WO1997008569A1 - Appareil d'imagerie - Google Patents

Appareil d'imagerie Download PDF

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Publication number
WO1997008569A1
WO1997008569A1 PCT/GB1996/002114 GB9602114W WO9708569A1 WO 1997008569 A1 WO1997008569 A1 WO 1997008569A1 GB 9602114 W GB9602114 W GB 9602114W WO 9708569 A1 WO9708569 A1 WO 9708569A1
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image
events
gamma
ofthe
pet
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PCT/GB1996/002114
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English (en)
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Philip Palin Dendy
Edwin Philip Wraight
Steven David Bloomer
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British Technology Group Limited
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Priority to IL12345796A priority Critical patent/IL123457A0/xx
Priority to EP96928600A priority patent/EP0846275A1/fr
Priority to AU68320/96A priority patent/AU6832096A/en
Priority to JP9509999A priority patent/JPH11511551A/ja
Publication of WO1997008569A1 publication Critical patent/WO1997008569A1/fr

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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 or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/02—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/03—Computed tomography [CT]
    • A61B6/037—Emission tomography

Definitions

  • the present invention relates to an imaging apparatus, and more particularly to an imaging apparatus which enables a dual-headed gamma camera to be used for routine diagnostic imaging of positron-emitting radionuclides, as well as single-photon-emitting radionuclides.
  • a dual-headed gamma camera Such a device is hereinafter referred to as a modified dual-headed gamma camera.
  • a modified dual-headed gamma camera is used to image positron-emitting radiopharmaceuticals in routine clinical applications.
  • the present invention describes a modified dual-head gamma camera which is capable of retaining the ability to image single-photon-emitting radiopharmaceuticals.
  • the invention thus enables opposing gamma camera detectors to operate in coincidence mode, thereby obviating the need for collimators.
  • the subsequent improvement in sensitivity over a collimated system is exploited by use of fast digital processing techniques to handle efficiently the relatively high rate of single events, and by the application of a fully three-dimensional image reconstruction technique to ensure efficient use of acquired data.
  • the invention may be inco ⁇ orated as a switchable unit in future gamma camera system designs, or it may be retro-fitted as a modular add-on to existing designs.
  • Radionuclide imaging is a means by which the physiology of a patient can be studied and is therefore used as a tool in patient diagnosis and treatment management.
  • the basis ofthe technique is to introduce a radioactive substance into the patient after attaching it chemically to a physiologically important pharmaceutical, and to use gamma ray emissions from this radiopharmaceutical to construct images of its distribution at subsequent times under certain controlled conditions.
  • the clinical information obtained from the test depends on the biochemistry of the radiopharmaceutical, and the types of radiopharmaceutical available are often limited by the chemical properties of the radioactive element to be attached.
  • radionuclide imaging is to study gamma ray emissions from radionuclides that emit single gamma ray photons of energies in the range 50-400keV.
  • the tomographic study of such emitters is usually termed SPET or SPECT, though static and non-tomographic whole body scans are common.
  • the equipment used to image these emissions is the gamma camera, which is a position sensitive gamma ray detector.
  • a gamma camera detects gamma rays using a scintillation material (usually a single crystal of thallium-doped sodium iodide).
  • Gamma rays incident on the scintillation material convert some of their energy into optical light which is detected by a number of photosensitive detectors (usually photomultiplier tubes optically coupled to the scintillator by a lightguide).
  • the detector obtains the directional information of each gamma ray by means of an absorbing collimator, which limits the solid angle of acceptance to a narrow range around a particular set of directions.
  • Radionuclide imaging is positron emission tomography (PET).
  • PET positron emission tomography
  • the radioactive element emits a positron when it decays, which then collides with a nearby electron in the patient.
  • the positron and electron undergo mutual annihilation resulting in the emission of two 51 IkeV gamma rays in near-opposite directions.
  • the near- collinear property of these photons is useful in determining the location of the emission point in the patient and collimation is not necessary if both gamma rays are detected.
  • the energy of the gamma rays is so much higher than those encountered in SPET, gamma cameras are not well-suited to detecting them.
  • PET radionuclides are therefore usually imaged with a dedicated PET imager, which consists of several rings of position-sensitive scintillator/photomultiplier detectors operating in coincidence mode (i.e. without a collimator).
  • the detectors usually bismuth germanate, BGO, scintillators backed by position-sensitive photomultipliers
  • BGO bismuth germanate
