Emission computed tomography

Emission Computed Tomography M i c h a e l E. Phelps Although there are many c o m m o n aspects to x-ray transmission and radionuclide emission computerized tomography, (ECT) there are added difficulties and a number of particular factors that form the basis of ECT. A t this time, the instrumentation design and application strategies in ECT are diverse and in a stage of rapid development. The approaches are divided into two major categories of single-photon counting (SPC), employing scanner and camera concepts with radionuclides of 9oreTc, 2~ ~231,etc. and annihilation coincidence detection ( A C D ) of positron-emitting radionuclides. Relatively uniform detection response with depth and approximate methods for photon attenuation correction have been demonstrated for S PC studies of the brain. However, A C D systems have a general advantage of more accurate methods for

photon attenuation correction, higher efficiency, and more uniform sensitivity and resolution with depth when applied to whole body studies. The deciding factor between these t w o approaches and the success of ECT in general depends to a large extent on the d e v e l o p m e n t of l a b e l e d c o m p o u n d s and procedures directed at the unique property of ECT in providing physiologic information not available from other techniques. ECT not only provides tomographic i m a g i n g a d v a n t a g e s , it a l l o w s a q u a n t i t a t i v e measurement not possible in the past. SPC approaches have an advantage in the availability of radionuclides and a disadvantage in providing true physiologic tracers. A C D approaches have the reverse situation at this time.

OMOGRAPHIC techniques for minimizing or removing the superposition of information with depth have been the goal of a number of systems for producing an image of a section of the body at a given depth. In the past, most tomographic images in diagnostic radiology and nuclear medicine '-3 were produced by employing the principles of "focal plane" tomography, first descr'bed by Bocage in 1921. More recently, computerized mathematical techniques have been developed that allow exact or nearly exact reconstruction of tomographic sections of the body. This technique is referred to by the popular name of computerized tomography (CT). Although the first examples of computerized transaxial reconstruction tomography were demonstrated in nuclear medicine by Kuhl and co-workers, 4 8 the successful development of the EMI (EMI Medical Limited, Middlesex, England)x-ray transmission scanner by Hounsfield has produced an incredible advancement in this technique in the past 5 years. While the tech-

nique of CT is clearly applicable to nuclear medicine imaging, its clinical application has lagged far behind x-ray transmission (TCT). Emission CT (ECT) is somewhat more difficult than TCT as discussed elsewhere) Many of the technical issues in ECT have been or are now being solved. Thus, it would appear at this time that ECT is in need of three things: (1) expanded technical effort, (2) clear definition of purpose, and (3) clinical results that provide information not obtainable by other techniques and that clearly demonstrate that its importance exceeds its cost. Development of ECT has followed two approaches: single-photon counting (SPC), which employs either multiple detector arrays with converging collimators or scintillation cameras for the detection of 99mTc, 2~ 'z:~I, ~:~I, etc., and annihilation coincidence detection (ACD) of positron-emitting radionuclides.

T I

From the Division of Nuclear Medicine, Center for the Health Sciences, Laboratory of Nuclear Medicine and Radiation Biology, UCLA, Los Angeles, Calif. Michael E. Phelps, Ph.D., Professor of Radiological Sciences, Department of Radiological Sciences, Center for the Health Sciences, UCLA, Los Angeles, Calif. Supported in part by ERDA Contract EY-76-C-03-0012. Reprint requests should be addressed to Michael E. Phelps, Ph.D., Department of Radiological Sciences, Center for the Health Sciences, UCLA, Los Angeles, CaliJ. 90024. r by Grune & Stratton, Inc. 1SSN:0001-2998. Seminars in Nuclear Medicine, Vol. 7, No. 4 IOctober), 1977

RECONSTRUCTION ALGORITHMS

The basis of reconstruction tomography can be illustrated with the example of transaxial CT (Fig. 1). Data are obtained from linear scans at discrete angles (or any equivalent scheme) around the cross-section of interest. These data are then processed in one of a number of mathematical algorithms to solve for the unknown cross-sectional structure. The mathematical approach is to reduce the threedimensional problem to one of two dimensions (2-D) by first assuming that the cross-section has no thickness (e.g., a 2-D plane). This 2-D 337

338

MICHAEL E. PHELPS Detector Field of -I-i-Linear Scan

View

Cross Sectional

~ .

=,

: '~Y

Slice

A

B

lllilllllI lllW'~qlli

II.......... II Ik ~ ~ ~ ~ II I I L :; ,II lllI._ -dill liililllll

D Fig. 1. Schematic ilustration of data collection format for transaxial ECT consisting of linear scans at discrete angles through 1 8 0 ~ around the object. A], A2 .... A are the discrete amounts of activity within the detector field of view at each linear scan point.

problem is then partitioned into a linear series of 1-D linear problems. For example, each datum point or detector reading, y, in a linear scan is assumed to represent a linear equation of the form: y = A1 +,42 . . . . . + A

(l)

where A 1 through A are the unknown amounts of activity along the line-of-view of the detector (Fig. 1). A linear set of these equations is generated during each linear scan. If enough data samples are taken in the linear and angular direction, a cross-sectional image can be reconstructed. This formulation of the problem is the same for both x-ray and emission CT. However, there are a number of assumptions in the reconstruction algorithm that pose additional problems in ECT. For example, each detector reading (y in equation 1) is assumed to be equal to the linear sum of activity in each picture element across a well-defined line at the position where the reading was taken (Fig. 1). This formulation neglects the fact that the photons emitted from the activity in the tissues are decreased to varying degrees by attenuation. Thus, photon attenuation r e p r e s e n t s another set of unknowns that must be solved or corrected for before the image can be reconstructed with conventional tomographic algorithms. Methods for compensating or corr e c t i n g for p h o t o n a t t e n u a t i o n and o t h e r physical factors will be discussed below.

Reconstruction tomography is not limited to the transaxial orientation just discussed. Rec o n s t r u c t i o n can be a c c o m p l i s h e d in a transaxial or a longitudinal direction or at any angle between these two extremes if data can be collected in a f o r m a t consistent with the mathematical assumptions of the algorithm. In fact, methods will probably be developed for reconstruction of the full 3-D object soon. In the broader sense, reconstruction tomography is a technique in which projections (either 1- or 2-D) are used to reconstruct mathematically the exact or nearly exact 3-D object. Present reconstruction algorithms can be separated into two categories: algebraic techniques, which generally use iterative schemes, and Fourier techniques, which use Fourier series, Fourier transforms, or convolution techniques. These techniques have been reviewed in several publications.'~ The most common reconstruction algorithm used today in CT is the convolution or linear superposition of filtered back-projections. This choice is primarily based on the computational and practical aspects of time, cost, and flexibility of implementing the reconstruction algorithms in a digital format. However, the choice of the optimum algorithm is in large part dependent on the particular application. Each type of algorithm has advantages and disadvantages applicable to each type of problem, and it should not be assumed that Fourier-based algorithms are optimal for all situations. PHYSICAL FACTORS AND ASSUMPTIONS

Some of the important physical factors and assumptions that provide optimal reconstructed image quality in ECT are: (1) detector reading proportional to linear sum of activity--correction or compensation for photon attenuation, no scattered radiation is detected, detector and associated electronics have linear response over dynamic range of count rates encountered; (2) uniform detection response with depth; (3) high detection efficiency to assure acceptable statistical accuracy; (4) linear and angular sampling consistent with detector resolution, required accuracy in reconstructed image, and desired image resolution; (5) activity, patient and organ remain in a stationary state during time of m e a s u r e m e n t ; and (6) a c c u r a t e mechanical positioning in linear and angular direction.

