslices, 86vo1xels)
153Gd/9wT~ 1.06 k 0.10 (corrected, 99 iterations)
11.1%(9621+1067, llslices, 93voxels)
*This was measured along the septal wall because a lesion was located in the lateral wall. For all other cases the measurement was made along the lateral wall.
the inferior wall. For 201T1,a greater potential for scatter from the liver, contaminating the inferior wall of the left ventricle, ~ the same concentration. is expected than with g g m Tfor The attenuation-correction results are quantified in Table I. The average anterior-to-inferior ratio was calculated and yielded 1.26 (sampled over 12 short axis slices) before correction and 1.06 (sampled over 13 slices) after attenuation correction. As in the 153Gd/ggmT~study, the attenuation-corrected reconstructions showed an improvement in uniformity after attenuation correction; however, errors for the average anterior-to-inferior ratio after correction are a little higher-approximately 10% compared to 9% in the 153Gd/ggmT~ study. The anterior-to-inferior ratios are further from one at the apex of the heart than they are at the base of the heart. It is possible that this is due to partial volume effects
and may also be due in part to scatter. Tlhe %rms error was 15.0% in a 113 voxel region sampled over 12 3.65-mm slices of the lateral wall before attenuation correction and 9.8% (13 slices and 126 voxels) after attenuation correction. Fig. 18 shows transaxial slices of the transmission reconstruction. The transmission source assembly was fitted with a 0.25-mm tin filter. The maximum transmission flux was 68 counts/pixel/projection view. The images in Fig. 18 and those in Fig. 15 appear different because the study displayed in Fig. 15 was performed with the phantom placed off-center, whereas the 153Gd/g9"T~study in Fig. 18 was performed with the phantom centered in the field of view. In Fig. 18, images of both the left and right contours of the phantom are blurred due to truncation in each transaxial image. The support region in Fig. 18 is not truncated by the image arra,y as it is in Fig. 15.
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Fig. 16. Results of 57Co/201Tl simultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom (Study 2). The two columns of images on the right are attenuation-corrected short axis slices of the cardiac insert. The two columns of short axis slices on the left are uncorrected. The order of slices starts from the apex of the heart and proceeds to the base reading from left to right and down. The lesion can be seen in the septal region.
Bull's-eye plots of the 57C~/201Tl simultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom (Study 2). On the left are the uncorrected and on the right are the attenuation-corrected bull's-eye plots. The lesion is seen as in Fig. 13 at 180".
Fig 17
Also, the texture of the images appears to have more noise than those in Fig. 15 because in Fig. 18 filtering was not applied before displaying the images. 3) 153GdPgmTcJaszczak Elliptical Cardiac Phantom Studies with Breasts (Lesion in Lateral Wall): Fig. 19 shows results of the 153Gd/99"T~simultaneous transmissiodemission study of the Jaszczak cardiac phantom with attached breasts. Profile plots show an improvement in the overall uniformity of activity in the myocardium, from superior to infenor, in the attenuation-corrected reconstructions. Reconstructions
were performed with both 30 iterations (not shown) and 99 iterations (shown in Fig. 19) and it was found that the profiles became more uniform with the increased number of iterations. It may be that in this circumstance more iterations are required because of the close proximity of the liver. It is recognized that the EM-ML algorithm requires more iterations to resolve the intensity for structures positioned close to each other, such as the myocardium and liver in this case. Overall, the attenuationcorrected bull's-eye plot in Fig. 20 shows more uniformity than the plot without attenuation correction. Notice the large
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Fig. 18. Reconstructed transmission maps in the 5iCo/20’Tl simultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom. The transaxial slices are ordered from superior to inferior. The phantom was not shifted as in Fig. 15 so the body contour of the phantom was truncated on both the right and left.
Fig. 19. Results of 153Gd/”9”T~simultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom with breasts (Study 3). The two columns of images on the right are attenuation-corrected short axis slices of the cardiac insert. The two columns of short axis slices on the left are uncorrected. The order of slices starts from the apex of the heart and proceeds to the base reading from left to right and down.
difference in the apparent defect size between that seen in Fig. 20 on the lateral side and that seen in Figs. 14 and 17 on the septal side, even though physically, the defect was the
same size for all three studies. This dissimilarity is due to the differences in reconstructing with 128 x 128 and 64 x 64 matrix resolutions.
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Fig. 20. Bull's-eye plots of the 153 Gd/""Tc simultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom with breasts (Study 3 ) . On the left are the uncorrected and on the right are the attenuation-corrected bull's-eye plots. The lesion is seen in the lateral wall at 0".
