CA2077735A1 - Over-sampling pet tomograph operating in stationary mode - Google Patents
Over-sampling pet tomograph operating in stationary modeInfo
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- CA2077735A1 CA2077735A1 CA 2077735 CA2077735A CA2077735A1 CA 2077735 A1 CA2077735 A1 CA 2077735A1 CA 2077735 CA2077735 CA 2077735 CA 2077735 A CA2077735 A CA 2077735A CA 2077735 A1 CA2077735 A1 CA 2077735A1
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- Prior art keywords
- gamma
- detectors
- ring
- pairs
- rings
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Classifications
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/29—Measurement performed on radiation beams, e.g. position or section of the beam; Measurement of spatial distribution of radiation
- G01T1/2914—Measurement of spatial distribution of radiation
- G01T1/2985—In depth localisation, e.g. using positron emitters; Tomographic imaging (longitudinal and transverse section imaging; apparatus for radiation diagnosis sequentially in different planes, steroscopic radiation diagnosis)
-
- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
- A61B6/02—Devices for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
- A61B6/03—Computerised tomographs
- A61B6/037—Emission tomography
Abstract
ABSTRACT OF THE DISCLOSURE
The PET apparatus comprises first and second coaxial and laterally adjacent rings of gamma-ray detectors surrounding a subject under study. The detectors of the first ring are angularly offset with respect to the detectors of the second ring. In the apparatus, (a) linear trajectories interconnect in a first direct plane pairs of detectors of the first ring, (b) linear trajectories interconnect in a second direct plane pairs of detectors of the second ring, and (c) linear trajectories interconnect in two cross planes pairs of detectors each formed of one detector of the first ring and one detector of the second ring.
Radioactive isotopes are injected into the subject and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions. When the gamma rays of a pair propagate in opposite directions along one of the trajectories, they are detected by the corresponding pair of detectors. A data acquisition circuit stores data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated, and a computer combines these data in a single plane of reconstruction so as to reconstruct in that plane a distribution of sources from which the detected pairs of gamma rays originate.
Adequate sampling is provided to achieve maximum resolution of the reconstructed image in stationary mode.
The PET apparatus comprises first and second coaxial and laterally adjacent rings of gamma-ray detectors surrounding a subject under study. The detectors of the first ring are angularly offset with respect to the detectors of the second ring. In the apparatus, (a) linear trajectories interconnect in a first direct plane pairs of detectors of the first ring, (b) linear trajectories interconnect in a second direct plane pairs of detectors of the second ring, and (c) linear trajectories interconnect in two cross planes pairs of detectors each formed of one detector of the first ring and one detector of the second ring.
Radioactive isotopes are injected into the subject and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions. When the gamma rays of a pair propagate in opposite directions along one of the trajectories, they are detected by the corresponding pair of detectors. A data acquisition circuit stores data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated, and a computer combines these data in a single plane of reconstruction so as to reconstruct in that plane a distribution of sources from which the detected pairs of gamma rays originate.
Adequate sampling is provided to achieve maximum resolution of the reconstructed image in stationary mode.
Description
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OVER-SAMPLING PET TOMOGRAPH
OPERATING IN STATIONARY MODE
FIELD OF THE INVENTION
The present invention relates to the field of positron emission tomography (PET). The subject invention is more specifically concerned with sampling requirements in positron emission tomography.
BACKGROUND OF THE INVENTION
Positron emission tomography (PET) uses radioactive isotopes decaying by emitting positrons (~t decay) as tracers for monitoring the spatial and temporal distribution of radioactivity within a body.
The method is useful to obtain information on dynamic processes such as the flow of a fluid within a pipe or on physiological processes within a living (animal or human) body. After emission by the radioactive nuclei, each positron annihilates with an electron to produce two gamma rays of 511 kev propagating in respective opposite directions 180 apart from each other. Diametrically opposite detectors can detect in coincidence these pairs of gamma rays, to thereby determine the trajectory on which the annihilation occurred. By superposing, by means of the 207773~
conventional techniques of tomography reconstruction, the plurality of trajectories detected in a plane by an array of detectors surrounding the subject under study, a distribution of annihilations (or sources of emission of pairs of gamma rays propagating in opposite directions) in the subject's body can be derived. A three-dimensional image can also be obtained by juxtaposing the two-dimensional images of the distribution of annihilations in many adjacent planes. The principles and advantages of PET in performing quantitative measurements of the biochemistry and the physiology "in vivo" of organs are well known in the art, have been extensively described in the literature and, therefore, do not need to be further discussed in the present specification.
In PET, the theoretical resolution limit that can be reached is of the order of 2-3 mm FWHM
(Full Width at Half Maximum of the distribution).
Those skilled in the art know that this limit originates from the ~' emission process and the annihilation positron-electron conducting to the production of the two gamma rays of 511 kev.
Presently, most PET tomographs have resolutions inferior (that is larger) to the theoretical limit.
The current trend is to design PET tomographs achieving resolutions equal or superior (that is smaller) to the theoretical limit and therefore capable of obtaining a maximum of spatial information.
Only a few prototypes, that reach a resolution close to the theoretical limit, have been built up to now, and such prototypes are not yet available on the 20~73~
market due to the technological and economical constraints which impose many compromises.
A typical PET camera comprises at least one ring of gamma-ray detectors forming a cylinder around the subject under study. With this configuration, any linear trajectory interconnecting a pair of diametrically opposite detectors and passing through the region of interest of the subject's body can be monitored. The measurements performed by a PET
camera therefore correspond to a series of linear trajectories interconnecting points, more specifically gamma-ray detectors of the cylinder. These trajectories can be grouped into parallel series, called parallel projections, usable in this form to reconstruct the images. The capacity of a PE~ camera to locate a gamma-ray source in one such projection defines a function of probability of presence, referred to in the art as "coincidence aperture function" (CAF), also called "line spread function".
The so called "intrinsic resolution" of a PET camera corresponds to the FWHM of this function. In practice, the intrinsic resolution is limited by the accuracy of the detectors in determining the point of impact of the gamma rays on the cylinder.
The resolution obtained in the reconstructed image or "image resolution" further depends on the sampling frequency (number of samples by unit linear distance) in the projections.
According to the sampling theorem of Nyquist, the best resolution of a detector arrangement cannot be smaller than two times the distance between two samples.
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Accordingly, the sampling period must be at least two times as small as the FWHM of the CAF to reach the maximum resolution.
The characteristics of resolution of the CAF's and of sampling frequency of given PET cameras depend on the type of the gamma-ray detectors and of the arrangement of these detectors on the cylinder.
Many approaches, differing by their costs and their performance, are currently used by the PET cameras to extract information about the position of the gamma-ray sources. One can note three categories:
(a) ANALOG POSITION DECODING:
With this approach, an estimate of the position of interaction of the gamma rays with the detectors is obtained from the signal of many photodetectors coupled to a large scintillation crystal (scintillator). In nuclear medicine, this is the method of detection currently used in the scintillation cameras. In PET, scintillators having high light conversion capabilities and a relatively short time constant are essential to ensure the statistical accuracy of the measurement of the position of interaction and allow the detection in coincidence. Presently, only NaI(Tl) adequately satisfies these stringent requirements. However, due to the low stopping power of NaI(Tl) for 511 kev gamma rays and due to the large probability of Compton scattering into the scintillation crystal, the achievable spatial resolution is limited at a value greater than 5 mm FWHM. Although this is far from the theoretical performance of PET, this approach has the advantage of achieving an essentially continuous sampling in the projections, limited only by the precision of the circuit digitizing the position of interaction of the gamma ray within the crystal, thereby automatically satisfying the Nyquist theorem.
Only one commercial PET camera uses that approach. It has been developed by UGM Medical Systems Inc. under the name PENN-PET and reaches a spatial resolution of 5-6 mm FWHM.
(b) POSITION DECODING ON A MATRIX OF CRYSTALS:
In order to solve the problems of limited resolution and of low detection efficiency inherent to the NaI(Tl) scintillators, most of the current PET
cameras use crystals made of bismuth germanium oxide (BGO). As BGO has a low light conversion efficiency, classical analog decoding is not possible. A
tentative solution consists of employing a matrix of discrete crystals instead of a unique crystal to enable classical analog decoding. This type of configuration tends to reduce spreading of the scintillation photons by channelling them toward a smaller number of photodetectors. As the statistics of the number of photons is low, the decoding is imperfect, therefore limiting the resolution. The best prototype detectors of this type achieve a resolution of the CAF between 3 to 4 mm FWHM.
One major drawback of the use of discrete crystals is the loss of continuous sampling.
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Theoretically, the best FWHM resolution of the CAF is of the order of half the thickness of the crystal, w/2. The sampling period obtained in the useful field of view (UFOV) of the tomograph by interleaving even and odd lines of coincidence is equal to half the distance between two detectors d/2. In practical detectors formed of discrete scintillation crystals, d ~ w and the theorem of Nyquist cannot be satisfied.
As the discrete crystals are made narrower in coded detection schemes, the statistical effects and the scattering of gamma rays between adjacent crystals contribute to substantially degrade the resolution of the CAF. However, even with the CAF resolution of current systems which can be substantially larger than ~/2, the Nyquist criterion cannot usually be satisfied. For instance, in one of the above mentioned prototypes of this type (W.M. Digby, M.
Dahlbom and E.J. Hoffman, "Detector, shielding and geometric design factors for a high-resolution PET
systemn, IEEE Trans. Nucl. Sci. NS-37 (1990) 664-670), ~/2 = 1.75 ~, d/2 = 2.0 ~ and the CAF cannot have a resolution better than 3.1 mm FWHM; the required sampling period is 1.55 mm or less. This sampling problem can be overcome by moving the cylinder during the data acquisition to improve the sampling density over the tomograph field ("clamshell", dichotomic and "wobble" motions are current). The commercial PET
cameras based on the principle of position decoding by means of discrete crystals all utilize a "wobble"
motion of the detector cylinder to satisfy the sampling theorem of Nyquist.