EMISSION COMPUTED TOMOGRAPHY

The detection of scattered radiation should be minimized since it produces nonlinear effects in equation 1. This is particularly true if the amount of scattered radiation detected is depthdependent (which is typically the case). The detector and associated electronics should exhibit minimum instability and dead-time losses over the dynamic range of count rate encountered or equation 1 will be violated. If multiple detectors, or position-sensitive detectors (e.g., a scintillation camera), are employed, corrections should be applied to compensate for variations in intrinsic efficiency and nonuniformity. The detector response should be uniform or nearly uniform with depth (Fig. 1), since this is assumed in most reconstruction algorithms. If the resolution does vary with depth, then methods must be developed to accomodate this, or one must accept the spatial averaging (e.g., resolution loss) and image artifacts that result. ACD has been shown to have nearly constant resolution with depth 9.~3,~4 in accordance with this assumption. Kuhl et al., s using converging collimators on the M A R K IV Scanner, have shown reconstructed line spread functions that v a r y a b o u t 10% full-width h a l f - m a x i m u m ( F W H M ) across an 18-cm diameter object. Recently, Jasczak et al.t5 and Budinger et al. 1~ have shown a nearly constant resolution with depth with the scintillation camera by using the geometric mean (i.e., square root of the product of directly opposing projections). These results were obtained with headsized phantoms in close proximity ( - 2 cm) to the face of the camera, a condition that is difficult to meet in a patient study. However, these results do indicate that with proper collimator design a relatively uniform response with depth for objects the size of the head can be achieved. More studies are needed to d e m o n s t r a t e whether these app r o a c h e s with S P C will p r o v i d e u n i f o r m response with depth for objects the size of the torso. Another fundamental assumption in the reconstruction algorithm is that the object remain in a stationary state during the course of measurement. If this stationary state is not maintained, motion artifacts will result. In ECT this means that the patient, organ, and injected radiopharmaceuticai must not change location significantly during the course of measurement. The redundant sampling (e.g., each datum point

339

in the scan profiles is repeatedly sampled at many different times throughout the scan) of the P E T T III, 17-z~ O R T E C / E C A T (Life Sciences Division, O R T E C Inc., Oak Ridge, Tenn.), and PETT IW ~ significantly reduces artifacts due to motion. This type of redundant sampling is also possible in the dual-headed scintillation camera developed by Muehllehner et al. ~2,~3 Stationary circular positron tomographs also provide protection against motion artifacts. Redundant sampling also minimizes artifacts from detector instability. Accurate mechanical positioning in both the linear and angular scans is extremely critical in reconstruction tomography. Huang 24 has shown that with an image resolution of 1 cm, a fixed displacement error of 0.5 mm during a scan will produce an artifact with a magnitude of as much as 10% of the local activity level. For example, if the image contains a hot spot that is 20 times greater than the general activity level, then an error twice (0.10 x 20) as large as the general activity level will result close to the hot spot (this error will decrease by l/r away from the hot spot).

Sampling The linear sampling required in CT is established by the sampling theorum 25 which states that data should be sampled in intervals, Ax, one-half the size of the resolution to be recovered in the final image. This is for ideal conditions, and in reality, sampling is frequently less than half the size of the desired resolution, R: Linear sampling distance, 2we < ~ R

(2)

For example, if l-cm image resolution is desired, then Ax should be _< 0.5 cm. The linear sampling should also be consistent with the spatial resolution of the detector. It can be shown that if the detector response is Gaussian in shape, then for an optimum signal-to-noise ratio, the final F W H M image resolution should be less than or equal to the F W H M inherent spatial resolution, r, of the detector. 26 With this result and equation 2: Linear sampling distance, Ax _< ~2r

(3)

The number of angles over which linear scans should be taken is given by: 11'27 Number of angles ~ 7rD/2Ax

(4)

340

MICHAEL E. PHELPS

B u t t ~ r w o ~ h ~Fmax P 3 2 ; n~e=~12;}

2*

3o

5~

6=

The examples given above illustrate the req u i r e m e n t s to achieve n e a r l y p e r f e c t reconstruction. If these requirements are not met, artifacts are produced in the final image. However, the significance of these artifacts depends on the magnitude of the total error or noise in the image (i.e., from statistical noise, motion errors, inaccuracies of photon attenuation correction, etc.). Therefore, in reality, the choice of sampling is also dependent on the overall accuracy requirements of the procedure, and practical sampling requirements may be less than those stated above. A note of caution should be given. Due to the particular way error propagation occurs in CT, the sampling capabilities in instrument design should be carefully examined before intuitive decisions are implemented.

9~

Attenuation Correction

10"

i!iiiiiil i!!i! 84

Photon attenuation produces a significant loss of information in all emission imaging, and ECT is no exception. Optimal and quantitative reconstruction requires an accurate method for

Fig. 2. Reconstructed i m a g e s from computer simulations showing effect of varying the size of the discrete angular sampling. A n g l e of 2 ~ is consistent w i t h eq. 3 and no artifacts are seen. As angle is increased noise begins to appear in i m a g e because object is u n d e r s a m p l e d in t h e angular direction. Left: grey scale image. Right: a v o l u m e m e t r i c display mode. Noiseless data was used for reconstruction (Courtesy of T. F. Budinger).

where D is the diameter of the object, and Ax is the linear sampling distance. Intuitively, equation 4 states that the angular sampling distance should be approximately equal to the linear sampling distance. This can be appreciated from equation 4 in that the number of angular projections is simply obtained by dividing the length of the 180 ~ arc around the object (TrD/2) into equally spaced increments of length Ax, the linear sampling distance. For example, if D is 32 cm and a linear sampling distance, Ax, of 5 mm was chosen, then equation 4 wouldpredict that the number of angular projections would be 100 or that samples should be taken every 1.8 ~ over the 180 ~ rotation. The effect of using fewer angular projections (i.e., larger angular increments) is illustrated in Fig. 2. Distortions also occur if angular projections are not recorded over a full angle of 180 ~, as illustrated in Fig. 3.

Fig. 3. C o m p u t e r s i m u l a t i o n s h o w i n g effect of t a k i n g data over a full angle o f 9 0 ~ , 135 ~ , and 1 8 0 ~ (A) A c t u a l phantom showing u n i f o r m d i s t r i b u t i o n w i t h four hot spots. (B} Reconstructed i m a g e w i t h data collected over 9 0 ~ ; note distortions, (C) Reconstructed i m a g e from data taken over 135 ~ ; I m a g e is i m p r o v e d over B but some distortions still remain. (D) Reconstructed i m a g e illustrating t h a t w h e n data are collected over a full 1 8 0 ~ , no distortions are apparent. 1

EMISSION COMPUTED TOMOGRAPHY

341

Table 1. Experimentally Determined Variation of Average Path Length Linear Attenuation Coefficients,/4~, for A C D in Human Subjects l~ Organ

Position

Head

4 cm above orbital meatil line 4 cm above tip of

Abdomen

Average Attenuation Coefficient. #~* 0 0 8 8 • 0.003 (• 0.089 • O,007 ( • 8%)

xiphoid Thorax

Fourth intercostal space

0,067 • 018

(•

9 Number in brackets is • 1 SD.

attenuation correction. The methods of correcting for photon attenuation are one of the major points of difference between SPC and ACD tomographic systems. The unique "electronic collimation" of ACD allows nearly exact corrections for photon attenuation.14.~9 Compensation for photon attenuation in ACD can be accomplished either by geometric computation or measured attenuation corrections in a simple, convenient, and accurate m a n n e r . 14,17 19 Attenuation in ACD is also a slowly varying function both because of the high energy of 511 keV and the fact that the coincidence requirement produces attenuation that is dependent on the total path length across the object. This is in contrast to present SPC methods in which attenuation must be corrected for each picture element. Measured variations in the magnitude of attenuation in ACD at different cross-sectional levels in the human 1'~ are shown in Table l. These small variations in attenuation can be compared to the point-to-point variations in

SPC attenuation, which can range from 180% in the head (bone-to-tissue) to the extreme of 580% in the thorax (bone-to-lung tissue). This refers to the variation in attenuation and should not be confused with the total amount of attenuations that occurs with SPC and ACD that will be discussed below. Budinger et al. 28 have recently reviewed eight different methods for photon attenuation for SPC. Their conclusion is that most of the methods are best when the photon energy is high (i.e., > 500 keV) and when the object is small and uniform. Kuhl et al. 8 have also shown good results for attenuation correction brain studies by using an empirical method (assuming head is a uniform cylinder of water). Waiters et al. 29 have employed an iterative correction for SPC attenuation for the brain that shows improvement in delineation of internal structures but also appears to increase image noise. Jasczak et a1.12 have recently shown that using the geometric mean of opposing views with the scintillation camera produces a relatively uniform resolution with depth and only small variation due to attenuation with depth. However, these results are strictly valid only for line sources measured at different depths one at a time and when all the path lengths through the object are the same, which is never the case with the human subject. Therefore, more work is needed to establish better photon attenuation correction methods for SPC. The effect of photon attenuation in ECT is shown in Fig. 4. It should be noted that any

Fig. 4. Reconstructed images at 4 cm below the top of the ziphoid using ~:~NH:~to image human liver w i t h PETT III. Crosssectional emission images are s h o w n to illustrate the distortion that occurs w h e n corrections are not applied for photon attenuation (A). The correct images are s h o w n in (B) w h e r e attenuation has been corrected for by using external transmission source and in (C) w h e r e a geometric correction has been applied. The structure in the images are the liver at the right, the pancreas, left kidney, and spleen. 1~

342

MICHAEL E. PHELPS

Table 2. Statistical Error Versus Number of Counts in ECT Image Number of Counts Required at Different

Image Resolutiont Statistical Error* (•

SD in %)

0.5 cm

1 cm

2 cm

3 cm 7.7 X 10 ~

•

1.7 • 108

2.1 •

10 r

2.6 • 106

•

4.5 • 107

5.6 • 10#

7.0 • 10 ~

2.1 • 10 z

•

1.9 X 107

2.4•

3 . 0 X 105

8.9 • 104

108

*Upper bound of statistical error. 1"For a 20-cm diameter object and with linear sampling distance, A x = V2 image resolution. If object diameter is increased to 3 0 crn, then number of counts required for each error value must be increased by 3.4. Number of counts required are from a upper bound error limit and using a ramp filter function.

inaccuracies in the attenuation correction m e t h o d are p r o p a g a t e d into the final rec o n s t r u c t e d i m a g e . T h e s e e r r o r s can be extremely large and lessen the quality of already photon limited ECT images.