  • Gamma ray events arriving at the detector rings at the same time are assumed to originate from the same annihilation event and the line between the two serves as the locus of possible emission sites within the patient. If many events are recorded, a tomographic image ofthe radiopharmaceutical distribution may be reconstructed.
  • SPET imaging with gamma-cameras is well established as a routine diagnostic tool and has a wide range of clinical applications in areas such as cardiology, oncology and neurology.
  • the single photon emitters tend to be heavy elements (e.g. technetium 99m Tc) and the range of pharmaceuticals to which these can be attached is limited.
  • Positron emitters tend to be lighter elements (e.g. fluorine l8 F) and are more readily attached to biologically important chemicals. For this reason, PET has been developed over the past two decades into a highly valuable research tool, providing physiological and medical imaging data not obtainable in any other way.
  • PET techniques can make a major contribution to diagnosis and patient management in routine medical practice (Schwaiger and Hutchins 1992, Broich et al. 1992 and Hawkins et al 1992). Whilst specialised centres derive great benefit from brain perfusion and functional imaging the most important areas for more general clinical application are in cardiology, demonstrating regional perfusion and myocardial viability more precisely than by any other technique, and in oncology where 18F-fluorodeoxyglucose (FDG) has been demonstrated to provide unique information on tumour staging and assessment of residual tumour viability following therapy. Despite this, PET radionuclides are not currently used as routine diagnostic imaging agents. The chief reason for this is cost.
  • a dedicated PET imager is currently approximately six times as expensive as a state-of the art, large field of view modified dual -headed gamma camera. This is largely because PET imagers consist of several detection rings, each comprising many small, high-cost position sensitive detectors, whereas the comparatively simple modified dual-headed gamma camera consists of only two large position-sensitive detectors.
  • a dedicated PET imager requires its own source of short-lived radionuclides on-site; for example a cyclotron.
  • a complete dedicated PET facility, with imager, cyclotron, and associated buildings costs can be very expensive and accordingly, dedicated PET imaging is restricted to research use in specialist centres. Routine diagnostic scans are simply too
  • an imaging apparatus having first and second gamma cameras, arranged such that the gamma cameras are substantially diametrically opposed, means for detecting an event, processing means for counting events and for manipulating data obtained from events so as to construct an image of an object, characterised in that the solid angle ofthe gamma cameras is not limited by collimation.
  • Preferably shielding vanes are provided to reduce the number of single events.
  • a thin layer of scintillation crystal material is provided, typically of the order of thickness of 0.095 m, which is the industry standard thickness.
  • Improved sensitivity may be obtained by removing the collimators which were present on prior art systems. This permits the modified dual-headed gamma camera to be used in coincidence mode without collimators and increases the sensitivity ofthe apparatus over existing collimated systems.
  • the count rate is improved by using digital event detectors and pulse separation techniques.
  • the invention is a unit that is used to enhance the performance coincidence detection and image reconstruction to a modified dual-headed gamma camera to enable it to be used effectively in routine imaging of PET radionuclides. It takes the position and energy signals from each detector and uses them to determine that an event has occurred, to distinguish overlapping events, and to perform the image reconstruction using an efficient technique.
  • the unit may be added as a stand-alone device onto existing gamma camera designs or it may be incorporated as a switch-in integral module in future designs. In either case, the invention can be easily disabled to allow the system to operate in single-photon mode.
  • Shielding vanes may be necessary to reduce the number of single events reaching the detectors.
  • Detector tuning and calibration, and gantry operation may be controlled by the host system.
  • DSP digital signal processing
  • Sensitivity and count rate performance may be further improved by application of a fully 3 -dimensional reconstruction technique because this uses approximately all ofthe time coincidence events; unlike previous systems, which have rejected many of these counts.
  • the invention described in this Application describes a viable low-cost alternative to the dedicated PET imager that addresses system performance and is therefore the first device that will facilitate the use of PET radionuclides in routine medical practice.
  • the imaging apparatus allows single-photon, as well as positron- emission tomographic, whole body and planar scans to be carried out on the same system.
  • the imaging apparatus' performance cannot compete with a dedicated PET scanner, and it can only image longer-lived radionuclides, such as 18 F, but can produce clinically useful results and with its low cost it is the only practical method for routine imaging of PET radiopharmaceuticals.
  • the first is a Multiple Wire Proportional Chamber (MWPC).