(c) DISCRETE CRYSTALS WITH INDEPENDENT READOUT:
In this approach, reading is carried out through individually coupled detectors to eliminate any error caused upon position decoding. The FWHM of the CAF is then very close to half the thickness of the scintillation crystals w/2. Obviously, sampling in stationary mode cannot satisfy the Nyquist theorem.
Although most of the PET tomographs of the first generation have been based on individually coupled discrete detectors, no high resolution tomograph using this approach has been commercialized yet. In the existing prototypes, a "clamshellN opening and closing motion of the detector cylinder i6 required to obtain adequate sampling.
In this context, one current development in positron emission tomography is oriented toward the design of cameras operating in stationary mode and having an image resolution close to that of the CAF.
The operation in stationary mode is advantageous in fast, dynamic studies, but is almost mandatory for heart studies where the synchronism with the movements of the heart is problematic.
Although PET cameras using analog decoding like the above mentioned PENN-PET already enable data acquisition in stationary mode, the resolution and sensitivity of these cameras are well lower than those using BGO scintillators. The research is therefore oriented toward the elaboration of discrete crystal cameras capable of reducing the sampling period in stationary mode.
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A first solution that has been proposed to operate in stationary mode is to reduce the dimensions of the detectors until the sampling theorem of Nyquist, corresponding to the theoretical S resolution limit of positron emission tomography, is satisfied. In practice, the thickness of the scintillation crystals must be reduced to 2 mm or less. The technical difficulties encountered in implementing this solution have up to now prevented the realization of tomographs using this solution.
A second solution consists of using a PET
camera formed of a plurality of concentric rings of detectors. The scintillation crystals are usually lS offset in the circumferential direction from one ring to the other. With this solution, the uniformity of the resolution through the UFOV of the tomograph is improved by reducing the out-of-center error of parallax. This is achieved by determining the radial coordinate of the point of interaction into the camera. The sampling in the UFOV of the tomograph is also improved in stationary mode. This solution however presents the dràwback that, in an arrangement comprising many concentric rings of detectors, the sensitivity associated with each trajectory varies;
indeed, the sensitivity of the detectors in the outer rings is reduced due to the attenuation of the radiation in the inner rings. Although sampling is improved, the statistical weight is not equally distributed. Proposals have been made to correct for this non-uniform statistics by making the depth of each concentric detector ring inversely proportional to the probability of interaction, but this solution results in a non-uniform resolution of the CAF.
A further solution consists of accurately measuring by means of photodiodes or other types of photodetectors the deepness of interaction within scintillation crystals tilted relative to the radial direction in order to obtain an essentially continuous sampling of the field in stationary mode. In practice, the accuracy of measurement that can be reached enables only a gross estimate of the position of interaction, and the improvement in sampling is not significantly different from that obtained with a PET
camera using two concentric detector rings.
OBJECTS OF THE INVENTION
A general object of the present invention is therefore to improve the performance, in stationary mode, of a PET tomograph.
Another object of the invention is to reduce the sampling period and increase the number of monitored trajectories in a PET tomograph operating in stationary mode.
A further object of the invention is to provide a PET tomograph capable of acquiring data satisfying the sampling theorem of Nyquist in stationary mode.
SUMMARY OF THE INVENTION
More specifically, in accordance with the present invention, there is provided a positron emission tomography apparatus comprising at least first and second laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of the first ring being offset in the circumferential direction with respect to the qamma-ray detectors of the second ring. In the PE~ apparatus, (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring. Accordingly, when radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each pair of gamma rays propagating in opposite directions along one of the first, second and third linear trajectories are detectable through the pair of detectors interconnected by this trajectory. A data acquisition circuit is connected to the gamma-ray detectors of the first and second rings to detect the pairs of gamma rays propagating in opposite directions along the first, second and third trajectories, and to acquire data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated. Finally, a computinq means combines in a single plane of reconstruction the acquired data indicative of the linear trajectories of the first and second direct planes and of the two cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in this plane of reconstruction a distribution of sources from which the detected pairs of gamma rays originate.
By offsetting the respective gamma-ray detectors of the first and second rings, and by combining in the plane of reconstruction the acquired data indicative of linear trajectories in the first and second direct planes and in the two cross planes, the number of distinct trajectories in the plane of reconstruction is significantly increased while the period of sampling is significantly reduced. This obviously facilitates satisfying of the sampling theorem of Nyquist.
The first and second rings of gamma-ray detectors define a ring assembly having a geometrical axis, and these gamma-ray detectors may be disposed angularly with respect to this geometrical axis and arranged in two staggered rows respectively defining the first and second rings.
In accordance with another preferred embodiment of the invention, each gamma-ray detector comprises a scintillation crystal to emit a flash of light upon detection of a gamma ray and a photodetector for converting this flash of light into an electric pulse, the photodetector comprises a silicon avalanche photodiode, and the scintillation crystal comprises bismuth germanium oxide.
According to a further preferred embodiment of the invention, the positron emission tomography apparatus comprises a number N of detector rings laterally adjacent to each other, N being an integer equal to or greater than 2, each detector has a dimension d in the circumferential direction, and the detectors of one ring are offset in the circumferential direction with respect to the detectors of the laterally adjacent ring or rings by a distance d/N. The combination in the single plane of reconstruction of data acquired from the N rings of gamma-ray detectors multiplies by N2 the number of distinct trajectories in the plane of reconstruction.
According to a further aspect of the invention, there is provided a positron emission tomography apparatus comprising at least three laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the ; gamma-ray detectors of each ring being offset in the circumferential direction with respect to the gamma-; ~ ray detectors of the adjacent ring or rings. The rings comprise at least two pairs of laterally adjacent first and second rings in which (a) first linear trajectories interconnect in a first direct ; 30 plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories ,, . . . ~ .
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interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring.
When radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each pair of gamma rays propagating in opposite directions along one of the linear trajectories is detectable through the pair of detectors interconnected by this trajectory. A data acquisition circuit connected to the gamma-ray detectors detects the pairs of gamma rays propagating in opposite directions along the linear trajectories, and acquires data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated. Then, a computing means combines, for each pair of rings, in a single plane of reconstruction the acquired data indicative of the linear trajectories of the first and second direct planes and of the two cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in this plane of reconstruction a distribution of sources from which the detected pairs of gamma rays originate.
According to an advantageous embodiment, the positron emission tomography apparatus comprises a number N of detector rings laterally adjacent to each other, N being an integer greater than 2, each detector has a dimension d in the circumferential direction, the detectors of one ring are offset in the circumferential direction with respect to the - .. .. . ... ..
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detectors of the laterally adjacent ring or rings by a distance d/2, and the combination in the planes of reconstruction of data acquired from the detectors multiplies by 4 the number of distinct trajectories and reduces the distance between these trajectories to d~4 in these planes of reconstruction.
In accordance with a further aspect of the invention, there is provided a positron emission tomography apparatus comprising a number N of laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of one ring being offset in the circumferential direction with respect to the gamma-ray detectors of the adjacent ring or rings. Therings comprise pairs of laterally adjacent first and second rings in which (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring. Radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electrcn to produce a pair of gamma rays propagating in opposite directions. Obviously, each pair of gamma rays propagating in opposite directions along one of the linear trajectories is detectable through the pair of detectors interconnected by this trajectory. A data acquisition circuit connected to the gamma-ray , detectors of the rings detects the pairs of gamma rays propagating in opposite directions along the trajectories, and acquires data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated. A computing means combines in a plurality of planes of reconstruction, each corresponding to a slice of the subject, the acquired data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in each plane of reconstruction a distribution of sources located in the corresponding slice and from which the detected pairs of gamma rays originate.
The objects, advantages and other features of the present invention will become more apparent upon reading of the following non restrictive description of preferred embodiments thereof, given by way of example only with reference to the accompanying drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
In the appended drawings:
Figure 1 is a perspective view of the camera of a PET apparatus in accordance with the present invention, including detector rings forming a cylinder around the subject under study;
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Figure 2 shows a first arrangement of gamma-ray detectors in the PET camera of Figure 1;
Figures 3a, 3b, 3c and 3d illustrate the trajectories in the first and second direct planes and in the two cross planes obtained with the arrangement of Figure 2;
Figure 4 is a second arrangement of gamma-ray detectors in the PET camera of Figure 1:
Figures 5a, 5b, 6a and 6b illustrate other arrangements for the gamma-ray detectors of the PET
camera of Figure l;
: 15 Figure 7 is a block diagram of an electronic circuit used in a PET apparatus according : to the subject invention; and ~: 20 Figure 8 is a representation of the above discussed "coincidence aperture function" (CAF) for a PET camera having a single ring of detectors, ~ being the distance between the centers of two detectors of the ring, d/2 the sampling period in the UFOV, w the ; 25 thickness of the scintillation crystals, and w/2 the maximum resolution in UFOV (FWHM of the CAF
;~ distribution).
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DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
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The PET apparatus in accordance with the present invention operates in stationary mode. As illustrated in Figure 1 of the appended drawings, it includes a PET camera 1 comprising a cylinder 2 formed of an assembly of laterally adjacent rings of gamma-ray detectors. As can be seen, the cylinder 2 surrounds the subject 4 under study.
The camera 1 further comprises two annular outer side walls 5 and 7, made of lead for shielding against scattered radiation and radiation from sources of radioactivity located outside of the field of view of the camera. The camera may further comprise (at 6 between the outer walls 5 and 7) one or several cylindrical inner walls made of lead or of a highly absorbing material for gamma-rays, for shielding against scattered radiation. A set of adjacent outer and/or inner walls defines an inner gap such as 3 through which the detected pairs of gamma rays (such as 26 and 27) propagate.