Accuracy T h e a c c u r a c y of the final r e c o n s t r u c t e d image is a function of almost every p a r a m e t e r involved in the instrument design, mathematical approach, tracer technique, attenuation correction, n u m b e r of d e t e c t e d photons, patient, organ and activity movement, etc. The magnitude of each of these is difficult, if not impossible to define in a general way. However, there are a number of relationships that relate to accuracy in reconstruction tomography that can be defined and used as guidelines. If the sampling requirement discussed above are met, errors due to the mathematical calculation have been shown by Shepp et al. 3~ to be insignificant compared to other errors in CT. It has been shown that for transaxial CT the magnitude of statistical error is inversely proportional to image resolution, R, to the third power. 11,31,32 This can be compared to conventional 2-D imaging in which the statistical accuracy is inversely proportional to R 2. Thus in CT we have: R-3 c~number of detected photons

(5)

if the same accuracy per resolution element is maintained when the resolution is changed. For example, to improve the resolution by a factor of 2 in the transaxial plane, one must increase the number of detected photons by a factor of 8. This relationship imposes stringent limitations in resolution improvements in both emission and transmission CT, since the time and/or patient

dose must increase by a factor of 8.* The number of counts required to achieve different image resolutions with 5%, 10%, or 15% statistical error for 20-cm diameter uniform object are given in Table 2. If resolution is also increased in the axial direction (i.e., making the slice thinner) time and/or dose* is inversely proportional to R, and thus, resolution improvement in both the plane and the axial direction would be proportional to R ~. If R is improved by 2 in both the plane and axial direction, time and/or dose* would be increased by 16. A more detailed description of these relationships is given by Phelps et al. 9 It is frequently stated that nuclear medicine scintigraphy is a field of photon-limited images. Where this is true, it applies to almost all imaging fields. In x-ray CT, more than a billion photons are detected, compared to ECT with about a million. However, the important objectcontrast in x-ray CT is only in the r a n g e o f 0 . 5 % -4%, as compared to ECT in which objectcontrast is 20%-2000%. Thus, the lower photon density in ECT is offset by its higher objectcontrast (i.e., 105 counts are required for 0.3% accuracy, while only a 100 counts are required for 10% accuracy). Photon density is often equated (incorrectly) to information density. Although this is usually true within a given field, it is not typically true for comparisons between different fields. For example, it is not a true statement that since roentgenographic images

*In x-ray CT, dose must be increased to improve resolution as given above, but in ECT, the patient dose can remain constant and the measurement time increased, since the activity injected into the patient typically produces the same dose whether the time required for the measurements is short or long.

EMISSION COMPUTED FOMOGRAPHY

have more photons than nuclear medicine ones they have more information. It has been shown that the error propagation in C T is p r o p o r t i o n a l to path length. ~9,24,3~In ECT this is further worsened by the fact that photon attenuation also increases as the path length across the object increases. Since all cross-sections of the human body can be described as either circles or ellipses, the path length increases from edge to center of the cross-section, and therefore, the accuracy in the final reconstructed image (after correction for attenuation) will be worse at the center than at the edges. This error propagation problem is particularly alarming, since one of the advantages of ECT is its ability to accurately detect and resolve internal structures as opposed to those in the surface where presently available s y s t e m s h a v e t h e i r g r e a t e s t sensitivity. However, this effect can be compensated for by systems whose efficiency increases towards the center of the object. The redundant sampling in the h e x a g o n a l design of the P E T T III, O R T E C / E C A T , and P E T T IV (also possible in the positron camera of Muehllehner) accomplishes this to a significant degree because the fan-beam geometry of these systems inherently increases efficiency from edge to center of the object, l* Photon attenuation also decreases statistical accuracy. Budinger et al./6 have recently shown by theoretical calculations that the total attenuation for the 140-keV photons from 99~Tc is less than for 511-keV photons when ACD is employed. However, it should be remembered that the major portion of this increase for 99mTc originates from the surface of the object, which is less important in terms of accuracy in ECT and for structures that lie interiorly. In fact, increased emission from the surface can obscure internal variations. The statistical accuracy is of course only one of many factors that determine overall image accuracy. The fan-beam geometry of ACD (i.e., the ability of one detector in ACD to record the radiation emitted in many different directions) affords a tremendous increase in geometric detection efficiency for ACD over SPC. ~4,~ However, the exact efficiency gain of ACD over SPC is dependent on instrument design and must await comparison between systems that are presently developed or under development.

343

A major source of inaccuracy in ECT results from the propagation of error from the attenuation correction methods into the final reconstructed image. This error is present in all ECT systems, since there is no perfect attenuation correction method. However, at this time, ACD attenuation correction methods produce less error and allow more exact attenuation correction than methods used in SPC. INSTRUMENTATION

A number of ECT systems have been or are being constructed at this time (Table 3). These systems can be separated into two categories: (1) systems that employ SPC, such as scanners and scintillation cameras for 9~mTc, ~~ ~23I, ~3~I, etc. and (2) systems that employ ACD of positron-emitting radionuclides. There are approximately an equal number of systems that have been constructed in these two categories. Due to the scarcity of detailed information about every one of these systems and the fact that they are in a rapid state of development and evaluation, the presentation of each will be limited to the more general aspects with examples shown of typical studies.

Single-Photon Counting Systems Mark IV. It is only fitting that a list of ECT systems begin with Kuhl's M A R K IV Scanner since Kuhl and his co-workers as early as the late 1950s began to develop the principles of ECT and have exemplified this through the construction, testing, and application of the M A R K I, M A R K II, M A R K III, and now M A R K IV. The M A R K IV scanner (Fig. 5) consists of a square array of 32 NaI detectors, which are each 7.6 cm in height, 2.5 cm in width, and 2.5 cm in depth? Each of the four detector banks is displaced with respect to each other by }/4 an interdetector spacing, such that when 360 ~ rotation is completed, the linear sampling resolution is 8 mm instead of the inherent 32 mm interdetector spacing. Via the use of a slip ring for providing power to the system and an optical readout, the gantry of the M A R K IV can continuously rotate. A 360 ~ rotation requires 50 sec during which data are binned into angular projections (7~2 ~ and reconstructed in about 30 sec with an orthogonal tangent algebraic iterative algorithm? Since the instrument operates

344

MICHAEL E. PHELPS

Table 3. Emission Computerized Tomographs

Institution or Company

Application

Design

Detection Modet

University of Pennsylvania (MARK IV)* University of Aberdeen University of Aberdeen and J & P Engineering Co. Donner Lab. and University of California U niversity of Michigan Searle and Baylor University Searle and Baylor University Union Carbide

Head Body Body Body Body Head Body Head

Square array scanner Square array scanner Dual head scanner Single head camera Single head camera Single head camera Dual head camera Twelve-sided scanner

SPC SPC SPC SPC S PC SPC SPC SPC

UCLA and ORTEC Inc. (ECAT) ORAU and ORTEC Inc. (ECAT) Washington University (PETT II1w IV) Massachusetts General Hospital and Cyclotron Corp (PC I and II) UCLA (CRTAPC) Brookhaven National-Montreal Neurological Institute 82 Lawrence-Donner Lab., Univ. of California Searle and University of Chicago University of California, San Francisco

Body Body Body Body Body Head Body Body Body

Hexagonal array Hexagonal array Hexagonal array Dual head multiple detector Circular array Circular array Circular array Dual head camera Dual head MWPC camera

ACD ACD ACD ACD ACD ACD ACD ACD ACD

* Presently at UCLA. t SPC represents single-photon counting as employed in conventional scanners and scintillation cameras. ACD represents annihilation coincidence detection of positron emission. No longer in use. New dual head camera is being constructed. w at Brookhaven National. ~[Presently at Montreal Neurological Institute.

at 50 sec per revolution, the continuous collecting and processing of data produces a new, updated cross-sectional image every 50 sec as the study progresses. The user can allow the system to continuously rotate while examining the progressive development of the image and can stop the rotation when the image has reached acceptable quality. In practice, a fiverevolution (4.2 min) scan is used in most clinical studies. The data are displayed with a memory buffered video display system in a 64 • 64

Fig. 5.

format for viewing and photographing. The system employs converging collimators on each detector, and attenuation corrections are applied by the use of an empirical relationship discussed in section, Physical Factors and Assumptions. This system has been investigated in detail by performing a wide variety of studies to determine the quantitative accuracy and linearity of the reconstructed data with excellent results? The M A R K IV is designed and applied exclusively to the head. The design concept of

The M A R K IV scanner in the clinical area of the Division of Nuclear Medicine, UCLA (courtesy of D. E. Kuhl).

