CA1245375A - Scintillation detector for tomographs - Google Patents
Scintillation detector for tomographsInfo
- Publication number
- CA1245375A CA1245375A CA000510983A CA510983A CA1245375A CA 1245375 A CA1245375 A CA 1245375A CA 000510983 A CA000510983 A CA 000510983A CA 510983 A CA510983 A CA 510983A CA 1245375 A CA1245375 A CA 1245375A
- Authority
- CA
- Canada
- Prior art keywords
- scintillator
- photodetector
- scintillation
- array
- detector
- Prior art date
- Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
- Expired
Links
Classifications
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/20—Measuring radiation intensity with scintillation detectors
- G01T1/2018—Scintillation-photodiode combinations
- G01T1/20185—Coupling means between the photodiode and the scintillator, e.g. optical couplings using adhesives with wavelength-shifting fibres
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/161—Applications in the field of nuclear medicine, e.g. in vivo counting
- G01T1/164—Scintigraphy
- G01T1/1641—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras
- G01T1/1644—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras using an array of optically separate scintillation elements permitting direct location of scintillations
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/20—Measuring radiation intensity with scintillation detectors
- G01T1/2018—Scintillation-photodiode combinations
- G01T1/20183—Arrangements for preventing or correcting crosstalk, e.g. optical or electrical arrangements for correcting crosstalk
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/20—Measuring radiation intensity with scintillation detectors
- G01T1/2018—Scintillation-photodiode combinations
- G01T1/20187—Position of the scintillator with respect to the photodiode, e.g. photodiode surrounding the crystal, the crystal surrounding the photodiode, shape or size of the scintillator
-
- 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
The present invention relates to a scintil-lation detector for a tomograph comprising at least two scintillators having different scintillation character-istics and being optically coupled to a photodetector.
The present invention relates to a scintil-lation detector for a tomograph comprising at least two scintillators having different scintillation character-istics and being optically coupled to a photodetector.
Description
2~S3~5 FIELD OF THE INVENTION
The present invention relates to a novel scintillation detector for a tomograph. By extent the invention also comprehends a detector array and a tomograph using such a scintillation detector.
BACKGROUND OF THE INVENTION
Nuclear medicine uses radiopharmaceutical products marked by radioactive isotopes emitting gamma radiation for obtaining information on the physiolog-ical processes of the human body. The progression of the radioactive products toward an organ or its accumu-lation in that organ are followed from outside the body by means of a gamma radiation detector, more or less sophisticated, the most common being the scintillation camera or gamma camera of the Anger type. The image ob'ained by such a camera represents the projection on a reference plane of the three dimensional distribution of the radiopharmaceu-tical product. A three dimensional image may be obtained by applying the well-known prin-ciples of t~e axial tomography.
Ano-ther approach, perhaps less popular but offering many advantages, uses as tracers atoms emit-ting positrons. The positrons annihilate themselves with electrons and generate two gammas of 511 keV e-.~ .
`'~;^
- 2 - ~ ~45~S
mitted at 180 relatively -to each other. By detecting coincidentally these two gammas with two diametrically opposite detectors, the trajectory on whlch the disin-tegration has occurred may be determined. By super-posing, by means of known techniques of tomographicreconstruction, the multiple trajectories measured by an array of detectors surrounding the source, the dis-tribution of the radioactivity in the volume enclosed by the array of detectors may be derived. The three dimensional image may be obtained by the juxtaposition of two-dimensior3al images of the radioactivity distribution in adjacent planes, or by direct recon-struction from the multiple inter-plane trajectories.
A typical tomograph comprises an array of individual detectors separated or not by septas. The detectors may be grouped in the array in one or more rings. The array surrounds the body -to be scanned and a suitable electronic circuitry processes the electric sig~3als generated by the detectors so as to obtain the desired image. Typically, the diameter of a detector ring varies from 50 to 100 cm, according to whether the apparatus is adapted for scanning the brain or the entire body. The majority of the existing cameras use (Bi4Ge3O12) scintillation detectors (hereinafter "BGO") coupled to photomultiplier tubes. Such cameras have a spatial resolution in the order of one centimeter.