The cylinder or ring assembly 2 (Figure 1) comprises two adjacent rings 8 and 9 of gamma-ray detectors (Figures 1 and 2). The rings 8 and 9 have the same diameter, and are coaxial and laterally adjacent. As illustrated in Figure 2, the detectors of ring 8 are offset in the circumferential direction with respect to the detectors 11 of ring 9 by a distance d/2 corresponding to half the distance d separating the centers of two consecutive gamma-ray detectors in the ring 8 or 9.
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In Figure 1, the rings 8 and 9 are circular. However, they can present other forms; for example they can be hexagonal, octagonal, etc.
In operation, radioactive isotopes emitting positrons are injected into the zone of interest of the subject's body, in a manner well known in the art. Each positron annihilates with an electron to produce a pair of gamma rays of 511 keV
propagating in opposite directions, 180 apart from each other. When a pair of gamma rays propagate in opposite directions along (a) linear trajectories interconnecting in a first direct plane pairs of detectors 10 of the ring 8, (b) linear trajectories interconnecting in a second direct plane pairs of detectors 11 of the ring 9, and (c) linear trajectories interconnecting in two cross planes pairs of detectors each formed of one detector 10 and one detector 11, it is detected by the corresponding pair of gamma-ray detectors.
Figure 3a shows the monitored linear trajectories between the pairs of detectors 10 in the first direct plane of coincidence. As can be seen in this particular example, each detector 10 is interconnected through linear trajectories with seven diametrically opposite detectors 10'. Obviously, each detector 10 can be interconnected through linear trajectories with fewer or more diametrically opposite detectors 10'.
Figure 3b illustrates the monitored linear trajectories interconnecting pairs of detectors 11 in .
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the second direct plane of coincidence. Again, linear trajectories interconnect each detector 11 with seven diametrically opposite detectors 11'. Obviously, each detector 11 can be interconnected through linear trajectories with fewer or more diametrically opposite detectors 11'.
Figure 3c illustrates the monitored trajectories in the two cross planes of coincidence.
In these two cross planes, linear trajectories interconnect each detector 10 with eight diametrically opposite detectors 11 ", and each detector 11 with eight diametrically opposite detectors 10''. Again, the number eight is only a non limitative example.
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The superimposition of the different trajectories of Figures 3a, 3b and 3c in a single plane of reconstruction is shown in Figure 3d. The body of the subject 4 is lying in the UFOV 15 of the PET apparatus delimited by the circle 13 (Figure 3d), where the density of trajectories is substantially higher and where a better sampling of the UFOV can therefore be obtained.
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Each trajectory represents in fact the volume situated between the two opposite detectors it interconnects. Each time two gamma rays are detected by a pair of opposite detectors within a predetermined short time window, it can be assumed that the source is located in the volume situated between these two detectors, which volume is represented by the linear trajectory interconnecting the pair of detectors.
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Each detector 10,11 advantageously comprises a discrete scintillation crystal or scintillator emitting a flash of light in response to a gamma radiation. In each detector, the scintillator is individually coupled to a photodetector capable of sensing the flash of light emitted by the scintillator and to convert it into an electric pulse.
Presently, the only detector of this type that can be utilized in the present invention is the one described in the article by R. Lecomte, J.
Cadorette, A. Jouan, M. Héon, D. Rouleau and G.
Gauthier, entitled "High resolution positron emission tomography with a prototype camera based on solid state scintillation detectors" published in IEEE
Transactions on Nuclear Science, Vol. 37, No. 2, April 1990, pages 805-811. The gamma-ray detector described in this article is made of the dual detection module RCA C30994 commercialized by EG & G Optoelectronics.
When the cylinder 2 consists of many rings of such RCA
C30994 detectors, an adjustment of the axial fraction of the package or of the configuration of the crystals would be necessary in order to maintain the periodicity in the axial direction. Also, an alternating configuration as illustrated in Figure 6b would solve the problems of periodicity but would result in a small loss of axial resolution. In the RCA C30994, the two scintillation crystals are made of BGO and the two photodetectors respectively coupled to these crystals are silicon avalanche photodiodes.
Although discrete and individuallycoupled detectors as described above are well adapted for use 207773~
in the present invention, the configurations proposed in Figures 2, 4, 5a, Sb, 6a and 6b can be easily adapted to the position decoding on a matrix of discrete crystals described in the preamble of the present specification. A system, commercialized under the name of Posicam and described in United States patent N 4,563,582 (Mullani) issued on January 7, 1986, United States patent N 4,642,464 granted to Nizar A. Nullani on February 10, 1987, and United States patent N4,864,138 (Mullani) dated September 5, 1989, already uses this approach with scintillators arranged in staggered rows in the axial direction.
Similar configurations can therefore be oriented in the circumferential direction of the rings. With this approach, the extension to more than two detector rings can be envisaged without any apparent problem.
Although the techniques used in PET to acquire data and to reconstruct these data are well known to those skilled in the art, a brief description of the procedure followed to perform these operations will be given in the following description in the context of the present invention.
Referring to Figure 7, the gamma ray is first absorbed by the scintillation crystal 40 of the detector 10,11. In response to the gamma ray, the scintillation crystal 40 produces a flash of light converted into an electric pulse by the photodetector 41.
The photodetector 41 of each detector 10,11 is associated to a circuit 30 that shapes any electric pulse from this photodetector. As illustrated in Figure 7, the shaping circuit 30 is included in a data acquisition circuit 31. A shaping circuit 30 is associated to each detector 10,11 as illustrated by the dashed line shaping circuit 30'.
The data acquisition circuit 31 further comprises a discrimination circuit 32 producing an output pulse in response to an electric pulse from the shaping circuit 30 that exceeds a predetermined amplitude.
A coincidence circuit 33 receives the output pulses from all the discrimination circuits 32 to coincide two detections of gamma rays by two diametrically opposite detectors when the two gamma-ray detections occur within a predetermined short time window. This information is then interpreted as a detection of a pair of gamma rays having propagated in opposite direction along the linear trajectory interconnecting the two detectors 10,11 of concern.
A circuit 34 then encodes the detection into an address transmitted to a trajectory memory circuit 35.
Each time a pair of gamma rays detected by the two detectors interconnected by a given trajectory are coincided by the circuit 33, this detection is encoded into the address of this trajectory by the circuit 34, and a number stored in the corresponding address of the trajectory memory circuit 35 is incremented by 1.
Therefore, the trajectory memory circuit 35 stores an histogram of the gamma-ray detections of each trajectory.
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Of course, the time required for acquiring a sufficient quantity of data depends on the intensity of activity in the region of interest of the subject's body, that is on the number of coincided detections of pairs of gamma rays by time unit.
The data acquired by the acquisition circ~it 31, more specifically the data stored in the trajectory memory circuit 35 are available for processing by a computer 36. For that purpose, the computer 36 includes a projection rebinning program 37, a projection filtering program 38, and a backprojection program 39.
Program 37 enables the computer 36 to perform rebinning of the acquired trajectory data from the memory circuit 35 into l-D projection vectors measured at different angles around the subject 4.
Rebinning of the data is a trivial operation adapted ~20 to each system, it depends mainly on the electronic ;~coding used upon storing the acquired data in the memory circuit 35. The projections collected at -~different angles can be seen as a set of equations from which the unknown 2-D activity distribution is solved for.
Program 38 allows the computer 36 to carry out backprojection. Backprojection is a mathematical process where the l-D projection are re-projected onto a 2-D image matrix. The resulting image will resemble the original distribution, but contains an artefact where the image, instead of being concentrated at the center of the 2-D matrix, is spread over the entire ' " ' , ' ' " ,' ', ' ', 207773~
image matrix. This artefact is a direct consequence of the backprojection and i6 sometimes referred to as l/r-smearing since its intensity decreases proportionally to l/r from the centre of the image.
This l/r-smearing can be removed by appropriately filtering the projections before backprojection. This filtering process can be performed in the spatial domain by the convolution of the l-D projections using a filter corresponding to the inverse function of the contribution l/r, or in the Fourier domain by multiplying the Fourier transform of the l-D projection with a ramp of function f(~
which is the transfer function corresponding to the Fourier transform of the l/r inverse filter.
As this process of reconstruction of the image is current in PET, it will not be further described in the present specification.
A distribution of sources from which the detected and coincided pairs of gamma rays originate in the plane of reconstruction is thereby obtained and displayed by a display unit 42. The computer 36 indeed statistically determines the regions of activity, where the sources of the coincided pairs of gamma rays are located.
The present invention therefore combines the data acquired through the detectors of the rings 8 and 9 and related to different direct and cross planes into a single plane of reconstruction to reconstruct an image of the distribution of activity .
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in a slice 14 (see Figure 1) of the subject's body.
The rings ~ and 9 are accordingly fused into this single plane of reconstruction.
In comparison with a PET camera including a single ring of detectors or with a classical PET
camera including at least two rings of detectors in which the detectors of the first ring are not offset in the circumferential direction with respect to the detectors of the second ring, one skilled in the art can appreciate that the present invention divides by 2 the angular distance between projections in the single plane of reconstruction and divides by 2 the linear distance between samples on these proiections.
As a result, the number of trajectories scanning the UFOV 15 is multiplied by 4 and the average sampling period is divided by 2 with the two stationary rings 8 and 9 of the present invention when compared to the number of trajectories obtained with only one of these stationary rings. More generally, the combination into a single plane of reconstruction of a number N of adjacent rings, of which the respective detectors are offset by a distance d/N, divides by N the sampling period and multiplies by N2 the number of trajectories ~ 25 scanning the UFOV 15.
;~ Since the intrinsic resolution of a detecting system formed of discrete scintillation crystals is typically 25 to 75% higher than half the thickness of these crystals, the improvement in sampling resulting from the combination of two detector rings in a single plane of reconstruction is : ,: .