EMISSION COMPUTED TOMOGRAPHY

345

Fig. 6, Representative studies with the MARK IV. (Top)~Qm]'c-pertechnetate with a patient with a cerebral hematoma. The region of interest (B) and histogram (C,D) are illustrated. (Bottom) Cross-sectional images of the cerebral metabolic rate for glucose using lSF-2-deoxyglucose administered intravenously in a normal human subject, Images show good delineation of the high metabolic activity in the superficial cortex and internal grey nuclei relative to l o w values in subcortical white matter. Scan parameters were : resolution of about 1.6 cm, scan time of 15 rain/slice, approximately 700,000 counts/image, and about 7 mCi were administered intravenously. Insert shows distribution of activity in the brain with a lateral scan taken with a conventional rectilinear scanner.

346

the M A R K IV is to apply maximum detection efficiency in the transaxial plane. The M A R K IV has been used with 99mTcO4 (Fig. 6) and 99mTc-RBC to investigate changes in the blood/brain barrier (BBB) and cerebral blood volume, s,33 respectively, in patients with a wide v a r i e t y of c e r e b r a l p a t h o l o g y . For example, studies have been carried out to examine the degree and extent of breakdown of the BBB and to determine the amount of vasodilatation or constriction in or surrounding the focal lesion, 8 or in diffuse injury. 34 Studies have also been carried out to examine local changes in BBB p e r m e a b i l i t y p r e - a n d p o s t t r e a t m e n t with steroids. 36 lz3I-iodoantipyrine has been investigated for the detection of focal ischemia in stroke patients? Although this method has clearly demonstrated examples of cerebral infarction, the single impulse injection method is limited, since antipyrine rapidly diffuses out of the tissue and recirculates towards a uniform distribution in the brain that does not represent cerebral blood flow (CBF). 8 However, different methods of administration and different tracers are being investigated for the determination of regional CBF. The M A R K IV scanner has been used with 99mTc-diphosphonate to image the cross-sectional distribution of the skull and with ~lInD T P A for cisternography. 8 More recently, this system has also been employed to image and to calculate the regional distribution of cerebral glucose metabolism with lSF-2-deoxyglucose in human volunteers. (Fig. 6) 34 Aberdeen scanner. Several SPC scanning tomographs have been developed by Mallard and co-workers in Aberdeen, Scotland. These scanners have been very similar in concept to the systems developed by Kuhl and co-workers. The present Aberdeen scanner 36 consists of two N a I detectors with converging collimators that scan in a linear direction and rotate in a similar manner to the M A R K II developed by Kuhl. The Aberdeen group, in cooperation with J • P Engineering in Scotland, are also constructing a multiple detector square array scanner similar to the M A R K IV design of Kuhl and coworkers. This system consists of a square array of N a I detectors and rotates in a continuous fashion about the patient and is to be used for whole-body studies. Donner Laboratories. Budinger et al. 1~

MICHAEL

E. P H E L P S

Fig. 7. The scintillation camera and rotating stand configuration employed by Donner Lab for ECT. The camera is held stationary and the patient is rotated through 3 6 0 ~ (courtesy ofT. F. Budinger).

have employed a scintillation c a m e r a with a p a r a l l e l - h o l e c o l l i m a t o r for p e r f o r m i n g transaxial CT. Data are collected by rotating the patient through a 360 ~ angle in front of the camera (Fig. 7). Data from the c a m e r a are digitized with analog to digital converters set to give a linear sampling distance of 4 mm. Depending on the type of study, anywhere from 36 to 144 angular projections are taken during a 360 ~ rotation (10 ~ ~ between projections). These data are then processed with either an iterative leastsquares or a filtered back-projection algorithm for reconstruction of the cross-sectional image. Correction for photon attenuation is discussed in the section, Physical Factors and Assumptions. The collected data are processed in either a H e w l e t t / P a c k a r d 5407 minicomputer or a CDC 7600 and displayed in a final 64 x 64 array format. This system has been employed for a wide variety of phantom studies to investigate resolution, contrast, sensitivity, attenuation correc-

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Fig. 8. (A) Cross-sectional images of the thorax of a 64year-old patient taken with the system shown in Fig. 7, subsequent to the intravenous injection of 4 mCi of =2~ Posterior wall and posterior septum regions show poor uptake consistent with EKG findings. Dome of the liver is seen in sections 2 and 3. (B) Cross-sectional images of 50-year-old patient showing abnormal accumulation of isotope in t w o tumors in left posterior region of brain. Location of tumors was confirmed at surgery (courtesy of T. F. Budinger).

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MICHAEL E. PHELPS

tion approach, sampling, and to evaluate different reconstruction algorithms, lo, 16,28,37The final image resolution in clinical studies is on the order of 2 cm. 16 Studies have been carried out with Z~ 129Cs, 7'%e-methionine for gated tomographic studies of the myocardium. 3s Z~ and ~29Cs (Fig. 8) have been used in patients to investigate their usefulness for tomographic detection of ischemic disease. 75Se-methionine has been used on a preliminary basis with dogs in evaluation of amino acid kinetics in the myocardium. Studies have also been carried out with 99mTcO~ (Fig. 8) in patients with focal cerebral lesions ~~ and with 99mTc-sulfur colloid for studies of the liver. The Donner system has been employed for both head and torso studies. Humongotron. The humongotron, developed by Keyes et al., '~9 consists of a scintillation camera mounted on a cantilevered c-arm which rotates 360 ~ with a selectable speed of rotation (Fig. 9). The camera can also be moved in the in and out direction to allow for different patient detector separation distances. The scintillation camera is connected both to an on-line computer (Medical Data System, trinary system) and to a standard console for conventional imaging. Patients are positioned in the humongotron by means of a cantilevered table that is adjustable in height and angle. In the transaxial tomographic mode angular projections are taken at equally spaced angles of 6 ~ or 12 ~ (most patient studies employ 12 ~ angles) during a full 360 ~ rotation. Data are collected at discrete angular increments or during continuous rotation. These investigators indicate that equivalent images are obtained by

Fig. 9.

The humongotron (courtesy of J. W. Keyes, Jr.).

Fig. 10. of a patient seen on the high uptake Keyes, Jr.).

Cross-sectional image with 99mTc-pertechnetate with a large intercranial aneurysm that was not conventional scintillation camera image due to of activity in temporal muscle (courtesy of J. W.

both approaches. Each linear scan profile is digitized with a linear sampling distance of 4 mm and reconstructed with a convolution algorithm and displayed in a 64 • 64 format. Attenuation correction is performed by using the "mean exponential technique" developed by Kay et al. 4~When 30 angular projections of data are employed, each transaxial image requires 30 sec for reconstruction with the aid of a h a r d w i r e floating-point p r o c e s s o r and hardwired multiply/divide modifications to the computer. Typically, eight 1.5-cm thick cross-sections of the head are obtained from a single rotation of the camera in 20 30 min with 20 mCi of 99mTc04. Keyes et al., 39 have used the humongotron to examine resolution, effects of photon attenuation, and sampling with phantoms. Typical resolution of this system is about 1.5 1.8 cm for brain studies. 39 To date, patient studies 39 carried out on a humongotron have dealt exclusively with patients with focal cerebral disease using 99mTc-pertechnetate (Fig. 10). Studies are presently underway to examine the use of the humongotron for whole-body studies, particularly heart and liver?' Searle Radiographics-Baylor (S-B). The S B system (Searle Radiographics, Chicago,

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349

produces some angular blurring) did not significantly affect the image when compared to discrete angular sampling. Data are processed with a small minicomputer to reconstruct nine 16-ram thick cross-sections of the brain, The typical resolution of the cross-sectional plane is about 1.4-2 cm ( F W H M ) in clinical studies ~5'~ with this p r o t o t y p e system. The effects of photon attenuation and varying resolution with depth are partially compensated for by using the geometric mean of directly opposing projections, ~ as discussed in section, Physical Factors

and Assumptions. Fig. 11. The Searle Radiographics-Baylor scintillation camera based tomograph (courtesy J. A. Burdine and R, J. Jaszczak).