  • the multiple wire proportional chamber is a form of position sensitive ionisation chamber comprising a large area gas filled chamber containing an array of fine wires.
  • MWPCs require a solid photon-electron converter within the chamber to make them suitable for detecting 51 IkeV gamma rays.
  • Ott (1993) reports that most MWPC detectors achieve a single-photon detection efficiency up to 20% at 51 IkeV. While this is better than gamma-camera efficiency, the detectors are very susceptible to scattered radiation and have no energy discrimination ability.
  • Fluoride scintillator in conjunction with a MWPC filled with UV -photon sensitive gas may provide an effective, low-cost alternative to BGO systems.
  • the second low cost alternative is a Cost-Reduced PET Imager.
  • the low data rate of MWPC detectors has discouraged some researchers from this approach.
  • a group in Geneva are working on a project to develop a rotating PET scanner using BGO block detectors, with the help of Siemens, (Townsend et al. 1993).
  • This system is a cut-down version of a static dedicated PET imager, and comprises two rotating, opposing sections of a PET detector operating in coincidence.
  • This system has the advantage that it uses detectors optimised for 51 IkeV. However, it cannot image single photon emitting radionuclides, and, although it is doubtless less expensive than a full dedicated PET imager, it is estimated to be more expensive than the present invention.
  • gamma camera systems make them an attractive possible alternative for low-cost and hence routine PET imaging.
  • the gamma camera is optimised to image at energies around 140keV and cannot operate at 51 IkeV without modification. If the modification is such that the original system set-up can be reverted to easily, these systems have the advantage that they have the dual capability of single-photon (SPET) work and routine PET imaging.
  • SPET single-photon
  • the first problem with imaging 51 IkeV radiation with a collimated gamma camera is that the probability of a photopeak interaction in the crystal is much lower than at 140keV.
  • Production gamma cameras usually have a crystal thickness of 9.5mm, at which the photopeak interaction efficiency for 51 IkeV radiation is only 11.8%, compared with 90% at 140keV (Anger 1964).
  • This combined with the inherently poor efficiency of the collimated method of image formation, results in a poor system sensitivity.
  • a typical level of administered activity needed to obtain an FDG scan in a reasonable time with a collimated gamma camera system is around 370MBq, resulting in a total radiation dose to the patient of approximately lOmSv. Such levels are too high to be recommended for routine use in many clinical situations.
  • a second problem comes from the relatively low attenuation coefficient for 511 keV gamma rays, compared with those for gamma rays at the energies for which the gamma camera is designed as can
  • the present invention has a sensitivity which is many times that ofthe collimated gamma camera and therefore can be used with lower activities and consequently lower radiation doses to the patient. It is therefore suited to routine diagnostic imaging. It does not suffer from weight restriction problems, because it does not need heavy collimators, and the spatial resolution is likely to be better than collimated systems because the present invention does not rely on collimator hole geometry.
  • the count-rate limiting factor is that the system must handle a large proportion of events where only one of the two gamma rays has been detected.
  • a true coincidence count rate of lOkc/s means that each detector is handling a singles count rate of around 200kc/s. At this singles rate, approximately 18% of events from a single detector overlap one another to some extent and the processing electronics must be able to handle this without loss of counts and with minimal loss of resolution. The systems used in early experiments were unable to achieve this.
  • the main advantage of the present invention over the collimated system is its efficiency, because although the chances of seeing a twin gamma ray event is only 12% of that of seeing one of the two gamma rays (as in the collimated case), there are no collimators and thus the solid angle is increased by around a factor of 1000. However, this solid angle must be exploited if the system is to be of practical use. Early experiments did not exploit the increased solid angle available because of problems in the reconstruction process. Only when an impracticably small reconstruction volume was adopted did the early attempts yield clinically useful count rates.
  • the present invention adopts a novel approach to overcome the aforementioned problems. It uses fast digital hardware to improve the pulse separation and a modified data- efficient reconstruction technique to include 100% of recorded data in the image reconstruction.
  • the resulting device has a count rate capability which is suitable for most routine clinical applications with the longer lived PET radiopharmaceuticals.
  • Figure 1 shows a cross-sectional schematic of a typical gamma camera
  • Figure 2 shows a Schematic view of a dedicated PET imager in operation
  • Figure 3 shows a graph of linear attenuation coefficient against energy
  • Figure 4 shows a schematic overview ofthe invention in use with a modified dual-headed gamma camera
  • Figure 5 shows a functional schematic view of PET unit
  • Figure 6 illustrates the formation of a coincidence trigger, and pulse pileup separation