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,. :
,~' 207773~
generally sufficient to satisfy the sampling theorem of Nyquist.
Similar results can also be obtained with the arrangement of Figure 4. In this Figure, the gamma-ray detectors define and angle ~ with respect to a perpendicular 16 to the plane 12 of the ring assembly. An arrangement in staggered rows is thereby obtained. The distance d', parallel to the plane 12, between the centers of a pair of adjacent detectors 10 and 11 is however slightly greater than d/2, the distance d separating two consecutive detectors in the ring 8 or 9 in Figure 2. The sampling period is therefore slightly higher than that of the arrangement of Figure 2. Due to the orientation of the detectors in the arrangement of Figure 4, the spatial resolution is also slightly degraded. However, the Nyquist theorem is very likely to be satisfied in stationary mode in most of the situations when combining the two rings of detectors.
Any other arrangement suitable to offset the adjacent detectors 10 and 11 of the rings 8 and 9 is likely to lead to similar results. Examples are shown in Figures 5a and 5b, as well as in Figure 6a and 6b. In Figure 5b, the crystals are square and oriented to form an arrangement in staggered rows both in the plane 12 and in the axial direction 17. With the arrangement of Figure 5b, axial sampling (direction 17) can be improved by coinciding into a same plane of reconstruction the trajectories between diametrically opposite pairs of detectors forming part of axial detector rows. Figures 6a and 6b illustrate ... . ..
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the use of the dual detection module RCA C30994 to construct a single ring of modules (Figure 6a) and many rings of modules (Figure 6b).
In all the arrangements, the concept of the present invention is based on the approximation that the 2 (or N) adjacent rings scan a common volume, that is the 2 (or N) adjacent rings form a single ring having a number of detectors multiplied by 2 (or N).
The PET apparatus in accordance with the invention presents in particular the following advantages.
- As stated in the foregoing description, when N = 2, the angular distance between projections is divided by 2 in the single plane of reconstruction, the linear distance between samples on these projections is also divided by 2, the number of trajectories scanning the UFOV 15 is multiplied by 4, and the average sampling period is divided by 2. With a single ring of detectors or with a classical PET
camera including at least two rings of detectors in which the detectors of the first ring are not offset ; ~ 25 in the circumferential direction with respect to the detectors of the second ring, sampling and data acquisition in at least four different positions of the ring is required to obtain the same results.
- The concept of the subject invention enables a data acquisition satisfying the sampling theorem of Nyquist in stationary mode by means of an array of discrete detectors.
28 2û77735 - The data acquired with a PET apparatus in accordance with the present invention can still be used to reconstruct classical images corresponding to the individual planes (two direct and one cross planes). Alternatively, by combining the data from the two direct planes and the two cross planes, an image can be reconstructed into a single plane to thereby improve data sampling. The reconstruction involving combination of data from the four planes tends to reduce the axial resolution. This reduction of axial resolution can be avoided by reconstructing the image of each individual plane according to the classical technique of reconstruction. Accordingly, with the present invention, the method of reconstruction can be chosen to optimize the resolution in the plane of reconstruction or the axial resolution. Concerning the reconstruction in the cross planes (Figure 3c), it should be pointed out that the number of distinct trajectories is multiplied by 2 if compared with the number of trajectories obtained with a single ring arrangement or with a classical PET camera including at least two rings of detectors in which the detectors of the first ring are not offset in the circumferential direction with respect to the detectors of the second ring (see for example Fiqure 3a).
- The concept of the subject invention enables improving the statistical accuracy of the data in much the same way as the combination of the two rings of a classical tomograph with no circumferential offset improves the statistical accuracy of the data in the single reconstructed plane, with the additional advantage of improving the sampling period to satisfy the Nyquist condition, as described above.
- If the ratio d/L (L being the width of the detectors as shown for example in Figure 4) tends toward "1", the sampling period and the uniformity in the sampling are simultaneously improved. A three-dimensional image of better quality can then be obtained. The principles of the combination of the axial planes as proposed in United States patent N
4,642,464 (Mullani) issued on February 10, 1987, United States patent 4,563,582 (Mullani) granted on January 7, 1986 and United States patent 4,864,138 (Nullani) dated September 5, 1989, can then be applied.
However, the PET apparatus in accordance with the present invention presents the following drawbacks.
- The combination of the two adjacent rings of detectors conducts to an increase in the axial resolution of the order of 100%. Such a loss of resolution is generally unacceptable with most of the tomographs currentlyjavailable on the market where the axial dimension of the scintillation crystals is at least twice as much as their width, that is ~ ~ 2d.
With the new generation of cameras in which the scintillators dimensions will tend to be the same in the plane and in the axial direction, that is d ~ L, the degradation of axial resolution will constitute an acceptable compromise to improve the sampling and resolution in the plane of reconstruction.
- The combination of the two adjacent rings of detectors results in an increase in the axial slice thickness of the order of 100~. Such an increase in the axial dimension of the unit cells of the reconstructed image will result in substantially more important partial volume effects which will affect the accuracy of the measurement of activity and possibly the contrast in the reconstructed image.
- In the case where the detectors are angularly rotated with respect to the circumference of the ring as shown in Figure 4, the effective width of the CAF will be slightly degraded. This loss of intrinsic resolution is however low and is unlikely to affect substantially the resolution in the reconstructed image.
It should finally be pointed out that, in a system comprising a plurality of laterally adjacent rings in which the gamma-ray detectors of each ring are offset in the circumferential direction with respect to the detectors of the one or two adjacent ring(s) by a distance d/2 (the detectors of the even and odd rings being aligned), each pair of consecutive rings can be used to reconstruct an image ir an associated plane of reconstruction. With this arrangement, the distance between the trajectories is reduced to d/~ and the number of distinct trajectories is multiplied by 4 in each plane of reconstruction.
The drawback is that the axial resolution is degraded by a factor ~ 2.
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Furthermore, with an assembly of many rings in which the detectors of the adjacent rings are offset by a distance d/2, axial sampling can still be carried out using different overlapping combinations of planes of coincidence. By this method, the distance between the adjacent planes of reconstruction can be maintained. This however causes an important overlapping between the planes of reconstruction and therefore a redundancy of the data used to reconstruct the adjacent planes.
Although the present invention has been described hereinabove by way of preferred embodiments thereof, such embodiments can be modified at will, within the scope of the appended claims, without departing from the spirit and nature of the subject invention.
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OVER-SAMPLING PET TOMOGRAPH
OPERATING IN STATIONARY MODE
FIELD OF THE INVENTION
The present invention relates to the field of positron emission tomography (PET). The subject invention is more specifically concerned with sampling requirements in positron emission tomography.
BACKGROUND OF THE INVENTION
Positron emission tomography (PET) uses radioactive isotopes decaying by emitting positrons (~t decay) as tracers for monitoring the spatial and temporal distribution of radioactivity within a body.
The method is useful to obtain information on dynamic processes such as the flow of a fluid within a pipe or on physiological processes within a living (animal or human) body. After emission by the radioactive nuclei, each positron annihilates with an electron to produce two gamma rays of 511 kev propagating in respective opposite directions 180 apart from each other. Diametrically opposite detectors can detect in coincidence these pairs of gamma rays, to thereby determine the trajectory on which the annihilation occurred. By superposing, by means of the 207773~
conventional techniques of tomography reconstruction, the plurality of trajectories detected in a plane by an array of detectors surrounding the subject under study, a distribution of annihilations (or sources of emission of pairs of gamma rays propagating in opposite directions) in the subject's body can be derived. A three-dimensional image can also be obtained by juxtaposing the two-dimensional images of the distribution of annihilations in many adjacent planes. The principles and advantages of PET in performing quantitative measurements of the biochemistry and the physiology "in vivo" of organs are well known in the art, have been extensively described in the literature and, therefore, do not need to be further discussed in the present specification.
In PET, the theoretical resolution limit that can be reached is of the order of 2-3 mm FWHM
(Full Width at Half Maximum of the distribution).
Those skilled in the art know that this limit originates from the ~' emission process and the annihilation positron-electron conducting to the production of the two gamma rays of 511 kev.
Presently, most PET tomographs have resolutions inferior (that is larger) to the theoretical limit.
The current trend is to design PET tomographs achieving resolutions equal or superior (that is smaller) to the theoretical limit and therefore capable of obtaining a maximum of spatial information.
Only a few prototypes, that reach a resolution close to the theoretical limit, have been built up to now, and such prototypes are not yet available on the 20~73~
market due to the technological and economical constraints which impose many compromises.
A typical PET camera comprises at least one ring of gamma-ray detectors forming a cylinder around the subject under study. With this configuration, any linear trajectory interconnecting a pair of diametrically opposite detectors and passing through the region of interest of the subject's body can be monitored. The measurements performed by a PET
camera therefore correspond to a series of linear trajectories interconnecting points, more specifically gamma-ray detectors of the cylinder. These trajectories can be grouped into parallel series, called parallel projections, usable in this form to reconstruct the images. The capacity of a PE~ camera to locate a gamma-ray source in one such projection defines a function of probability of presence, referred to in the art as "coincidence aperture function" (CAF), also called "line spread function".
The so called "intrinsic resolution" of a PET camera corresponds to the FWHM of this function. In practice, the intrinsic resolution is limited by the accuracy of the detectors in determining the point of impact of the gamma rays on the cylinder.
The resolution obtained in the reconstructed image or "image resolution" further depends on the sampling frequency (number of samples by unit linear distance) in the projections.
According to the sampling theorem of Nyquist, the best resolution of a detector arrangement cannot be smaller than two times the distance between two samples.
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Accordingly, the sampling period must be at least two times as small as the FWHM of the CAF to reach the maximum resolution.
The characteristics of resolution of the CAF's and of sampling frequency of given PET cameras depend on the type of the gamma-ray detectors and of the arrangement of these detectors on the cylinder.