Ill.; Fig. 11) consists of a single scintillation camera mounted on a rotating gantry that revolves through a full angle of 360 ~ during a tomographic study. 15 During the continuous rotation of the gantry data are recorded every 4 ~ of angular rotation. In agreement with Keyes et al., 39 studies carried out with this system also indicate that the continuous rotation (which

Reconstructions are p e r f o r m e d by using either the two-dimensional Fourier transform method or by convolution. The data are reconstructed and displayed in a 64 x 64 format. The S-B tomograph has been used with phantom studies to evaluate methods for correcting for variation in resolution and sensitivity with depth, and evaluation of image contrast with phantom studies.15 Patient studies have exclusively dealt with the brain and have compared the contrast ratio with 99mTco4between the ECT and conventional mode for patients with focal cerebral disease (Fig. 12). These

Fig. 12. Images using ~ from a 46-year-old patient with cerebral metastatic carcinoma from lung. (Top) Conventional views of the brain with a scintillation camera. (Bottom) Three cross-sectional tomographic images taken with system shown in Fig. 11. Lesion is seen in the right parietal area (courtesy of J. A. Burdine).

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MICHAEL E. PHELPS

Fig. 13. The Union Carbide ECT system for the brain: Uniret-1 (courtesy of Union Carbide Corp.).

studies indicated that there was approximately a factor of 3 advantage in contrast in the ECT mode. 1,~ S e a r l e R a d i o g r a p h i c s and Baylor have stopped studies with the prototype system and are nearing completion of a new SPC C T ? 2 The new tomograph will consist of two directly opposing large field-of-view scintillation cameras and will be designed for whole-body studies. The data from this system will be collected and processed with an Inter-Data 8-32 minicomputer, and attenuation corrections will employ an iterative technique described by Walters et al. z9 This system is expected to be completed in mid-1977 and will be installed at the Baylor College of Medicine for clinical evaluation by Burdine and co-workers? 2 The resolution of this system is expected to be that of conventional cameras. 42 Union Carbide. This s y s t e m (Union Carbide Imaging Systems, Norwood, Mass.) consists of a 12-sided array of 20.3 cm long • 12.7 cm wide • 2.54 cm deep NaI detectors with individual phototubes and converging collimators. The detectors scan in a rectilinear fashion and move in and out (i.e., toward and away from the patient). This system is a headonly unit with a field of view of 21 cm, tilting gantry (~20 ~ with automatic and selectable capabilities for sequential scans, and variable scan times (2-5 min/slice). The scanning motion, data collection and reconstruction ( N 2 min/slice), display and postimage processing are accomplished with aid of a minicomputer (Data General Eclipse S-200). TV monitors are provided for displaying and photographing images presented in a 128 • 128 • 16 format.

This system allows simultaneous data collection and processing (Fig. 13). The first system has been installed in St. Elizabeth Hospital in Brighton, Mass. for clinical evaluation, but at this time very little information is available.

Annihilation Coincidence Detection Systems Massachusetts General Positron Camera (PC-I, PC-H). The PC-I was developed by Brownell and co-workers 3'~3'~'~ and consists of two opposing banks of NaI detectors. Each bank of d e t e c t o r s contains 127 individual crystals (2 • 2 • 3 cm in size). The design is such that each detector on one bank is connected in coincidence with the directly opposing detector and its 24 nearest neighbors on the opposite bank. Thus, even though with 127 detectors on each bank, there would be a possible total number of coincidences of 16,129 (127 • 127), only 2549 total coincidences are employed. The outputs of each bank of detectors are subsequently processed by coincidence circuits ( - 2 0 nsec resolving time), identified, coded, and transferred to a buffer memory that is subsequently read in to a PDP-9 computer for data analysis. These data are then typically displayed in a 128 • 128 array, with 64 gray levels for viewing and photographing. The camera heads are mounted on a gantry that can move in and out to change the patient detector separation distance, move up and down to increase sampling in the vertical direction, and also rotate around the patient. The PC-I has been used in three modes of imaging: (1) Conventional 2-D imaging, in which only the straight-across and slightly angulated

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Fig. 14. The Massachusetts General positron camera, PC-II (courtesy of G. L. Brownell and J, A. Correia).

lines of response (LOR) are employed. In this mode the detector banks are set at their widest separation distance to minimize the angulation errors. (2) Longitudinal focal plane tomography in which all 2549 coincidence LOR are employed by sorting them into longitudinal images that contain the plane of interest in focus, with the planes behind and in front blurred. In this mode the detector banks are moved in close to maximize the angles of coincidence LOR (i.e., improve sharpness of focus). This is the most efficient imaging mode of the system, since it

Fig. 15.

351

employs all the coincidence LOR. (3) Transaxial reconstruction tomography in which straight across LOR and some angulated nearest neighbor LOR are used when the system is rotated 180~ around the subject. These data are then processed in a convolutionbased reconstruction algorithm for the display of cross-sectional tomographic images. Inherent resolution of less than 1.5 cm has been reported with this system, 43but due primarily to sampling limitation, lower image resolution has been employed. Photon attenuation is corrected for by placing an external sheet of positron activity in front of the banks of detectors and recording the data with and without the subject between the two detector banks. A new positron camera, PC-II has been constructed by Brownell and co-workers and is presently under evaluation (Fig. 14). This system is very similar to the PC-I except that it has more crystals (144 detectors on each bank) that are placed i n a square array to better accomodate transaxial tomographic imaging. It also has a linear scanning motion in both the horizontal and vertical direction to remove the sampling limitations that were present in the PC-I. The commercial version of PC-II, which is being developed by Cyclotron Corporation (Berkeley, Calif.), is shown in Fig. 15. The PC-I and PC-II are employed in studies of the brain using ~'~NH3 for blood flow, ~'~CO2 (continuous inhalation) for blood flow, 1602 for oxygen metabolism, 6SGa-ATP for detection of brain tumors, and ~3N2 (Fig. 16) for lung and

The Cyclotron Corporation positron camera (Courtesy of Cyclotron Corporation).

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MICHAEL E. PHELPS

Fig. 16. Lung study using an equilibrium state with ~3N2. Nine cross-sectional levels are s h o w n from the top to the lower portion of the lung. The heart and diaphragm are seen as a black silhouette. Attenuation was corrected using a set of transmission images (courtesy of G. L. IBrownell and J. A. Correia).

brain flow studies in human subjects. 3,43-4'~ ~Cglucose has also been used for cerebral metabolism. This system has also been employed for studies of the h e a r t using 13NH:~ and 82Rb43'4~,46 and for the lung (infarction) using 6SGa microspheres 43 and ~SF for bone imaging? The latter studies have been carried out in animals. 8~Rb has also been employed to image myocardial infarcts in dogs. 43,46The PC-I and II are employed for whole-body studies (Fig. 16).

Fig. 17. Dual:headed scintillation camera positron t o m o g r a p h developed cooperatively by Searle Radiographics- University of Chicago (courtesy of G. Muehllenhener and P. V. Harper).

Searle Radiographics-University of Chicago positron camera. This system (Fig. 17) has been developed in a cooperative effort between M e u h l l e n h n e r and H a r p e r with their coworkers. It consists of two directly opposing large field-of-view scintillation cameras that are mounted on a gantry for continuous rotation through 180 ~ The cameras use the highspeed electronics of the Searle large field-ofview camera with some modifications to opti-

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353

Fig. 18, (A) Cross-sectional images of blood v o l u m e distribution in brain subsequent to inhalation of a p p r o x i m a t e l y 2 mCi of '~CO. I m a g e s w e r e taken a p p r o x i m a t e l y 15 rain after a d m i n i s t r a t i o n and 1.8-mitbon total counts w e r e collected over an i m a g i n g t i m e of 2 0 min. (B) Longitudinal sections of brain subsequent to the single breath inhalation of a p p r o x i m a t e l y 2 mCi of I C O . Data collection was initiated 10 rain after inhalation, and a p p r o x i m a t e l y a m i l l i o n total counts w e r e collected. Images f r o m left to right proceed longitudinally f r o m posterior to anterior (courtesy of G. M u e h l l e n h e n e r and P, V. Harper.),

mize detection of 51 I-keV photons. In addition, the NaI crystals are 1-inch thick to improve detection etEciency. The outputs of the two camera heads are processed for both pulseheight analysis and coincidence with a 15 nsec coincidence resolving time. All possible lines of coincidence between opposing detector heads are collected in a serial format on magnetic

tape. These data are then sorted appropriately for different imaging formats. This system has high sampling resolution in both the transverse and axial direction. As discussed for the BrownelI PC systems, this system also allows imaging in the three t'orm a t s of two-dimensional projections, focal plane tomography, and transaxial tomography.