  • Figure 7 shows diagrammatically that the point source response function is not spatially invariant for a detector pair working in coincidence, in which shaded areas represent the point response functions for a point near the edge ofthe FOV (a) compared with that near the centre ofthe FOV (b); and
  • Figure 8 shows an axial slice of shell phantom containing two point sources: a) shows a simulated back-projection b) shows a forward projection from initial estimate c) shows a final deconvolved image d) shows an original simulated emitter distribution.
  • Figure 1 shows diagrammatically a gamma camera comprising a plurality of photomultiplier tubes 2, a lightguide 3, a scintillator 4 and a collimator 5.
  • Figure 2 is a diagrammatical view of a patient 6 in a position sensitive detector ring 7 in a PET imager.
  • FIG. 4 shows how energy and position signals from detectors 10 of a gamma camera 11 are fed into a processing unit 12.
  • the signals represent raw analogue energy and position signals (E, X+, X-, Y+, Y-) from an Anger position matrix 14.
  • E, X+, X-, Y+, Y- raw analogue energy and position signals
  • the signals have the pulse-shaping particular to the make of gamma camera 11, but unless pole-zero cancellation has been implemented the signals will probably exhibit the characteristic time constant ofthe scintillator. This is 230ns for Na ⁇ (Tl) .
  • the event trigger is carried out using the energy sum signal from each detector A and B (channel A and B respectively).
  • Suitably conditioned signals i.e. with a suitable risetime
  • conditioned signals are fast-digitised using a flash converter running at a clock speed of 40MHz or more.
  • the resolution ofthe flash converter needs to be such that final integrated signal resolution is not unduly impaired by over-course digitisation.
  • Adjacent sets of samples are tested on each clock cycle to see if there is a gradient greater than a predetermined threshold until both channels show a positive result (e.g. trigger point Q as shown in Figure 6), at which point a coincidence event is detected.
  • the duration ofthe coincidence time window determines the number of random coincidences recorded (i.e. events which are causally unconnected).
  • the coincidence window needs to be as small as possible but at the same time keeping a sufficient signal-to-noise ratio. Digitisation at the speeds suggested above facilitate the use of coincidence windows of less than 100ns.
  • the event stream from each detector suffers from overlapping events. At a singles count rate of 200kc/s, 18% ofthe events in each stream overlap one another at least once. It is therefore necessary to take this into account when integrating under the position and energy signals. This can be done by using the single channel trigger to determine when each pulse starts. By integrating the signals into accumulators and storing the sums at the onset of each trigger it is possible to separate the overlapping events by using knowledge ofthe characteristic decay ofthe signals. Details of pulse integration are shown in Figure 5.
  • Figure 6 depicts one possible sequence of overlapping events on energy channels A and B. Only two events are coincident. P, Q and R are the first three trigger events detected by the gradient test. At trigger P (originating from channel A), channels A and B start being integrated into a separate accumulator, until trigger Q occurs. The integrated sums for channel A and B are stored temporarily. Note that trigger Q is also a coincidence trigger, because it is detected in both channels A and B. The accumulators for channels A and B are then cleared and start integrating again, until the next trigger (or a suitable time longer than the typical signal envelope). In this example a third trigger, R, occurs because of a lone pulse on channel B. At this point, the accumulator contents for the integral Q-R are stored temporarily and this trigger/integration process continues.
  • the system attempts to separate the events on channels A and B.
  • the coincident event is "sitting on the back" of a preceding event.
  • the stored integral P-Q for that channel may be used to determine the residual tail of the preceding event upon which the event of interest rests. This can be done by multiplying the integral by a value stored in a lookup table containing the characteristic decay of a signal.
  • the residual tail is then subtracted from the integral stored for Q-R, which is also used to determine the residual tail which was lost because of the trigger from a following event on channel B. This residual tail is than added to the result to give the value ofthe coincident event energy signal for channel A.
  • the energy value for channel B is arrived at in exactly the same way. Applying the same algorithm has the same effect.
  • the position signals for channels A and B are separated in a similar fashion, except that they take their trigger points levels from the energy channels. In this way it is possible to detect coincident events and handle pulse pileup without loss of events or undue loss in resolution. Baseline shifts caused by the build up of the delayed component of the scintillator output can also be corrected digitally during the integration process.
  • DSPs digital signal processors
  • the denominator serves to remove energy dependence in the result.
  • Some systems may use the total position sum (X+ + Y+ + X- + Y-) here.
  • the Y coordinate is achieved in a similar fashion.