Many approaches, differing by their costs and their performance, are currently used by the PET cameras to extract information about the position of the gamma-ray sources. One can note three categories:
(a) ANALOG POSITION DECODING:
With this approach, an estimate of the position of interaction of the gamma rays with the detectors is obtained from the signal of many photodetectors coupled to a large scintillation crystal (scintillator). In nuclear medicine, this is the method of detection currently used in the scintillation cameras. In PET, scintillators having high light conversion capabilities and a relatively short time constant are essential to ensure the statistical accuracy of the measurement of the position of interaction and allow the detection in coincidence. Presently, only NaI(Tl) adequately satisfies these stringent requirements. However, due to the low stopping power of NaI(Tl) for 511 kev gamma rays and due to the large probability of Compton scattering into the scintillation crystal, the achievable spatial resolution is limited at a value greater than 5 mm FWHM. Although this is far from the theoretical performance of PET, this approach has the advantage of achieving an essentially continuous sampling in the projections, limited only by the precision of the circuit digitizing the position of interaction of the gamma ray within the crystal, thereby automatically satisfying the Nyquist theorem.
Only one commercial PET camera uses that approach. It has been developed by UGM Medical Systems Inc. under the name PENN-PET and reaches a spatial resolution of 5-6 mm FWHM.
(b) POSITION DECODING ON A MATRIX OF CRYSTALS:
In order to solve the problems of limited resolution and of low detection efficiency inherent to the NaI(Tl) scintillators, most of the current PET
cameras use crystals made of bismuth germanium oxide (BGO). As BGO has a low light conversion efficiency, classical analog decoding is not possible. A
tentative solution consists of employing a matrix of discrete crystals instead of a unique crystal to enable classical analog decoding. This type of configuration tends to reduce spreading of the scintillation photons by channelling them toward a smaller number of photodetectors. As the statistics of the number of photons is low, the decoding is imperfect, therefore limiting the resolution. The best prototype detectors of this type achieve a resolution of the CAF between 3 to 4 mm FWHM.
One major drawback of the use of discrete crystals is the loss of continuous sampling.
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Theoretically, the best FWHM resolution of the CAF is of the order of half the thickness of the crystal, w/2. The sampling period obtained in the useful field of view (UFOV) of the tomograph by interleaving even and odd lines of coincidence is equal to half the distance between two detectors d/2. In practical detectors formed of discrete scintillation crystals, d ~ w and the theorem of Nyquist cannot be satisfied.
As the discrete crystals are made narrower in coded detection schemes, the statistical effects and the scattering of gamma rays between adjacent crystals contribute to substantially degrade the resolution of the CAF. However, even with the CAF resolution of current systems which can be substantially larger than ~/2, the Nyquist criterion cannot usually be satisfied. For instance, in one of the above mentioned prototypes of this type (W.M. Digby, M.
Dahlbom and E.J. Hoffman, "Detector, shielding and geometric design factors for a high-resolution PET
systemn, IEEE Trans. Nucl. Sci. NS-37 (1990) 664-670), ~/2 = 1.75 ~, d/2 = 2.0 ~ and the CAF cannot have a resolution better than 3.1 mm FWHM; the required sampling period is 1.55 mm or less. This sampling problem can be overcome by moving the cylinder during the data acquisition to improve the sampling density over the tomograph field ("clamshell", dichotomic and "wobble" motions are current). The commercial PET
cameras based on the principle of position decoding by means of discrete crystals all utilize a "wobble"
motion of the detector cylinder to satisfy the sampling theorem of Nyquist.
(c) DISCRETE CRYSTALS WITH INDEPENDENT READOUT:
In this approach, reading is carried out through individually coupled detectors to eliminate any error caused upon position decoding. The FWHM of the CAF is then very close to half the thickness of the scintillation crystals w/2. Obviously, sampling in stationary mode cannot satisfy the Nyquist theorem.
Although most of the PET tomographs of the first generation have been based on individually coupled discrete detectors, no high resolution tomograph using this approach has been commercialized yet. In the existing prototypes, a "clamshellN opening and closing motion of the detector cylinder i6 required to obtain adequate sampling.
In this context, one current development in positron emission tomography is oriented toward the design of cameras operating in stationary mode and having an image resolution close to that of the CAF.
The operation in stationary mode is advantageous in fast, dynamic studies, but is almost mandatory for heart studies where the synchronism with the movements of the heart is problematic.
Although PET cameras using analog decoding like the above mentioned PENN-PET already enable data acquisition in stationary mode, the resolution and sensitivity of these cameras are well lower than those using BGO scintillators. The research is therefore oriented toward the elaboration of discrete crystal cameras capable of reducing the sampling period in stationary mode.
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A first solution that has been proposed to operate in stationary mode is to reduce the dimensions of the detectors until the sampling theorem of Nyquist, corresponding to the theoretical S resolution limit of positron emission tomography, is satisfied. In practice, the thickness of the scintillation crystals must be reduced to 2 mm or less. The technical difficulties encountered in implementing this solution have up to now prevented the realization of tomographs using this solution.
A second solution consists of using a PET
camera formed of a plurality of concentric rings of detectors. The scintillation crystals are usually lS offset in the circumferential direction from one ring to the other. With this solution, the uniformity of the resolution through the UFOV of the tomograph is improved by reducing the out-of-center error of parallax. This is achieved by determining the radial coordinate of the point of interaction into the camera. The sampling in the UFOV of the tomograph is also improved in stationary mode. This solution however presents the dràwback that, in an arrangement comprising many concentric rings of detectors, the sensitivity associated with each trajectory varies;
indeed, the sensitivity of the detectors in the outer rings is reduced due to the attenuation of the radiation in the inner rings. Although sampling is improved, the statistical weight is not equally distributed. Proposals have been made to correct for this non-uniform statistics by making the depth of each concentric detector ring inversely proportional to the probability of interaction, but this solution results in a non-uniform resolution of the CAF.
A further solution consists of accurately measuring by means of photodiodes or other types of photodetectors the deepness of interaction within scintillation crystals tilted relative to the radial direction in order to obtain an essentially continuous sampling of the field in stationary mode. In practice, the accuracy of measurement that can be reached enables only a gross estimate of the position of interaction, and the improvement in sampling is not significantly different from that obtained with a PET
camera using two concentric detector rings.
OBJECTS OF THE INVENTION
A general object of the present invention is therefore to improve the performance, in stationary mode, of a PET tomograph.
Another object of the invention is to reduce the sampling period and increase the number of monitored trajectories in a PET tomograph operating in stationary mode.
A further object of the invention is to provide a PET tomograph capable of acquiring data satisfying the sampling theorem of Nyquist in stationary mode.
SUMMARY OF THE INVENTION
More specifically, in accordance with the present invention, there is provided a positron emission tomography apparatus comprising at least first and second laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of the first ring being offset in the circumferential direction with respect to the qamma-ray detectors of the second ring. In the PE~ apparatus, (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring. Accordingly, when radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each pair of gamma rays propagating in opposite directions along one of the first, second and third linear trajectories are detectable through the pair of detectors interconnected by this trajectory. A data acquisition circuit is connected to the gamma-ray detectors of the first and second rings to detect the pairs of gamma rays propagating in opposite directions along the first, second and third trajectories, and to acquire data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated. Finally, a computinq means combines in a single plane of reconstruction the acquired data indicative of the linear trajectories of the first and second direct planes and of the two cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in this plane of reconstruction a distribution of sources from which the detected pairs of gamma rays originate.
By offsetting the respective gamma-ray detectors of the first and second rings, and by combining in the plane of reconstruction the acquired data indicative of linear trajectories in the first and second direct planes and in the two cross planes, the number of distinct trajectories in the plane of reconstruction is significantly increased while the period of sampling is significantly reduced. This obviously facilitates satisfying of the sampling theorem of Nyquist.
The first and second rings of gamma-ray detectors define a ring assembly having a geometrical axis, and these gamma-ray detectors may be disposed angularly with respect to this geometrical axis and arranged in two staggered rows respectively defining the first and second rings.
In accordance with another preferred embodiment of the invention, each gamma-ray detector comprises a scintillation crystal to emit a flash of light upon detection of a gamma ray and a photodetector for converting this flash of light into an electric pulse, the photodetector comprises a silicon avalanche photodiode, and the scintillation crystal comprises bismuth germanium oxide.
According to a further preferred embodiment of the invention, the positron emission tomography apparatus comprises a number N of detector rings laterally adjacent to each other, N being an integer equal to or greater than 2, each detector has a dimension d in the circumferential direction, and the detectors of one ring are offset in the circumferential direction with respect to the detectors of the laterally adjacent ring or rings by a distance d/N. The combination in the single plane of reconstruction of data acquired from the N rings of gamma-ray detectors multiplies by N2 the number of distinct trajectories in the plane of reconstruction.
According to a further aspect of the invention, there is provided a positron emission tomography apparatus comprising at least three laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the ; gamma-ray detectors of each ring being offset in the circumferential direction with respect to the gamma-; ~ ray detectors of the adjacent ring or rings. The rings comprise at least two pairs of laterally adjacent first and second rings in which (a) first linear trajectories interconnect in a first direct ; 30 plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories ,, . . . ~ .
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interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring.
When radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each pair of gamma rays propagating in opposite directions along one of the linear trajectories is detectable through the pair of detectors interconnected by this trajectory. A data acquisition circuit connected to the gamma-ray detectors detects the pairs of gamma rays propagating in opposite directions along the linear trajectories, and acquires data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated. Then, a computing means combines, for each pair of rings, in a single plane of reconstruction the acquired data indicative of the linear trajectories of the first and second direct planes and of the two cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in this plane of reconstruction a distribution of sources from which the detected pairs of gamma rays originate.