354

However, to date, only the latter two tomographic imaging formats have been employed. B e c a u s e all c o m b i n a t i o n s of c o i n c i d e n c e between opposing banks of detectors can be identified and recorded, sampling flexibility is inherent in this system for focal plane and transaxial tomography. In addition, algorithms similar to that developed by Chang et al. 47 are being employed to produce longitudinal tomographic images that compensate for superposition of information (i.e., reconstruction of longitudinal planes). 48 This system is therefore capable of simultaneously generating multiple longitudinal planes in a stationary position. Studies are also being carried out to improve the efficiency of transaxial reconstruction tomography with this system by employing the angular lines of response that are outside a selected transaxial plane of interest, but which cross through it at different angles. 48 The resolution of this system is about 1-1.4 cm. Patient studies have been carried out with ~3NH3 and SlRb (with comparisons to 2~ for the detection and delineation of myocardial ischemia; 49 68Ga-EDTA for the detection of focal cerebral lesions; bone studies with baboons using 18F;22.23 and cerebral blood volume studies 48 (Fig. 18) in human subjects using inhalation of 1lCO. P E T T H and HI. The positron emission transaxial tomographs, PETT II and PETT III, were developed by Phelps, Hoffman, Ter Pogossian, and co-workers at Washington University.9,14,17 20,20 The P E T T III (now at Brookhaven National Laboratory, Long Island, N.Y.) consists of a hexagonal array of 48 5 • 7.5 cm NaI detectors in which each detector on opposing banks is connected in multiple coincidence (Fig. 19). This system was specifically designed for performing quantitative high-" contrast transaxial CT. The gantry performs a linear scan of all banks in synchrony (5 cm distance) and discrete angular rotation (3 ~ ) through a full angle of 60 ~ for a complete scan. Data are sorted, corrected for photon attenuation, and reconstructed with a convolutionbased reconstruction algorithm with an Inter Data minicomputer for display in either a 50 • 50 or 100 • 100 format on a memorybuffered video. Both linear and angular sampling is variable and selectable to allow optimal sampling for reconstruction. Slit shielding is

MICHAEL E. PHELPS

Fig. 19. PErT III (presently at Brookhaven National Laboratory).

provided to minimize coincidence count rate from scattered radiation and singles count rate that produces random coincidences. The coincidence resolving time of this system is approximately 25 nsec. The image resolution is variable from about 1.4 to 2.5 cm F W H M , but reconstruction image resolution was typically _> 2cm. Photon attenuation is corrected either by using a geometric correction (used for the head and abdomen), or by measuring the actual attenuation with an external ring source of positron activity. 9,14,19 PETT III has been employed for the following studies. (1) Brain: ~CO for cerebral blood volume (Fig. 20) and laNH3 for cerebral perfusion in patients with stroke, tumors, and normal volunteers; 17,~s,5'.~2 ~CO for cerebral blood volume (Fig. 20) and ~:~NH:~ for cerebral perfusion in patients with stroke, tumors, and normal volunteers; ~7,,s.61..52 11C_glu_ cose for cerebral metabolism in patients with stroke and normal volunteers; 9,2~ 6SGa-EDTA has been used in the study of patients with stroke and tumors. ~7-~,5~ (2) Heart: ~CO for gated studies of the ventricular chambers with volunteers; 5~ ~3NH3 for the study of myocardial perfusion in human volunteers and patients with myocardial infarction? 1 HC-palmitic acid for the detection and sizing of myocardial infarct in dogs, '~3-~ normal volunteers, and patients with myocardial infarction. ~ (3) Liver: ~3NH3 for the study of p e r f u s i o n d i s t r i b u t i o n in n o r m a l volunteers. (See Fig. 4) 18,19 ECAT-UCLA. The Emission Computerized Axial Tomograph, ECAT, (Fig. 21)

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355

l,\ Fig. 20. Cross-sectional images of blood volume distribution of the brain from a human volunteer following single breath inhalation of 20 mCi of L~CO. Resolution was approximately 2 cm and imaging time varied from 8 to 12 rain per slice with 300K-80O-K counts per slice. Study was performed with P ETT III. ~8

was jointly developed by Phelps and Hoffman with O R T E C Inc. and consists of an hexagonal array of 66 3.8 • 7.5 cm N a I detectors in which all detectors on opposing banks are connected in multiple coincidence? 6 Linear (4 cm) and angular (5 ~ ) scanning are carried out through an angle of 60 ~ The data are then sorted, cot-

Fig. 21.

r e c t e d for p h o t o n a t t e n u a t i o n , and reconstructed with a convolution-based algorithm with an on-line P D P l l / 4 5 m i n i c o m p u t e r . Cardiac gating is also provided in the system. Data are stored on floppy disc and are displayed with a memory-buffered video display system in a 100 x 100 or 200 • 200 format with 64 gray

The ORTEC ECAT in the Nuclear Medicine Clinic, UCLA.

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MICHAEL E. PHELPS

Fig. 22. Rectilinear and tomographic scans taken with UCLA-ECAT. Top: (A) Whole-body rectilinear scan with human subject subsequent to single breath inhalation of 10 mCi of 11CO showing the simultaneous recording of A-P and 60 ~ left and right oblique views of the blood volume distribution with HC-carboxyhemoglobin. Total scan time, 20 rain; resolution, 1.4 cm. (B) Cross-sectional images of cerebral blood volume starting at 6 cm above the orbital meatus, OM, (top left hand corner) and moving left to right to 3 cm below the O i in 15-mm steps. Top is anterior and left side is at left. The major structures seen in the images are superior saggital sinus, straight sinus, and high blood volume around the Sylvian fissure and internal capsule, the lateral sinus, the left and right corotid and cavernous sinuses, sigmoid sinuses and the high blood volume in the nasal area. Note that the superior saggital sinus is shifting to the right in images 3 and 4. Resolution is 9.5 mm (compare size of vessels to fig. 20). Bottom: Study of blood pool in heart chambers using HCO administered by single breath inhalation. (Left) Transmission images; at bottom are three limited field rectilinear transmission scans of the thorax in the rectilinear scan mode. The X indicates the position chosen with the joy stick to initialize the tomographic sequence. Above are shown six transmission transaxial images from base of heart towards the apex in 15-mm steps. Anterior is at top, and left and right are reversed. (Right) Emission images; at bottom is shown the emission limited field rectilinear scan with the positions selected (shown as the X) for initializing the sequence of ECT. Above are shown six ECT images that correspond to the same positions as transmission images shown at left. Image 1 shows the right ventricle, left ventricle, left ventricular o u t f l o w tract, aorta, superior vena cavae. The second image shows the right ventricular outflow tract, right atrium, left ventricle, left atrium, pulmonary artery, and aorta. The intraventricular septum is clearly shown in images 4, 5, and 6, and in image 6, the inferior vena cavae is appearing. The resolution is 1.4 cm, and the images were ungated. Approximately 10 mCi of HCO was administered with a total scan time for all slices of approximately 20 min with 1-3.5 million counts per image.