  • the spatial coordinates are then used to determine the energy and spatial corrections which need to be applied to the energy and position ofthe events in each detector. This can also be done digitally using lookup tables. Energy and spatial errors are an unavoidable feature of gamma camera design and there are many techniques to determine and correct for them.
  • an energy discrimination level should be adopted for each coincidence event. Both single events making up a coincidence must be included in the window for the event to be accepted.
  • the choice ofthe position and width of the energy window will depend on the energy resolution of the system and on the usefulness of Compton-scattered radiation. Some radiation will Compton-scatter within the detector such that only part ofthe energy of a gamma ray is deposited in the scintillator 4, others will have scattered in the patient and therefore lost useful spatial information, some will have scattered in both patient 6 and detector 7. However, many of these events may possess correct spatial information because they have not been scattered in the patient.
  • the orientation ofthe line of response (LOR) between the two coordinates can be calculated.
  • the detector orientation and separation must be known and these values therefore need to be obtained from the host system at intervals during the acquisition.
  • the event is characterised by the energy of each gamma ray and a direction through space and can be used in the image reconstruction process. The reconstruction process is described in the next section.
  • the event information may either be stored in sequential 'list' mode for subsequent processing, or used in real ⁇ time back-projection if the processing equipment has sufficient capability. Data Efficient Image Reconstruction Technique.
  • Fully three-dimensional (3D) image reconstruction uses gamma-ray events with lines-of-flight having components in the axial direction as well as those travelling in the transaxial plane.
  • Single-photon (SPECT) systems employing parallel-hole collimators and typical multi-ring PET scanners with inter-ring septa use only those gamma-rays travelling at angles close to a transaxial plane and are inherently inefficient for this reason.
  • Jeongs 1980, Colsher 1980, Daube-Witherspoon and Muehllehner 1987) and most attempts at its solution comprise a process where each individual event, or a set of filtered
  • 2D projections is back-projected through a 3D volume, followed by the deconvolution of the system point source response function from the backprojected image in Fourier space.
  • the 3D back-projected image g(x) can be thought of as the convolution of the radionuclide distribution f(x) with the backprojected image of a point source measured with the same system, h(x,_x'), where h(x, x') denotes the contribution to the image at position x due to a point source positioned at x' :
  • the true radionuclide distribution, f(x). can be isolated by deconvolution via use of the discrete Fast Fourier Transform. Analytic expressions for the point response function (or the appropriate filter for the deconvolution) have been derived (Colsher 1980).
  • An iterative approach which begins by reconstructing a small region of the backprojected image using an appropriate value of limiting q may be used.
  • the region chosen is small enough so that the point-response function changes little within it.
  • a surrounding region of the backprojected region is then chosen, and for which a smaller value of q is appropriate.
  • the original image is deliberately convolved with the point- response function corresponding to the new angle, is then added to the new region and the result is deconvolved with the new point response function. The process iterates until the entire region of interest is reconstructed.
  • Another approach can be summarised as follows: first, a low-statistics image is reconstructed using a limited angular acceptance which meets the spatial invariance condition ofthe point response function.
  • the point response function is the deconvolved from this back-projection (usually by prefiltering the projections) giving a first estimate of the radionuclide distribution.
  • This distribution is then forward-projected to provide an estimate ofthe projections not measured because of the limited geometry ofthe detectors. Scaling and adding the estimated missing projections to those actually measured allows the result, which incorporates about 90% ofthe detected data, to be successfully filtered with an appropriate spatially-invariant point response function.
  • LOR line of response
  • the positron may have moved some way before annihilating, or the two gamma rays may not be quite collinear due to the motion ofthe centre of mass ofthe positron-electron pair. Or this may occur because the two detected gamma rays forming the coincidence are causally unconnected. Nevertheless, in a real system the point of origin is assumed to lie on the LOR, and the fact it may not is a contributing factor to noise.
  • the first step in the backproj ection process is the calculation of the points of intersection ofthe LOR with the bounding box. This helps to define the range of voxels through which the LOR has to be tracked.