According to an advantageous embodiment, the positron emission tomography apparatus comprises a number N of detector rings laterally adjacent to each other, N being an integer greater than 2, each detector has a dimension d in the circumferential direction, the detectors of one ring are offset in the circumferential direction with respect to the - .. .. . ... ..
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detectors of the laterally adjacent ring or rings by a distance d/2, and the combination in the planes of reconstruction of data acquired from the detectors multiplies by 4 the number of distinct trajectories and reduces the distance between these trajectories to d~4 in these planes of reconstruction.
In accordance with a further aspect of the invention, there is provided a positron emission tomography apparatus comprising a number N of laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of one ring being offset in the circumferential direction with respect to the gamma-ray detectors of the adjacent ring or rings. Therings comprise pairs of laterally adjacent first and second rings in which (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring. Radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electrcn to produce a pair of gamma rays propagating in opposite directions. Obviously, each pair of gamma rays propagating in opposite directions along one of the linear trajectories is detectable through the pair of detectors interconnected by this trajectory. A data acquisition circuit connected to the gamma-ray , detectors of the rings detects the pairs of gamma rays propagating in opposite directions along the trajectories, and acquires data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated. A computing means combines in a plurality of planes of reconstruction, each corresponding to a slice of the subject, the acquired data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in each plane of reconstruction a distribution of sources located in the corresponding slice and from which the detected pairs of gamma rays originate.
The objects, advantages and other features of the present invention will become more apparent upon reading of the following non restrictive description of preferred embodiments thereof, given by way of example only with reference to the accompanying drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
In the appended drawings:
Figure 1 is a perspective view of the camera of a PET apparatus in accordance with the present invention, including detector rings forming a cylinder around the subject under study;
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Figure 2 shows a first arrangement of gamma-ray detectors in the PET camera of Figure 1;
Figures 3a, 3b, 3c and 3d illustrate the trajectories in the first and second direct planes and in the two cross planes obtained with the arrangement of Figure 2;
Figure 4 is a second arrangement of gamma-ray detectors in the PET camera of Figure 1:
Figures 5a, 5b, 6a and 6b illustrate other arrangements for the gamma-ray detectors of the PET
camera of Figure l;
: 15 Figure 7 is a block diagram of an electronic circuit used in a PET apparatus according : to the subject invention; and ~: 20 Figure 8 is a representation of the above discussed "coincidence aperture function" (CAF) for a PET camera having a single ring of detectors, ~ being the distance between the centers of two detectors of the ring, d/2 the sampling period in the UFOV, w the ; 25 thickness of the scintillation crystals, and w/2 the maximum resolution in UFOV (FWHM of the CAF
;~ distribution).
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DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
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The PET apparatus in accordance with the present invention operates in stationary mode. As illustrated in Figure 1 of the appended drawings, it includes a PET camera 1 comprising a cylinder 2 formed of an assembly of laterally adjacent rings of gamma-ray detectors. As can be seen, the cylinder 2 surrounds the subject 4 under study.
The camera 1 further comprises two annular outer side walls 5 and 7, made of lead for shielding against scattered radiation and radiation from sources of radioactivity located outside of the field of view of the camera. The camera may further comprise (at 6 between the outer walls 5 and 7) one or several cylindrical inner walls made of lead or of a highly absorbing material for gamma-rays, for shielding against scattered radiation. A set of adjacent outer and/or inner walls defines an inner gap such as 3 through which the detected pairs of gamma rays (such as 26 and 27) propagate.
The cylinder or ring assembly 2 (Figure 1) comprises two adjacent rings 8 and 9 of gamma-ray detectors (Figures 1 and 2). The rings 8 and 9 have the same diameter, and are coaxial and laterally adjacent. As illustrated in Figure 2, the detectors of ring 8 are offset in the circumferential direction with respect to the detectors 11 of ring 9 by a distance d/2 corresponding to half the distance d separating the centers of two consecutive gamma-ray detectors in the ring 8 or 9.
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In Figure 1, the rings 8 and 9 are circular. However, they can present other forms; for example they can be hexagonal, octagonal, etc.
In operation, radioactive isotopes emitting positrons are injected into the zone of interest of the subject's body, in a manner well known in the art. Each positron annihilates with an electron to produce a pair of gamma rays of 511 keV
propagating in opposite directions, 180 apart from each other. When a pair of gamma rays propagate in opposite directions along (a) linear trajectories interconnecting in a first direct plane pairs of detectors 10 of the ring 8, (b) linear trajectories interconnecting in a second direct plane pairs of detectors 11 of the ring 9, and (c) linear trajectories interconnecting in two cross planes pairs of detectors each formed of one detector 10 and one detector 11, it is detected by the corresponding pair of gamma-ray detectors.
Figure 3a shows the monitored linear trajectories between the pairs of detectors 10 in the first direct plane of coincidence. As can be seen in this particular example, each detector 10 is interconnected through linear trajectories with seven diametrically opposite detectors 10'. Obviously, each detector 10 can be interconnected through linear trajectories with fewer or more diametrically opposite detectors 10'.
Figure 3b illustrates the monitored linear trajectories interconnecting pairs of detectors 11 in .
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the second direct plane of coincidence. Again, linear trajectories interconnect each detector 11 with seven diametrically opposite detectors 11'. Obviously, each detector 11 can be interconnected through linear trajectories with fewer or more diametrically opposite detectors 11'.
Figure 3c illustrates the monitored trajectories in the two cross planes of coincidence.
In these two cross planes, linear trajectories interconnect each detector 10 with eight diametrically opposite detectors 11 ", and each detector 11 with eight diametrically opposite detectors 10''. Again, the number eight is only a non limitative example.
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The superimposition of the different trajectories of Figures 3a, 3b and 3c in a single plane of reconstruction is shown in Figure 3d. The body of the subject 4 is lying in the UFOV 15 of the PET apparatus delimited by the circle 13 (Figure 3d), where the density of trajectories is substantially higher and where a better sampling of the UFOV can therefore be obtained.
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Each trajectory represents in fact the volume situated between the two opposite detectors it interconnects. Each time two gamma rays are detected by a pair of opposite detectors within a predetermined short time window, it can be assumed that the source is located in the volume situated between these two detectors, which volume is represented by the linear trajectory interconnecting the pair of detectors.
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Each detector 10,11 advantageously comprises a discrete scintillation crystal or scintillator emitting a flash of light in response to a gamma radiation. In each detector, the scintillator is individually coupled to a photodetector capable of sensing the flash of light emitted by the scintillator and to convert it into an electric pulse.
Presently, the only detector of this type that can be utilized in the present invention is the one described in the article by R. Lecomte, J.
Cadorette, A. Jouan, M. Héon, D. Rouleau and G.
Gauthier, entitled "High resolution positron emission tomography with a prototype camera based on solid state scintillation detectors" published in IEEE
Transactions on Nuclear Science, Vol. 37, No. 2, April 1990, pages 805-811. The gamma-ray detector described in this article is made of the dual detection module RCA C30994 commercialized by EG & G Optoelectronics.
When the cylinder 2 consists of many rings of such RCA
C30994 detectors, an adjustment of the axial fraction of the package or of the configuration of the crystals would be necessary in order to maintain the periodicity in the axial direction. Also, an alternating configuration as illustrated in Figure 6b would solve the problems of periodicity but would result in a small loss of axial resolution. In the RCA C30994, the two scintillation crystals are made of BGO and the two photodetectors respectively coupled to these crystals are silicon avalanche photodiodes.
Although discrete and individuallycoupled detectors as described above are well adapted for use 207773~
in the present invention, the configurations proposed in Figures 2, 4, 5a, Sb, 6a and 6b can be easily adapted to the position decoding on a matrix of discrete crystals described in the preamble of the present specification. A system, commercialized under the name of Posicam and described in United States patent N 4,563,582 (Mullani) issued on January 7, 1986, United States patent N 4,642,464 granted to Nizar A. Nullani on February 10, 1987, and United States patent N4,864,138 (Mullani) dated September 5, 1989, already uses this approach with scintillators arranged in staggered rows in the axial direction.
Similar configurations can therefore be oriented in the circumferential direction of the rings. With this approach, the extension to more than two detector rings can be envisaged without any apparent problem.
Although the techniques used in PET to acquire data and to reconstruct these data are well known to those skilled in the art, a brief description of the procedure followed to perform these operations will be given in the following description in the context of the present invention.
Referring to Figure 7, the gamma ray is first absorbed by the scintillation crystal 40 of the detector 10,11. In response to the gamma ray, the scintillation crystal 40 produces a flash of light converted into an electric pulse by the photodetector 41.
The photodetector 41 of each detector 10,11 is associated to a circuit 30 that shapes any electric pulse from this photodetector. As illustrated in Figure 7, the shaping circuit 30 is included in a data acquisition circuit 31. A shaping circuit 30 is associated to each detector 10,11 as illustrated by the dashed line shaping circuit 30'.
The data acquisition circuit 31 further comprises a discrimination circuit 32 producing an output pulse in response to an electric pulse from the shaping circuit 30 that exceeds a predetermined amplitude.
A coincidence circuit 33 receives the output pulses from all the discrimination circuits 32 to coincide two detections of gamma rays by two diametrically opposite detectors when the two gamma-ray detections occur within a predetermined short time window. This information is then interpreted as a detection of a pair of gamma rays having propagated in opposite direction along the linear trajectory interconnecting the two detectors 10,11 of concern.
A circuit 34 then encodes the detection into an address transmitted to a trajectory memory circuit 35.
Each time a pair of gamma rays detected by the two detectors interconnected by a given trajectory are coincided by the circuit 33, this detection is encoded into the address of this trajectory by the circuit 34, and a number stored in the corresponding address of the trajectory memory circuit 35 is incremented by 1.
Therefore, the trajectory memory circuit 35 stores an histogram of the gamma-ray detections of each trajectory.
, ~, . ~ . ... . .