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levels or color. The patient is placed on a bed that can be moved in the up and down and in the axial directions either manually or under automatic control of the on-line computer, so that a complete study can be carried out automatically. This system was designed as a highcontrast, high-efficiency, high-resolution, quant i t a t i v e f a s t - s c a n n i n g , single-slice s y s t e m representing an improved and optimized version of PETT III. Slit sheilding is provided to minimize coincidence from scattered radiation and singles count rate that produce random coincidence. The coincidence time resolution is 17 nsec. The E C A T can also perform rectilinear scans, either whole-body or limited field, (Fig. 22) in which three simultaneous views (0 ~ and • ~ ) are recorded and displayed. The rectilinear scan mode can be used alone or can be used to select planes for tonaography by simply indicating the levels of choice on the rectilinear scan with a joy stick (Fig. 22). Both the tomographic and rectilinear scans can be corrected for attenuation with either an exterr~al ring source of positron activity or an automatic geometric correction. TMtg Scan times are selectable from 10 sec to multiple minutes per slice in the transaxial mode. Resolution is also variable and selectable in both the transaxial and the rectilinear mode from about 9 mm to 1.8 cm. The electronics design employs high-speed modular electronics for high data rates, low dead-time, flexability, and expandability. This system has been applied in a transaxial mode to the following studies. (1) Brain: 11CO for cerebral blood volume in normal volunteers (Fig. 22), patients with stroke and tumor; ~aNHa for blood perfusion in patients with stroke and tumors; 68Ga-EDTA for BBB evaluation of patients with stroke and brain tumors; ~SF-2deoxyglucose in normal volunteers for the measure of stimulated response, patients with stroke and focal cerebral lesions. '~r'S* Studies have also been carried out with each of these tracers in rhesus monkeys and dogs to determine the quantitative uptake and distribution of these tracers. (2) Heart: ~CO for gated studies of ventricular chambers and vascular in normal volunteers (Fig. 22); 'aNH3 for the measure of perfusion distribution in normal volunteers and patients with myocardial infarc-

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Fig. 23.

PETT IV (courtesy of M. M. Ter-Pogossian).

tion; l~C-paimitic acid in p a t i e n t s with myocardial infarction; ~SF-2-deoxyglucose in n o r m a l v o l u n t e e r s and in p a t i e n t s with myocardial infarction. Each of these tracers has also been used in dogs and rhesus monkeys to quantitate the uptake, distribution, and retention. (3) Liver: 62CUS, ~ G a (OH)~, and ~3NH3 for imaging of the liver in dogs and patients with liver disease. (4) Bone: ~SF in patients with metastatic bone lesions and in dogs. All of the above tracers have been used also for whole-body or limited field rectilinear imaging in human subjects and animals. O R A U - E C A T . An O R T E C - E C A T positron tomograph has also been installed at Oak Ridge Associated Universities (ORAU) and is presently being employed for studies with 6SGaEDTA in patients with stroke and cerebral tumors; 68Ga-citrate and ~SF for absecess and bone studies in dogs? 9 Studies are also underway to use ECAT for the evaluation of 11Camino acid in cancer patients. P E T T IV. The PETT IV (Fig. 23) is being developed by Ter Pogossian, Mullani, Higgins, C u r r i e , and c o - w o r k e r s at W a s h i n g t o n University. zl This system is similar in design to the PETT III, with a notable exception that the 48 detector array contains Nal crystals that are 5 cm thick by 17 cm in length in the axial direction. This is a 4-slice system in which the slices are defined by 3 lead slit shields that are approximately 8 mm thick. The identification of events in each of the 4 slices is determined by a position logic from phototubes placed on opposing ends of the detectors. Data are collected, sorted, corrected for photon attenuation, and

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MICHAEL E. PHELPS

Fig. 24. Cross-sectional images of distribution of cerebral blood v o l u m e (CBV) f r o m study w i t h PETT IV. A p p r o x i m a t e l y 15 m C i of " C O w a s a d m i n i s t e r e d by inhalation. Resolution w a s 1.6 cm in the transaxial plane and 1.6 cm in slice thickness. Total i m a g i n g t i m e for 16 slices w a s 16 m i n w i t h a m a x i m u m n u m b e r of counts of 8 7 4 , 0 0 0 (#1) to a m i n i m u m n u m b e r of counts of 9 5 , 0 0 0 (#16). The position of the 15 levels are indicated on the f i g u r e and are o b t a i n e d by indexing the patient four times. The large blood vessels and general d i s t r i b u t i o n of C B V are seen in the images (courtesy of M. M. Ter-Pegossian).

reconstructed by an on-line Inter Data 7/32 minicomputer for display on a video buffer display system. This system is designed for wholebody studies. The construction of this system has recently been completed and initial evaluations are in progress. The PETT IV will be used for studies similar to those described for PETT III (Fig. 24).

B r o o k h a v e n - M o n t r e a l Circular Tomograph. This system was initially developed by Robertson et al. ~~ approximately 12 years ago and consists of 32 N a I detectors (3 x 2.5 cm) in a circular array. The initial success of this system was limited because of a lack of a suitable r e c o n s t r u c t i o n algorithm. It was subsequently loaned to the Montreal Neurological Institute and is now interfaced to a PDP-12 for data reconstruction either with or without a small amount of rotation (6 ~ small angle rotation is employed to improve linear sampling and image resolution). This system has been used for the first dynamic positron tomography with images taken every 4 5 sec with 68Ga-EDTA

(intravenous injection) and every 20 30 sec with rrKr (inhalation). The regional time curves from these two tracers are then used to estimate cerebral vascular mean transit time and CBF. Each level studied requires a separate administration of tracer. Multiple-level studies are not indicated. Although they would be possible with rrKr, the build-up of ~SGa-EDTA would be a problem. Approximately 120 patients with cerebral tumors, arterial-venous malformation, and stroke have been studied. Selected results in patients have demonstrated abnormal flow patterns with minimal abnormality in CT scans (stroke) and indicated the "steal" phenomenon in a patient with an AVM. al Some encouraging results are shown with the system (Fig. 25), even though the technique is limited in temporal, spatial, and statistical resolution21 These investigators are planning to i m p r o v e the present resolution ( - 6 cm without rotation and 3 cm with rotation) by a factor of 2 by replacing the 32 N a I detectors with 64 Bismuth Germanate detectors of half the size. 61 This system is for the head only.

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Fig. 25. Brookhaven-Montreal circular positron tomograph (top.) (Bottom) Cross-sectional images of cerebral blood flow employing inhalation of 77Kr. Left and right images are pre and post anastomotic operation showing the reinstitution of cerebral blood flow to the ischemic area indicated by arrow (courtesy of L. Yamamoto).

1.1 cm, the actual image resolution can be somewhat better than given above (probably about 3.8 and 1.8 cm). Data from the C R T A P C are collected, sorted, corrected for attenuation by an external ring of activity, and reconstructed with a minicomputer for display and photographing. The C R T A P C has been used in a variety of phantom studies ~2 to examine the effects of contrast, resolution, and efficiency and has been employed in preliminary human and dog studies to evaluate the use of '3N-alanine for studies of the pancreas and heart (Fig. 27) and 81Rb for heart, e3 This system is designed for whole-body dynamic and static studies. Lawrence-Donner Laboratories ( L D L ) positron camera. The LDL positron camera was developed by Derenzo, Budinger, and coworkers j6.64,65 at the Lawrence-Berkeley Donner Laboratory (Fig. 28). This system basically consists of 280 NaI crystals (8 mm wide • 3 cm high x 5 cm deep) in a circular geometry with a 31-cm diameter field of view (FOV). The total number of coincidence lines of response is 14,000. This system is completely stationary with major design objectives of performing dynamic

CRTAPC. A Circular Ring T r a n s v e r s e Axial Positron Camera (CRTAPC, Fig. 26) is being developed by Cho and co-workers at UCLA. 62,63 This system consists of a circular ring of 2 x 3.8 cm thick NaI crystals. Each detector on the circular ring is connected in coincidence with 23 detectors on the opposite side such that there are a total of 736 total coincident lines of response (LOR). The coincidence L O R are sorted into 64 angular projections (i.e., 2.8 ~ angular sampling) with 12 linear scan points each. The field of view covers a 30-cm diameter and with a linear sampling resolution of about 2.5 cm (i.e., the interdetector spacing). If the system is rotated over an arc length of half the interdetector spacing (i.e., a 2.8 ~ rotation) the data can be sorted into 64 projections that now have 23 points per linear scan or a linear sampling resolution of about 1.25 cm. Since image resolution is typically twice the value of the linear sampling, the resolution without and with the half arc rotation would be about 5 and 2.5 cm, respectively. However, since the inherent detector resolution is about

Fig. 26. The Circular Ring Transverse Axial Positron Camera (CRTAPC) at UCLA (courtesy of Z. H. Cho).

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Fig. 27. Tomographic human heart study with CRTAPC employing ~3N-L-alanine injected intravenously, which shows clear visualization of the left ventricle. Gating is not employed (courtesy of Z. H. Cho).

tomographic studies and allowing for high tomographic resolution. The transaxial resolution has been determined using a computer simulation from data collected with 16 detectors. The resolution at the center of the FOV is circular with a F W H M of about 8 m m and becomes elliptical away from center with about 8 • 13 m m and 8 • 19 m m resolutions at 10 and 15 cm from the center of FOV. 1~'64 This variation in resolution results from the particular sampling

Fig. 28: The Lawrence-Donner Laboratories circular ring positron camera (courtesy of S. E. Derenzo and T. F. Budinger).