The anterior-to-inferior ratios, and the %rms errors, are summarized in Table I. For 30 iterations, the average anteriorto-inferior ratio was calculated for nine short axis slices, yielding 1.24 without correction and 1.16 with attenuation correction. For 99 iterations, the average anterior-to-inferior ratio was 1.22 without correction and 1.07 with attenuation correction. The improvement in the attenuation-corrected anterior-to-inferior ratio with 99 iterations is indicative of improvement in convergence with a greater number of iterations. For 30 iterations, the %rms error was 12.8% in a 70-voxel region sampled over nine 7.3-mm slices of the septal wall without attenuation correction and 12.5% (nine slices and 68 voxels) with attenuation correction. (Note the measurements were made in the septal wall instead of the lateral wall because of a lesion in the lateral wall.) For 99 iterations, the %rms error was 10.4% in a 96 voxel region sampled over 10 7.3mm slices in the septal wall without attenuation correction and 9.4% (nine slices and 87 voxels) with attenuation correction. Fig. 21 shows transaxial slices of the transmission reconstruction. The transmission source assembly utilized a 0.25-mm tin filter. The maximum transmission flux was 1090 counts/pixel/projection view. The phantom was positioned at the center of rotation. In the image, the left and right edge of the phantom boundary and the breasts are blurred due to truncation. No filtering was applied before the transmission images were displayed. 4 ) 153 GdPgmTc Jaszczak Elliptical Cardiac Phantom Studies with Breasts (No Lesion): Fig. 22 shows the bull's-eye plots for the 153Gd/ggmTcsimultaneous transmission/emission study of the Jaszczak cardiac phantom with attached breasts and without a lesion. The attenuation-corrected bull's-eye plot shows more uniformity than the plot without attenuation correction. The anterior-to-inferior ratios, and the %rms errors, are summarized in Table I. For 30 iterations, the average anterior-toinferior ratio was calculated for 21 3.65" short axis slices, yielding 1.30 without correction and 1.03 with attenuation correction. For 99 iterations, the average anterior-to-inferior ratio was 1.29 without correction and 1.06 with attenuation correction. For 30 iterations, the %rms error was 17.3% in a 190-voxel region sampled over 21 3.65-mm slices of the lateral wall without attenuation correction and 11.1% with
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attenuation correction. For 99 iterations, the %rms error was 14.6% in a 86-voxel region sampled over 11 7.3" slices of the lateral wall without attenuation correction and 11.1% (11 slices and 93 voxels) with attenuation correction. For this case, the increase in iterations did not improve the ratios as it did in the previous study. Remember, the liver activity was a factor of two less in this study than in the previous study. Therefore, it is possible that the EM-ML algorithm did not need as many iterations to resolve differences between the cardiac insert and the liver. 5) 153GdPgmTcH o f i a n Brain Phantom Study with Alumina Bone Simulator: The results of the Hoffman brain phantom are shown in Fig. 23. Notice the high resolution in the transmission reconstructions. The alumina sleeve is clearly visible across the ll4-in-thick region but there is loss of resolution as the sleeve narrows from 1/4 in at the bottom to 1/16 in at the top of the transmission image. The loss of resolution causes a partial volume effect that decreases the intensity in the thinner regions of the alumina attenuator. Notice that it is difficult to differentiate between the water and the lucite in the phantom. For this particular study the transmission flux had a maximum count of 1933 counts/pixel/projection view. In the images and profiles it can be seen that the attenuationcorrected reconstructions show an improvement in quantitation for interior pixels in the phantom as demonstrated by the increased brightness of the pixel values and the more uniform peak intensities in the profiles across the phantom. IV. DISCUSSION The results of this study show that a three-detector SPECT system with a lightweight low-dose high-resolution collimated line source mounted to the rotating gantry, with a controller that moves the line source in synchrony with one of three fan-beam collimated detectors, can acquire simultaneous transmission and emission data that produce attenuation-corrected source distributions. The design choices dictated by incorporating a transmission-computed tomography system into a threedetector SPECT system pose problems of transmission data truncation and crosstalk between transmission and emission data windows. These technological problems can be overcome by carefully designing data preprocessing algorithms and iterative reconstruction algorithms capable of providing images for diagnostic viewing in less than 5 min. An important conclusion derived from this work is the evidence that the use of iterative reconstruction algorithms, which solve the system of linear equations corresponding only to the truncated transmission data, can be used to reconstruct attenuation-corrected emission distributions of cardiac radiopharmaceuticals even though the transmission image may be significantly distorted. The attenuation factors, which are exponentials of partial line integrals of the attenuation distribution, are calculated with less than 10% error from the truncated attenuation reconstruction for those attenuation factors that have the greatest influence upon the heart emission measurements [26]. This is because the partial line integrals of attenuation coefficients are formed from ray sums that satisfy, in a least squares sense, the measured projection data.
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Fig. 21. Reconstructed transmission maps in the 153Gd/99mTcsimultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom with breasts. The transaxial slices are ordered from superior to inferior. The phantom was not shifted off-center.
Fig. 22. Bull's-eye plots of the 153Gd/99mT~simultaneous transmission and emission imaging of the Jaszczak Cardiac Torso Phantom with breasts and without lesion (Study 4). On the left are the uncorrected and on the right are the attenuation-corrected bull's-eye plots.
The lightweight low-dose high-resolution collimated line source is mounted to the rotating gantry with a controller that keeps the line source in the focal line of the opposing fan-beam collimator as it moves along the body contour. The useful line source strength can range from 15 to 60 mCi for 153Gd and from 5 to 20 mCi for 57C0.The added transmission dose to the patient is small (