Of course, the time required for acquiring a sufficient quantity of data depends on the intensity of activity in the region of interest of the subject's body, that is on the number of coincided detections of pairs of gamma rays by time unit.
The data acquired by the acquisition circ~it 31, more specifically the data stored in the trajectory memory circuit 35 are available for processing by a computer 36. For that purpose, the computer 36 includes a projection rebinning program 37, a projection filtering program 38, and a backprojection program 39.
Program 37 enables the computer 36 to perform rebinning of the acquired trajectory data from the memory circuit 35 into l-D projection vectors measured at different angles around the subject 4.
Rebinning of the data is a trivial operation adapted ~20 to each system, it depends mainly on the electronic ;~coding used upon storing the acquired data in the memory circuit 35. The projections collected at -~different angles can be seen as a set of equations from which the unknown 2-D activity distribution is solved for.
Program 38 allows the computer 36 to carry out backprojection. Backprojection is a mathematical process where the l-D projection are re-projected onto a 2-D image matrix. The resulting image will resemble the original distribution, but contains an artefact where the image, instead of being concentrated at the center of the 2-D matrix, is spread over the entire ' " ' , ' ' " ,' ', ' ', 207773~
image matrix. This artefact is a direct consequence of the backprojection and i6 sometimes referred to as l/r-smearing since its intensity decreases proportionally to l/r from the centre of the image.
This l/r-smearing can be removed by appropriately filtering the projections before backprojection. This filtering process can be performed in the spatial domain by the convolution of the l-D projections using a filter corresponding to the inverse function of the contribution l/r, or in the Fourier domain by multiplying the Fourier transform of the l-D projection with a ramp of function f(~
which is the transfer function corresponding to the Fourier transform of the l/r inverse filter.
As this process of reconstruction of the image is current in PET, it will not be further described in the present specification.
A distribution of sources from which the detected and coincided pairs of gamma rays originate in the plane of reconstruction is thereby obtained and displayed by a display unit 42. The computer 36 indeed statistically determines the regions of activity, where the sources of the coincided pairs of gamma rays are located.
The present invention therefore combines the data acquired through the detectors of the rings 8 and 9 and related to different direct and cross planes into a single plane of reconstruction to reconstruct an image of the distribution of activity .
207773~
in a slice 14 (see Figure 1) of the subject's body.
The rings ~ and 9 are accordingly fused into this single plane of reconstruction.
In comparison with a PET camera including a single ring of detectors or with a classical PET
camera including at least two rings of detectors in which the detectors of the first ring are not offset in the circumferential direction with respect to the detectors of the second ring, one skilled in the art can appreciate that the present invention divides by 2 the angular distance between projections in the single plane of reconstruction and divides by 2 the linear distance between samples on these proiections.
As a result, the number of trajectories scanning the UFOV 15 is multiplied by 4 and the average sampling period is divided by 2 with the two stationary rings 8 and 9 of the present invention when compared to the number of trajectories obtained with only one of these stationary rings. More generally, the combination into a single plane of reconstruction of a number N of adjacent rings, of which the respective detectors are offset by a distance d/N, divides by N the sampling period and multiplies by N2 the number of trajectories ~ 25 scanning the UFOV 15.
;~ Since the intrinsic resolution of a detecting system formed of discrete scintillation crystals is typically 25 to 75% higher than half the thickness of these crystals, the improvement in sampling resulting from the combination of two detector rings in a single plane of reconstruction is : ,: .
.. ' ' ' . . '~
,. :
,~' 207773~
generally sufficient to satisfy the sampling theorem of Nyquist.
Similar results can also be obtained with the arrangement of Figure 4. In this Figure, the gamma-ray detectors define and angle ~ with respect to a perpendicular 16 to the plane 12 of the ring assembly. An arrangement in staggered rows is thereby obtained. The distance d', parallel to the plane 12, between the centers of a pair of adjacent detectors 10 and 11 is however slightly greater than d/2, the distance d separating two consecutive detectors in the ring 8 or 9 in Figure 2. The sampling period is therefore slightly higher than that of the arrangement of Figure 2. Due to the orientation of the detectors in the arrangement of Figure 4, the spatial resolution is also slightly degraded. However, the Nyquist theorem is very likely to be satisfied in stationary mode in most of the situations when combining the two rings of detectors.
Any other arrangement suitable to offset the adjacent detectors 10 and 11 of the rings 8 and 9 is likely to lead to similar results. Examples are shown in Figures 5a and 5b, as well as in Figure 6a and 6b. In Figure 5b, the crystals are square and oriented to form an arrangement in staggered rows both in the plane 12 and in the axial direction 17. With the arrangement of Figure 5b, axial sampling (direction 17) can be improved by coinciding into a same plane of reconstruction the trajectories between diametrically opposite pairs of detectors forming part of axial detector rows. Figures 6a and 6b illustrate ... . ..
207773~
the use of the dual detection module RCA C30994 to construct a single ring of modules (Figure 6a) and many rings of modules (Figure 6b).
In all the arrangements, the concept of the present invention is based on the approximation that the 2 (or N) adjacent rings scan a common volume, that is the 2 (or N) adjacent rings form a single ring having a number of detectors multiplied by 2 (or N).
The PET apparatus in accordance with the invention presents in particular the following advantages.
- As stated in the foregoing description, when N = 2, the angular distance between projections is divided by 2 in the single plane of reconstruction, the linear distance between samples on these projections is also divided by 2, the number of trajectories scanning the UFOV 15 is multiplied by 4, and the average sampling period is divided by 2. With a single ring of detectors or with a classical PET
camera including at least two rings of detectors in which the detectors of the first ring are not offset ; ~ 25 in the circumferential direction with respect to the detectors of the second ring, sampling and data acquisition in at least four different positions of the ring is required to obtain the same results.
- The concept of the subject invention enables a data acquisition satisfying the sampling theorem of Nyquist in stationary mode by means of an array of discrete detectors.
28 2û77735 - The data acquired with a PET apparatus in accordance with the present invention can still be used to reconstruct classical images corresponding to the individual planes (two direct and one cross planes). Alternatively, by combining the data from the two direct planes and the two cross planes, an image can be reconstructed into a single plane to thereby improve data sampling. The reconstruction involving combination of data from the four planes tends to reduce the axial resolution. This reduction of axial resolution can be avoided by reconstructing the image of each individual plane according to the classical technique of reconstruction. Accordingly, with the present invention, the method of reconstruction can be chosen to optimize the resolution in the plane of reconstruction or the axial resolution. Concerning the reconstruction in the cross planes (Figure 3c), it should be pointed out that the number of distinct trajectories is multiplied by 2 if compared with the number of trajectories obtained with a single ring arrangement or with a classical PET camera including at least two rings of detectors in which the detectors of the first ring are not offset in the circumferential direction with respect to the detectors of the second ring (see for example Fiqure 3a).
- The concept of the subject invention enables improving the statistical accuracy of the data in much the same way as the combination of the two rings of a classical tomograph with no circumferential offset improves the statistical accuracy of the data in the single reconstructed plane, with the additional advantage of improving the sampling period to satisfy the Nyquist condition, as described above.
- If the ratio d/L (L being the width of the detectors as shown for example in Figure 4) tends toward "1", the sampling period and the uniformity in the sampling are simultaneously improved. A three-dimensional image of better quality can then be obtained. The principles of the combination of the axial planes as proposed in United States patent N
4,642,464 (Mullani) issued on February 10, 1987, United States patent 4,563,582 (Mullani) granted on January 7, 1986 and United States patent 4,864,138 (Nullani) dated September 5, 1989, can then be applied.
However, the PET apparatus in accordance with the present invention presents the following drawbacks.
- The combination of the two adjacent rings of detectors conducts to an increase in the axial resolution of the order of 100%. Such a loss of resolution is generally unacceptable with most of the tomographs currentlyjavailable on the market where the axial dimension of the scintillation crystals is at least twice as much as their width, that is ~ ~ 2d.
With the new generation of cameras in which the scintillators dimensions will tend to be the same in the plane and in the axial direction, that is d ~ L, the degradation of axial resolution will constitute an acceptable compromise to improve the sampling and resolution in the plane of reconstruction.
- The combination of the two adjacent rings of detectors results in an increase in the axial slice thickness of the order of 100~. Such an increase in the axial dimension of the unit cells of the reconstructed image will result in substantially more important partial volume effects which will affect the accuracy of the measurement of activity and possibly the contrast in the reconstructed image.
- In the case where the detectors are angularly rotated with respect to the circumference of the ring as shown in Figure 4, the effective width of the CAF will be slightly degraded. This loss of intrinsic resolution is however low and is unlikely to affect substantially the resolution in the reconstructed image.
It should finally be pointed out that, in a system comprising a plurality of laterally adjacent rings in which the gamma-ray detectors of each ring are offset in the circumferential direction with respect to the detectors of the one or two adjacent ring(s) by a distance d/2 (the detectors of the even and odd rings being aligned), each pair of consecutive rings can be used to reconstruct an image ir an associated plane of reconstruction. With this arrangement, the distance between the trajectories is reduced to d/~ and the number of distinct trajectories is multiplied by 4 in each plane of reconstruction.
The drawback is that the axial resolution is degraded by a factor ~ 2.
207773~
Furthermore, with an assembly of many rings in which the detectors of the adjacent rings are offset by a distance d/2, axial sampling can still be carried out using different overlapping combinations of planes of coincidence. By this method, the distance between the adjacent planes of reconstruction can be maintained. This however causes an important overlapping between the planes of reconstruction and therefore a redundancy of the data used to reconstruct the adjacent planes.
Although the present invention has been described hereinabove by way of preferred embodiments thereof, such embodiments can be modified at will, within the scope of the appended claims, without departing from the spirit and nature of the subject invention.
. : ,; ', ~' .