Certain models can reach a resolution of 4 to 6 mm lZ~L5375 FW~M. These resolution values are not the inherent theoretical limits fixed by the positron range in tissues and the non-colinearity of emission of anni-hilation gamma-rays, but rather represen'c a compromise resulting from physical and technological restrains.
The improvement of the resolution of a tomo-graph up to three millimeters ~WHM, which is close to the theoricical limit, is highly desirable~ However, the parallax error which exists in a detector ring has, up to now prevented such improvement out of the region very close to the center of the tomograph.
The parallax error may briefly be defined as the lack of information on the radial position of interaction of a gamma ray in a given detector of the ring. The position of interaction in a detector is a function of probability. In some cases, a gamma ray may pass through a detector without interacting therein and interact in an adjacent detector. Therefore, when a detector generates an output signal, indicating the occurrence of an interaction, the gamma ray may come from anywhere within the channel defined by the pro-jection of the volume of the detector, with a distribu-tion given by -the probability of intexaction of the gamma in this detector (the so-called "aperture function").
At first sight, a simple way -to resolve the parallax problem is to reduce the depth of the detec-~2~S3~;
tors to lower the volume of the projection channel to,in turn, reduce the incertitude region and the parallax error. However, a thinner de-tector implies that more gamma rays will pass throughout without interacting, resulting in a loss of efficiency which may not be acceptable for clinical applications. In a similar manner, the increase OL the ring diameter will reduce the paralla~ error, involving a reduction of efficiency of the device and an increase of the costs due to the larger number of detectors necessary to construct a bigger ring.
An alternative solution which has been adopted in several of the commercially available tomographs consists of inserting septas of a heavy metal (Tungsten, Gold or Uranium) between the detectors to reduce the possibility of a gamma ray passing from one detector into another. To stop efficiently a gamma ray of 511 keV, the septas must be sufficiently thick (more than one mm). ~lowever, in a high resolution sys-tem where the detectors are typically 3 or 4 milli-meters thick, the drop of efficiency of 25 to 50~ which would results from the use of such septas, is obviously undesirable.
OBJECTS AND STATEMENT OE THE INVENTION
An object of the pxesent invention is a ~245~75 scintillation detector for a tomograpn, the detector having an increased resolution.
Another object of the invention is an array of scintillation detectors for a tomograph, the array having an increased resolution.
A further object of thls invention is a tomo-graph with an improved resolution.
The objects of this invention are achieved by providing a scintillation detector sensible to the position of interaction of a gamma ray therein. In other words, the position of interaction of the gamma ray in the detector may be determined with a certain precision, for reducing the parallax error.
In one embodiment, the detector comprises two scintillators having different scintillation character-istics and optically coupled to each other. To one of the scintillators is connected a photodetector which generates an electrical signal in response to a flash of light produced by one of the scintillators due to an interaction of a gamma ray. Since the scintillators have different scintillation characteristics, different signals will be generated by the photodetector depend-ing whether the gamma has interacted in the first or the second scintillator. By using known signal discrim-ination techniques, the scintillator in which theinteraction has occurred, may be determined.
For further increasing the resolution the qg ~ ~ ,.. , . .. ~
, - 6 - ~ 37~
detector may be formed of more than two scintillators.
It should be understood that the term "light"
includes not only visible light but also other types of electromagnetic radiations such as ultraviolet light or others.
The concept behind the scintillation detector of this invention is not restricted on]y to the detec-tion of gamma radiation. When other types of radiation are to be detected, appropriate scintillators respon-sive to the emitted radiation must be used for theconstruction of the detector.
Such variations of this invention are well within the reach and the knowledge of a man skilled in the art and for that reason they will not be explored in details here.