  • the point of intersection ofthe LOR with each of the inter-voxel planes it crosses is calculated and the distance of this intersection from the point that the LOR enters the bounding box is stored along with the index ofthe plane it is crossing. If the LOR is parallel to a voxel side, it will have a length in that voxel of 1.0. If it passes through diametrically opposite corners of a voxel it will have a length in that voxel of 1.732. These distances are then sorted sequentially as follows.
  • the shortest distance stored is the length ofthe LOR lying within the first voxel. This length is added to the image memory at the address of that voxel. The next largest distance minus the previous distance is the length ofthe LOR passing through the adjacent voxel, and is added to the image memory at that voxel's address. This cycle is repeated until the entire LOR has been accounted for in the reconstruction volume.
  • the backprojected image represents the distribution of emission points smeared with the backprojected image of a point source (the point response function). For a detector with limited geometry, the point response function will vary with position, and the distribution of emission points for an extended source cannot be retrieved simply by deconvolving the backprojected image with some nominal point response without generating artefacts in the result.
  • the backprojected image of events meeting the condition of point response function invariance can yield a low-statistics estimate ofthe distribution of emission sites.
  • This low-statistics image can then be used to forward project generated events into those solid angles ofthe system where events are not normally detected.
  • this forward-projected image is scaled and added to the true backprojected image, the result can be deconvolved with the point response function obtained for a 4 ⁇ detector.
  • the forward projection process works in the program as follows. First, a special patient scan is carried out where the angle of acceptance is limited to the maximum that can be achieved for a particular imaging volume and detector size. In practice, this could form a subset ofthe events detected in normal acquisition, and a copy of these could be stored separately during the acquisition: a separate acquisition may not be required. These ⁇ - limited events are backprojected as described above, and the resultant image is then deconvolved with the corresponding point response function. This response function is determined empirically by generating events under the same conditions from a point-source placed at the centre ofthe imaged volume. A frequency filter, such as a Harming window, (see e.g.
  • the deconvolved low-statistics image is then used as the generator for artificial, forward- projected events.
  • the low-statistics generating image is prepared for event generation by first setting voxels with contents lower than a suitable threshold. The remaining voxels are then scaled so that their total equals the number of events to be generated.
  • the forward-projection process then steps through the non-zero voxels and uses their centres as emission points. Event direction cosines are generated as before.
  • the deconvolution ofthe empirically-determined point response functions from the backprojected data is carried out in the Fourier domain.
  • the discrete 3- dimensional Fast Fourier Transform FFT
  • the Nyquist critical frequency is set by the imaging matrix.
  • a filter such as the Harming filter (as used by Colsher 1980), with the cut-off frequency set to the Nyquist frequency is used prior to carrying out the final inverse transform to reduce the effect of noise on the deconvolution.
  • the backprojected image (which is really a convolution) then effectively has a sha ⁇ edge which gives artefacts when deconvolved. If there is zero-padding of the data, it is necessary to deconvolve one edge of the data using information from the opposite edge. The only way to alleviate this problem is to ensure that the important region of the data is at least a point response function's width inside the image boundary. This is an important factor in the implementation of this technique.
  • An important potential use of a gamma-camera PET system is oncological whole- body scanning and static imaging.
  • detectors In whole body or static planar imaging, detectors do not rotate about the patient and therefore there is further restriction to the directions ofthe lines of response (LORs) available for use in image reconstruction.
  • LORs lines of response
  • the static and whole body imaging capability of such a system are considered to be a particularly attractive features and may be implemented in a preferred system.
  • the reconstruction algorithm uses as many of these LORs as possible.
  • the LORs pass through the imaging volume at different angles they carry information about the depth distribution of the emitting radiopharmaceutical so as to extract this information.
  • a solution is proposed where a subset of coincident events, which have a limited range of acceptance angles, are used directly to form a low-statistics, real-time image. This is really a form of electronic collimation.
  • the subset of gamma ray events might be that used to determine the first estimate for the forward-projection process if the solid angle is small enough to form a suitable image.
  • Activity used in the system sensitivity calculations is taken to be the annihilation event rate in the phantom.
  • the system spatial resolution is estimated at around 7.5 mm throughout the imaging volume.
  • Figure 7 shows axial slices ofthe reconstruction of a simulated emission phantom during stages in the reconstruction process.