Claims (13)
1. A positron emission tomography apparatus comprising:
at least first and second laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of the first ring being offset in the circumferential direction with respect to the gamma-ray detectors of the second ring, wherein (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring;
whereby, when radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each said pair of gamma rays propagating in opposite directions along one of said first, second and third linear trajectories being detectable through the pair of detectors interconnected by said one trajectory;
a data acquisition circuit connected to the gamma-ray detectors of the first and second rings to detect the pairs of gamma rays propagating in opposite directions along the first, second and third trajectories, and to acquire data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated; and a computing means for combining in a single plane of reconstruction said acquired data indicative of the linear trajectories of the first and second direct planes and of the two cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in said plane of reconstruction a distribution of sources from which the detected pairs of gamma rays originate.
at least first and second laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of the first ring being offset in the circumferential direction with respect to the gamma-ray detectors of the second ring, wherein (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring;
whereby, when radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each said pair of gamma rays propagating in opposite directions along one of said first, second and third linear trajectories being detectable through the pair of detectors interconnected by said one trajectory;
a data acquisition circuit connected to the gamma-ray detectors of the first and second rings to detect the pairs of gamma rays propagating in opposite directions along the first, second and third trajectories, and to acquire data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated; and a computing means for combining in a single plane of reconstruction said acquired data indicative of the linear trajectories of the first and second direct planes and of the two cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in said plane of reconstruction a distribution of sources from which the detected pairs of gamma rays originate.
2. A positron emission tomography apparatus as recited in claim 1, comprising a number N of detector rings laterally adjacent to each other, N being an integer equal to or greater than 2, wherein each detector has a dimension d in the circumferential direction, and wherein the detectors of one ring are offset in the circumferential direction with respect to the detectors of the laterally adjacent ring or rings by a distance d/N.
3. A positron emission tomography apparatus as recited in claim 1, in which the first and second rings of gamma-ray detectors define a ring assembly having a geometrical axis, and in which said gamma-ray detectors are disposed angularly with respect to said geometrical axis and are arranged in two staggered rows respectively defining the first and second rings.
4. A positron emission tomography apparatus as recited in claim 1, wherein each gamma-ray detector comprises a scintillation crystal to emit a flash of light upon detection of a gamma ray, and a photodetector for converting said flash of light into an electric pulse, and wherein the photodetector comprises a silicon avalanche photodiode.
5. A positron emission tomography apparatus as recited in claim 1, wherein each gamma-ray detector comprises a scintillation crystal to emit a flash of light upon detection of a gamma ray, and a photodetector for converting said flash of light into and electric pulse, and wherein the scintillation crystal comprises bismuth germanium oxide.
6. A positron emission tomography apparatus as recited in claim 1, further comprising a display unit connected to said computing means for displaying an image of said distribution of gamma-ray sources in the single plane of reconstruction.
7. A positron emission tomography apparatus as recited in claim 2, wherein N = 2 and wherein said combination in the single plane of reconstruction of data acquired from the first and second rings having respective gamma-ray detectors offset by a distance d/2 in the circumferential direction multiplies by 4 the number of distinct trajectories in said plane of reconstruction.
8. A positron emission tomography apparatus as recited in claim 2, wherein said combination in the single plane of reconstruction of data acquired from said N rings of gamma-ray detectors multiplies by N2 the number of distinct trajectories in said plane of reconstruction.
9. A positron emission tomography apparatus as recited in claim 1, comprising a number of detector rings laterally adjacent to each other, N being an integer greater than 2, wherein each detector has a dimension d in the circumferential direction, and wherein the detectors of one ring are offset in the circumferential direction with respect to the detectors of the laterally adjacent ring or rings by a distance d/2.
10. A positron emission tomography apparatus comprising:
at least three laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of each ring being offset in the circumferential direction with respect to the gamma-ray detectors of the adjacent ring or rings, wherein said rings comprise at least two pairs of laterally adjacent first and second rings in which (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring: whereby, when radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each said pair of gamma rays propagating in opposite directions along one of said linear trajectories being detectable through the pair of detectors interconnected by said one trajectory;
a data acquisition circuit connected to said gamma-ray detectors to detect the pairs of gamma rays propagating in opposite directions along said linear trajectories, and to acquire data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated; and a computing means for combining, for each pair of rings, in a single plane of reconstruction said acquired data indicative of the linear trajectories of the first and second direct planes and of the two cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in said plane of reconstruction a distribution of sources from which the detected pairs of gamma rays originate.
at least three laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of each ring being offset in the circumferential direction with respect to the gamma-ray detectors of the adjacent ring or rings, wherein said rings comprise at least two pairs of laterally adjacent first and second rings in which (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring: whereby, when radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each said pair of gamma rays propagating in opposite directions along one of said linear trajectories being detectable through the pair of detectors interconnected by said one trajectory;
a data acquisition circuit connected to said gamma-ray detectors to detect the pairs of gamma rays propagating in opposite directions along said linear trajectories, and to acquire data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated; and a computing means for combining, for each pair of rings, in a single plane of reconstruction said acquired data indicative of the linear trajectories of the first and second direct planes and of the two cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in said plane of reconstruction a distribution of sources from which the detected pairs of gamma rays originate.
11. A positron emission tomography apparatus as recited in claim 10, comprising a number N of detector rings laterally adjacent to each other, N being an integer greater than 2, wherein each detector has a dimension d in the circumferential direction, and wherein the detectors of one ring are offset in the circumferential direction with respect to the detectors of the laterally adjacent ring or rings by a distance d/2.
12. A positron emission tomography apparatus as recited in claim 11, wherein said combination in the planes of reconstruction of data acquired from said detectors multiplies by 4 the number of distinct trajectories and reduces the distance between said trajectories to d/4 in said planes of reconstruction.
13. A positron emission tomography apparatus comprising:
a number N of laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of one ring being offset in the circumferential direction with respect to the gamma-ray detectors of the adjacent ring or rings, said rings comprising pairs of laterally adjacent first and second rings in which (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring:
whereby, when radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each said pair of gamma rays propagating in opposite directions along one of said linear trajectories being detectable through the pair of detectors interconnected by said one trajectory;
a data acquisition circuit connected to the gamma-ray detectors of the rings to detect the pairs of gamma rays propagating in opposite directions along said trajectories, and to acquire data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated; and a computing means for combining in a plurality of planes of reconstruction, each corresponding a slice of said subject, said acquired data indicative of the linear trajectories of said direct and cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in each plane of reconstruction a distribution of sources, located in the corresponding slice and from which the detected pairs of gamma rays originate.
a number N of laterally adjacent rings of gamma-ray detectors capable of surrounding a subject under study, the gamma-ray detectors of one ring being offset in the circumferential direction with respect to the gamma-ray detectors of the adjacent ring or rings, said rings comprising pairs of laterally adjacent first and second rings in which (a) first linear trajectories interconnect in a first direct plane pairs of gamma-ray detectors of the first ring, (b) second linear trajectories interconnect in a second direct plane pairs of gamma-ray detectors of the second ring, and (c) third linear trajectories interconnect in two cross planes pairs of detectors each formed of one gamma-ray detector of the first ring and one gamma-ray detector of the second ring:
whereby, when radioactive isotopes are injected into the subject under study and emit positrons each annihilating with an electron to produce a pair of gamma rays propagating in opposite directions, each said pair of gamma rays propagating in opposite directions along one of said linear trajectories being detectable through the pair of detectors interconnected by said one trajectory;
a data acquisition circuit connected to the gamma-ray detectors of the rings to detect the pairs of gamma rays propagating in opposite directions along said trajectories, and to acquire data indicative of the linear trajectories of the direct and cross planes along which the detected pairs of gamma rays have propagated; and a computing means for combining in a plurality of planes of reconstruction, each corresponding a slice of said subject, said acquired data indicative of the linear trajectories of said direct and cross planes along which the detected pairs of gamma rays have propagated, in order to reconstruct in each plane of reconstruction a distribution of sources, located in the corresponding slice and from which the detected pairs of gamma rays originate.
Priority Applications (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CA 2077735 CA2077735A1 (en) | 1992-09-08 | 1992-09-08 | Over-sampling pet tomograph operating in stationary mode |
Applications Claiming Priority (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CA 2077735 CA2077735A1 (en) | 1992-09-08 | 1992-09-08 | Over-sampling pet tomograph operating in stationary mode |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CA2077735A1 true CA2077735A1 (en) | 1994-03-09 |
Family
ID=4150381
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA 2077735 Abandoned CA2077735A1 (en) | 1992-09-08 | 1992-09-08 | Over-sampling pet tomograph operating in stationary mode |
Country Status (1)
| Country | Link |
|---|---|
| CA (1) | CA2077735A1 (en) |
Cited By (2)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| WO2014141040A3 (en) * | 2013-03-14 | 2014-11-27 | Koninklijke Philips N.V. | Positron emission tomography and/or single photon emission tomography detector and mri system |
| CN109009197A (en) * | 2018-08-06 | 2018-12-18 | 南京航空航天大学 | A kind of bicrystal item for PET detection passes through and meets line of response detection system and method |
-
1992
- 1992-09-08 CA CA 2077735 patent/CA2077735A1/en not_active Abandoned
Cited By (4)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| WO2014141040A3 (en) * | 2013-03-14 | 2014-11-27 | Koninklijke Philips N.V. | Positron emission tomography and/or single photon emission tomography detector and mri system |
| US9599731B2 (en) | 2013-03-14 | 2017-03-21 | Koninklijke Philips N.V. | Positron emission tomography and/or single photon emission tomography detector |
| CN109009197A (en) * | 2018-08-06 | 2018-12-18 | 南京航空航天大学 | A kind of bicrystal item for PET detection passes through and meets line of response detection system and method |
| CN109009197B (en) * | 2018-08-06 | 2020-07-07 | 南京航空航天大学 | Double-crystal strip crossing coincidence response line detection system and method for PET detection |
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