A plurality of detectors according to this invention are mounted together, and grouped together, preferably in one-dimensional or two-dimensional ar-rays. In a tomograph, each detector is formed ~y a plurality of scintillators and a photodetector, the photodetectors being mounted at the periphery of the ring. When a plurality of rings are used, they are mounted side by side so as to obtain images in a plurality of adjacent planes. Alternatively, two-dimensional arrays of detectors can be used, eachdetector being formed by a plurality of scintillators with the photodetectors mounted on top of the array.
.
~LZ4537S
The tomograph has the form of a cylindrical array of detectors.
Pxeferably, the scintillation detectors are optically isolated from each other in the array.
A tomograph according to this invention typically includes an array of scintillatlon detectors to which is connected a signal processing system, for analyzing the electric signals generated by the photo-detector so as to construct an image on a monitor or a representation in any other form of the organ or the body ~hich is scanned.
The present invention comprises in a most general aspect a scintillation detector for a tomo-graph, the scintillation detector being adapted for detecting radiation, the detector comprising:
- a first scintillator;
- a second scintillator optically coupled to the first scintillator, the scintillators having different scintillation characteristics; and - a 2hotodetector optically coupled to the second scintillator, a radiation interaction in the second scintillator generating a flash of light which is detected by the photodetector, a radiation inter-action in the first scintillator generating a flash of light which is transmitted through the second scintil-lator and detected by the photodetector.
The invention further comprehends an array of -124S;~S
scintillation detectors lor a tomograph, the scintilla-tion detectors being adapted for detecting radiation, each scintillation detector including:
- a first scintillator;
- a second scintillator optically coupled to the first scintillator, the scintillators having dif-ferent scintillation characteristics; and - a pho-todetector optically coupled -to the second scintillator, a radiation interaction in the second scintillator generating a flash of light which is detected by the photodetector, a ra~iation inter-action in the first scintillator generating a flash of light which is transmi.tted through the second scintil-lator and is detected by the photodetector.
The invention further comprehends a tomograph for obtaining information on a human ~ody or an animal, said tomograph comprising:
- radiation detecting rneans, which includes an array of scintillation detectors comprising a plurality of scintillation detectors, each de-tector including:
a) a first scintillator;
b) a second scintillator optically coupled -to the first scintillator, the scintillators having different scintillation characteristics; and c) a photodetector optically coupled to the second scintillator, a radiation interaction in the ....
i37S
second scintillator generating a flash of light which is detected by -the photodetec'cor which generates, in turn, an electric signal, a radiation interaction in the first scintillator generating a flash of light which is transmitted through the second scintillator and is detected by the photodetector which generates, in turn, an electric signal;
- processing means operatively connected to the photodetectors of the scintillation detectors of said array for processing the signal generated by the photodetectors to provide said information.
BRIEF DESCRIPTION OF THE DRAWINGS
Figure 1, is a perspective view of a prior art array of scintillation detectors forming a ring;
Figure 2, is an enlarged perspective view of a portion of the array shown in Figure l;
Figure 3, is a schematical view of a detector ring, illustrating the parallax error phenomena, Figure 4, is a schematical view of a scintil-lation detector according to the present invention cou-pled to a signal discrimination circuit;
Figure 4a ~ 4b: Figure 4a is a diagram of the signals generated by a photodetector in response to an interaction of a gamma ray in three different scintillators. It shows the difference in decay time of ~24S375 the scintillation light in each of the scintillators.
Figure 4`~ is the diagram of the corresponding signals at the output of an integrating amplifier; and Figure 5, is a diagram of aperture functions lllustrating the resolution improvement obtained with the scintillation detector of this invention.
DESCRIPTION OF A ~RIOR ART DEVICE
_ A typical detector array 10 for a tomograph is illustrated in Figure 1. Array 10 is constituted by a plurality of individual scintillation detectors 12 grouped in a ring 14. Ring 14 is sandwiched between two conventional lead shielding rings 16.