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Abstract

Appareil d'imagerie (1) comprenant une camera gamma à deux têtes modifiée pour utilisation en tomographie par émission de positrons. Des caméras gamma sont utilisées dans la tomographie par émission de positrons (PET). Des collimateurs sur chaque caméra gamma réduisent l'intensité du signal que les caméras gamma sont capables de détecter. L'invention concerne une caméra gamma modifiée à deux têtes, capable de donner une représentation visuelle d'événement isolés d'émission de photons, sans limite quant à l'angle solide. Ce nouveau système est plus sensible que les anciens systèmes. L'utilisation de techniques de traitement numérique de signaux (DSP) rapides et de déconvolution au moyen d'une fonction de réponse par point permet d'obtenir une image améliorée d'un objet.
PCT/GB1996/002114 1995-08-25 1996-08-27 Appareil d'imagerie WO1997008569A1 (fr)

Priority Applications (4)

Application Number Priority Date Filing Date Title
IL12345796A IL123457A0 (en) 1995-08-25 1996-08-27 Imaging apparatus
EP96928600A EP0846275A1 (fr) 1995-08-25 1996-08-27 Appareil d'imagerie
AU68320/96A AU6832096A (en) 1995-08-25 1996-08-27 An imaging apparatus
JP9509999A JPH11511551A (ja) 1995-08-25 1996-08-27 イメージング装置

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GB9517491.8 1995-08-25
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WO2010048363A2 (fr) * 2008-10-24 2010-04-29 University Of Washington Estimation de ligne de réponse pour détecteur de tomographie à émission de positon haute définition
US8718973B2 (en) 2011-09-09 2014-05-06 Kabushiki Kaisha Toshiba Method, device, and system for calculating a geometric system model using an area-simulating-volume algorithm in three dimensional reconstruction
CN106057050A (zh) * 2016-08-16 2016-10-26 东北大学 一种基于gate平台的荧光ct系统模拟方法
US11061147B2 (en) 2019-03-01 2021-07-13 University Of Washington Accurate photon depth-of-interaction decoding and calibration of multiplexed detector modules

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KR101265256B1 (ko) 2010-12-30 2013-05-20 한국원자력연구원 방사선원의 3차원 분포 영상 처리 방법 및 이를 이용한 시스템
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Cited By (9)

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EP1331493A1 (fr) * 2000-10-04 2003-07-30 Kabushiki Kaisha Toshiba Dispositif de detection de rayonnement pour dispositif de diagnostic de medecine nucleaire et procede de detection associe
EP1331493A4 (fr) * 2000-10-04 2007-03-14 Toshiba Kk Dispositif de detection de rayonnement pour dispositif de diagnostic de medecine nucleaire et procede de detection associe
EP3153889A1 (fr) * 2000-10-04 2017-04-12 Toshiba Medical Systems Corporation Dispositif de détection de radiations et procédé de détection de radiations pour appareil de diagnostic médical nucléaire
WO2010048363A2 (fr) * 2008-10-24 2010-04-29 University Of Washington Estimation de ligne de réponse pour détecteur de tomographie à émission de positon haute définition
WO2010048363A3 (fr) * 2008-10-24 2010-07-08 University Of Washington Estimation de ligne de réponse pour détecteur de tomographie à émission de positon haute définition
US8716669B2 (en) 2008-10-24 2014-05-06 University Of Washington Line of response estimation for high-resolution PET detector
US8718973B2 (en) 2011-09-09 2014-05-06 Kabushiki Kaisha Toshiba Method, device, and system for calculating a geometric system model using an area-simulating-volume algorithm in three dimensional reconstruction
CN106057050A (zh) * 2016-08-16 2016-10-26 东北大学 一种基于gate平台的荧光ct系统模拟方法
US11061147B2 (en) 2019-03-01 2021-07-13 University Of Washington Accurate photon depth-of-interaction decoding and calibration of multiplexed detector modules

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EP0846275A1 (fr) 1998-06-10
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AU6832096A (en) 1997-03-19
CA2230470A1 (fr) 1997-03-06
IL123457A0 (en) 1998-09-24

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