MICHAEL E. PHELPS

Fig. 29. The multiwire proportional chamber positron camera (courtesy of L. Kaufman).

scheme and geometric response employed in this system. The L D L tomograph also has variable slice resolution that is selectable from several m m to 15 mm, and the gantry can tilt an angle of • 20 ~ Completion of the L D L tomograph is scheduled for mid-1977, and it will be employed for whole-body studies. 66 MWPC. The multiwire proportional c h a m b e r ( M W P C ) p o s i t r o n c a m e r a was developed by Kaufman, Perez-Mendez, and coworkers 67'68 and consists of an opposed pair of l a r g e - a r e a multiwire proportional c h a m b e r s separated by about 50 cm (Fig. 29). Each of the chambers is 48 • 48 sq cm sensitive area and are connected in coincidence. Each chamber contains a honey-comb configuration of lead converters in close proximetry to a grid pattern of wires that are used for the detection and positioning of each detected event. When a photon strikes the lead honey-comb converters, conversion electrons are ejected that ionize the counting gas and are positioned within discrete cells of the M W P C c h a m b e r . No pulse-height analysis is employed. This system has been demonstrated to have high spatial resolution of approximately 7 mm F W H M with a 10-cm lucite scattering material. ~8 However, due to the poor time resolution ( ~ 3 0 0 nsec) it is limited in count rate capability to about 300 cps. 67 While the M W P C has a large geometric efficiency, the intrinsic stopping power of the M W P C is very low.

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This MWPC has been employed primarily for focal plane tomography similar to that employed by Brownell et al. 3 and Muehllenhner et al. 22 However, these investigators are presently evaluating the use of longitudinal reconstruction tomography with an iterative algorithm2 9 At present, the MWPC is used with ~SGa-DTPA for the study of focal brain lesions. SUMMARY

At the present time, the instrumentation for emission computed tomography is in the stage of rapid growth and development and therefore does not lend itself to explicit definition. Systems are being developed in academic institutions and by commercial companies for both SPC and ACD types of CT. Presently, approximately an equal number of ECT systems are in each of these two categories (Table 3). Each of these systems contain different design and application strategies that will broaden our understanding and better define the optimal approach to ECT. No explicit comparisons are provided between all of the systems described in this article, since a standardized method for comparison has not been developed. In addition, the systems described vary in stage of development from being in the middle of construction to completion and application. Since ECT systems vary to such a great degree in design, there are some general aspects that deserve mentioning for the future comparison. Phantom studies must be carried out that realistically portray the environment (i.e., human studies) in which the systems will be finally employed. It is a natural tendency of all instrumentation developers to design test phantoms that are optimal for their systems. These test phantoms may or may not reflect the actual problem for which they will be applied. However, the phantoms should reflect a similar condition for photon absorption and scatter, object size, activity distributions, activity densities, and type of study to be performed (i,e., static, dynamic, quantitative, qualitative, etc.). Spatial resolution and efficiency should be quoted under the same and realistic conditions. Clinical trials must be carried out for final evaluation of system parameters and methodological approaches. Since ECT involves a mathematical

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reconstruction of an image, there is a great deal of latitude in image formation, particularly in terms of resolution and contrast. System design is always composed of many compromises and trade-offs. Improvements in one system parameter or imaging mode is done at the expense of another; for example, resolution improvement at the expense of efficiency and increased image noise, tomographic efficiency at the expense of 2-D imaging capability and vice versa, faster scan time at the expense of lower resolution and lower accuracy, etc. Thus, the user must clearly appreciate each of the factors and make his decision in accordance with his priorities. The relative merits of single-slice versus multiple-slice systems must be considered carefully for each application. It should not be taken as an obvious conclusion that 2, 4, 10 etc., slice systems are 2, 4, 10, etc. times better than a single-slice ECT. One must compare the overall efficiency and time for each system to carry out a complete exam. For example, a single-slice system may have 2 10 times the efficiency, scan speed, and capability per slice of a multiple-slice system. The parameters that affect image quality (scattered radiation, count rate capabilities, random coincidence, etc.) and the flexibility and importance of different imaging modes must be considered. It should also be remembered that CT has a very particular type of error propagation and artifact generation that is a function of system design. For example, the data collected for each transaxial plane are interdependent and also independent of surrounding planes. Therefore, from this point-ofview, a system that rapidly collects the data for each plane is preferred over a slower system that simultaneously collects the data for multiple planes, even though both systems may require the same time for a total organ exam. For example, rapidly collecting the interdependent data for each transaxial plane reduces the probability of distortions from patient, organ, or activity movement. Of course, a system with maximum efficiency per plane, multiple plane and multiple imaging modes is advantageous if the design is cost effective and optimal image quality is maintained. The commercial availability at this time is equally split between the two techniques of SPC

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and ACD tomography. Searle Radiographics, J & P Engineering Co., and Union Carbide Imaging have recently indicated that they will be supplying systems in the former category. While ORTEC Inc. and Cyclotron Corp. (also possibly Searle Radiographics, although at this time they are undecided) are supplying systems in the latter category. There are advantages and disadvantages to both SPC and ACD tomography. More definitive data are needed to clarify the individual points of concern. As has been discussed by many investigators, ACD tomography has the advantages of higher efficiency in transaxial and longitudinal reconstruction tomography, exact or near-exact correction for photon attenuation, and uniform resolution and sensitivity that is relatively independent of object size. However, the degree to which these factors allow one technique to dominate another and the role of these criteria in the overall technique are yet to be clearly established. The deciding factors for both of these techniques or between them can probably be better appreciated by trying to define the purpose of ECT. 7~ For example, will it be sufficient to add the advantages (some of which are yet to be determined) of ECT to the conventional studies that are now performed in nuclear medicine? Will it be diagnostically important at the added cost of ECT to improve image contrast and spatial description of lesions? Will ECT improve detection rates? Will it provide more definitive information about the disease entity and change the course of patient management? These are all questions that are still unanswered today but which must be dealt with in the face of the rapid growth of x-ray CT, ultrasound, and the increasing concern over rising medical costs. The question should also be asked as to what ECT allows one to do that couldn't be done before or that can't be done by other techniques. Thus, along with instrumentation development there must be a development of purpose. The unique feature of ECT is the ability to quantitatively measure tissue activity concentration of selected tracers for the measurement of specific physiologic functions. The achievement of this goal rests not only on the development of truly quantitative instrumentation, but also on the development of

MICHAEL E. PHELPS

radiopharmaceuticals and appropriate physiologic models consistent with the criteria of ECT. '5~,7~Thus, if we accept this approach, the type and availability of suitable radiopharmaceuticals will play a major role in determining the success of ECT and choice of instrumentation. SPC has a distinct advantage in that readily available radiopharmaceuticals can be used directly. This is offset to some degree by the availability of positron-emitting isotopes of 6SGa (from 6SGe generator), ~SF, and the possible development of other generator systems. However, new compounds will have to be developed for the measurement of physiologic or functional processes in the body to exploit fully the unique capabilities of ECT. The isotopes of 11C, 13N, and ~50 provide the capability to synthesize a wide variety of labeled natural substrates, analogs, and drugs by biosynthetic (enzymatic and photosynthetic) as well as classical chemical organic reactions. Compounds labeled with these positron-emitting radionuelides can potentially provide the means to measure a broad variety of metabolic, transport, and hemodynamic processes not possible by any other technique. This aspect lends support to the concept of developing an accelerator-based generator system for the routine production of these radionuclides. Much of the success of positron tomography will depend on whether this barrier for delivery on a routine clinical basis is broken. While this seems like a difficult task, significant advances in any field rarely come easily. ECT clearly and uniquely provides the capability to develop a new and important diagnostic technique of physiologic or function tomography as opposed to the morphologic tomography with x-ray CT. The potential significance of ECT, if it can be brought to full fruition is unquestionable, whether it be by SPC or ACD. ECT should be recognized as a technique that can provide a dramatic advancement of the original concept of nuclear medicine, that is the use of radionuclides for the measure of metabolism and physiologic function. ECT is in need of three things to answer and clarify many questions that exist today: (1) expanded technical effort, (2) clear definition of purpose, and (3) clinical results that provide information

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not obtainable by other techniques and that clearly demonstrate that its importance exceeds its cost.

ACKNOWLEDGMENT

I would like to thank all of the investigators who kindly provided figures, photographs, and time for discussion.

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