Detector ring 10 is of a size to accomodate a human body 17 which has previously been injected with a substance producing an emission of gamma rays in oppo-site directions, at 180 rrom each other. The gamma rays are coincidently detected by two opposed scintil-lation detectors 12 to determine the trajectory of the gamma rays.
Suitable electronic detection and processing circuitrY is used to construct an image of the organ in which the radioactive substance is accumulated, in the plane of the detector ring 14, from t'ne signals generated by the detectors of array 10.
Referring to Figure 2, illustrating a group of three adjacent scintillation detectors 12a, 12b and 12c, the detectors comprising scintillators 18a, 18b and 18c, respectively, known in the art. When a gamma ray passes through scintillator 18b, it interacts therein and produces a flash of light detected by a photodetector 20 (u~ually a photomultiplier tube), mounted on top of scintillator 18b. Tungsten septas 22 may be inserted between the detectors so as to prevent the passage of gamma rays from one detector to another.
Referring to Figure 3, when gamma rays are emitted from the human body 17, near the periphery of the detector ring 14, they penetrate the scintillator 18b at an incident angle which increases as the point of emission of the gamma rays is near the periphery of ring 14. In the example given in Figure 3, the incident angle is of 30, but the following also holds true for other values of incident angles.
When a gamma ray penetrates scintillator 18b the position of interaction in the scintillator is a function of probability. In extreme cases, the gamma ray may pass through scintillator 18b without inter-acting, penetrate crystal 18a, in the absence of Tungsten septas, and interact in scintillator 18a.
Similarly, a gamma ray may pass through scintillator 18c and interact in scintillator 18b. Therefore, when a detector generates an output signal, a gamma ray which has interacted therein, may have been emitted anywhere i ~A
,,,~,;
' - 12 - ~ ~ ~53~
within the zone identified by the reference letter A.
Zone A has a width D which corresponds to the uncer-tainty on the position of the source for an incidence angle of 30. This incertitude which has previously been defined as the parallax error, is obviously unde-sirable and increases as the position of emission of the gamma ray is near the periphery of ring 14.
DESCRIPTION OF A PREFERRED EMBODIMENT
Figure 4 illustrates schematically a detector 23 according to the present invention which comprises three scintillators 24, 26 and 28 respectively, opti-cally coupled to each other through optical contacts 30. Each scin-tillator has different scintillation characteristics. A photodetector 32 such as an avalanche photodiode is mounted to scintillator 28.
As an example, avalanche photodiodes manufac-tured by RCA (trademark) and sold under the part numberC30994E, may be used for the construction of scintilla-tion detectors according to this invention.
The assembly of scintillators 24, 26 and 28 defines a light guide which transmits the flash of light generated in response to an interaction of a gamma ray, in any one of the scintillators to the - 13 - ~ 37~
photodetector 32.
Scintillation detector 23 is connected to an amplification and signal discrimination circuit 34 comprising an in-tegrating amplifier 36 connected to photodetector 32. A pulse shape analyzer 38 is con-nected to amplifier 36.
Since -the sci.ntillators 24, 26 and 28 have different scintillation characteristics, when a gamma interacts in detector 23, it suffices to observe the decay time of the output signal generated by photode-tector 32 or the rise time of the integrated signal at the output of integrated amplifier 36 to determine in which scintillator the interaction has occured. Figures 4a and 4b are diagrams of the output signals from photodetector 32 and from integrating amplifier 36, respectively, produced in response to an interaction in each scintillator of detector 23. The decay and the rise times of the signals associated with each scin-tillator are di.fferent which allows to determine in which scintillator the gamma ray has interacted.
Electronic circuit 34 for discriminating signals having different rise or decay times is well known in the art and, for that reason, it will not be described i.n details here.
.. ~, .
- 13a ~ S3~
~ y constructing each detector of a plurality of individual scintillators, results, for all practical purposes, in a reduction of the depth of the detector without a substantial reduction in the efficiency thereof.
~LX~53~;
Figure 5 shows the resolution improvement which may be obtained with the detector according to the present invention. It may be observed that, for a prior art detector formed by a single scintillator having a depth of 20 mm irradiated at an angle of 30, the resolution is of 5.1 mm. However, when a detector according to the present invention, formed by 4 scin-tillator crystals having each a depth of 5 mm is used, the resolution is of 2.2 mm, a significant improvement.
However, the overall depth of the detector is still 20 mm which implies that there is little or no drop in tne efficiency.
A small loss of efficiency may be expected in the multiscintillator system according to this inven-tion resulting from the use of scintillation crystalswhich have less ability to stop gamma rays than the BGO
crystal being one of the most efficient. For example, with a two scintillators detector, BGO/GSO (Gd2SiO5), a drop of efficiency of about 5% may be expected, which is tolerable.
A plurality of scintillation detectors are mounted side by side without interacting optically with each other. This may be achieved by optically isolating the detectors from each other. The scintillation detec-tors form a one-dimensional or a two-dimensional array.
In each detector of the array, the photodetector is mounted on one end of the scintillator assembly and ~,..
. .
r - "
aligned with the scintillators. The array may have the shape of a ring, or any other shape surrounding the body to be examined, with the photodetectors extending on the periphery of the ring. With such an arrangement, a plurality of adjacent rings may be placed side by side along the same axis to obtain at the same time images in a plurality of adjacent planes. Alternative-ly, a two dimensional array may have the shape of a cylinder or any other shape surrounding the body to be examined with the photodetectors extending on the periphery of the cylinder, therefore allowing the three-dimensional image of a complete volume to be obtained simultaneously.
A tomograph according to this invention comprises one or more detector rings, or a cylinder or any other shape surrounding the body to be examined, formed by a two-dimensional array of detectors, to which is connected a signal processing and analysis circuitry, generally ~nown in the art. This circuitry permits to analyze the signals generated by the photo-detectors so as to construct an image on a monitor or in any other form of the specimen under observation.
It should be understood that the scope of the present invention is not intended to be limited to the specific preferred embodiment illustrated in the draw-ings and described above.
,~
,.
. ~, .
The present invention relates to a novel scintillation detector for a tomograph. By extent the invention also comprehends a detector array and a tomograph using such a scintillation detector.
BACKGROUND OF THE INVENTION
Nuclear medicine uses radiopharmaceutical products marked by radioactive isotopes emitting gamma radiation for obtaining information on the physiolog-ical processes of the human body. The progression of the radioactive products toward an organ or its accumu-lation in that organ are followed from outside the body by means of a gamma radiation detector, more or less sophisticated, the most common being the scintillation camera or gamma camera of the Anger type. The image ob'ained by such a camera represents the projection on a reference plane of the three dimensional distribution of the radiopharmaceu-tical product. A three dimensional image may be obtained by applying the well-known prin-ciples of t~e axial tomography.
Ano-ther approach, perhaps less popular but offering many advantages, uses as tracers atoms emit-ting positrons. The positrons annihilate themselves with electrons and generate two gammas of 511 keV e-.~ .
`'~;^
- 2 - ~ ~45~S
mitted at 180 relatively -to each other. By detecting coincidentally these two gammas with two diametrically opposite detectors, the trajectory on whlch the disin-tegration has occurred may be determined. By super-posing, by means of known techniques of tomographicreconstruction, the multiple trajectories measured by an array of detectors surrounding the source, the dis-tribution of the radioactivity in the volume enclosed by the array of detectors may be derived. The three dimensional image may be obtained by the juxtaposition of two-dimensior3al images of the radioactivity distribution in adjacent planes, or by direct recon-struction from the multiple inter-plane trajectories.
A typical tomograph comprises an array of individual detectors separated or not by septas. The detectors may be grouped in the array in one or more rings. The array surrounds the body -to be scanned and a suitable electronic circuitry processes the electric sig~3als generated by the detectors so as to obtain the desired image. Typically, the diameter of a detector ring varies from 50 to 100 cm, according to whether the apparatus is adapted for scanning the brain or the entire body. The majority of the existing cameras use (Bi4Ge3O12) scintillation detectors (hereinafter "BGO") coupled to photomultiplier tubes. Such cameras have a spatial resolution in the order of one centimeter.
Certain models can reach a resolution of 4 to 6 mm lZ~L5375 FW~M. These resolution values are not the inherent theoretical limits fixed by the positron range in tissues and the non-colinearity of emission of anni-hilation gamma-rays, but rather represen'c a compromise resulting from physical and technological restrains.
The improvement of the resolution of a tomo-graph up to three millimeters ~WHM, which is close to the theoricical limit, is highly desirable~ However, the parallax error which exists in a detector ring has, up to now prevented such improvement out of the region very close to the center of the tomograph.
The parallax error may briefly be defined as the lack of information on the radial position of interaction of a gamma ray in a given detector of the ring. The position of interaction in a detector is a function of probability. In some cases, a gamma ray may pass through a detector without interacting therein and interact in an adjacent detector. Therefore, when a detector generates an output signal, indicating the occurrence of an interaction, the gamma ray may come from anywhere within the channel defined by the pro-jection of the volume of the detector, with a distribu-tion given by -the probability of intexaction of the gamma in this detector (the so-called "aperture function").
At first sight, a simple way -to resolve the parallax problem is to reduce the depth of the detec-~2~S3~;
tors to lower the volume of the projection channel to,in turn, reduce the incertitude region and the parallax error. However, a thinner de-tector implies that more gamma rays will pass throughout without interacting, resulting in a loss of efficiency which may not be acceptable for clinical applications. In a similar manner, the increase OL the ring diameter will reduce the paralla~ error, involving a reduction of efficiency of the device and an increase of the costs due to the larger number of detectors necessary to construct a bigger ring.
An alternative solution which has been adopted in several of the commercially available tomographs consists of inserting septas of a heavy metal (Tungsten, Gold or Uranium) between the detectors to reduce the possibility of a gamma ray passing from one detector into another. To stop efficiently a gamma ray of 511 keV, the septas must be sufficiently thick (more than one mm). ~lowever, in a high resolution sys-tem where the detectors are typically 3 or 4 milli-meters thick, the drop of efficiency of 25 to 50~ which would results from the use of such septas, is obviously undesirable.
OBJECTS AND STATEMENT OE THE INVENTION
An object of the pxesent invention is a ~245~75 scintillation detector for a tomograpn, the detector having an increased resolution.
Another object of the invention is an array of scintillation detectors for a tomograph, the array having an increased resolution.
A further object of thls invention is a tomo-graph with an improved resolution.
The objects of this invention are achieved by providing a scintillation detector sensible to the position of interaction of a gamma ray therein. In other words, the position of interaction of the gamma ray in the detector may be determined with a certain precision, for reducing the parallax error.
In one embodiment, the detector comprises two scintillators having different scintillation character-istics and optically coupled to each other. To one of the scintillators is connected a photodetector which generates an electrical signal in response to a flash of light produced by one of the scintillators due to an interaction of a gamma ray. Since the scintillators have different scintillation characteristics, different signals will be generated by the photodetector depend-ing whether the gamma has interacted in the first or the second scintillator. By using known signal discrim-ination techniques, the scintillator in which theinteraction has occurred, may be determined.
For further increasing the resolution the qg ~ ~ ,.. , . .. ~
, - 6 - ~ 37~
detector may be formed of more than two scintillators.
It should be understood that the term "light"
includes not only visible light but also other types of electromagnetic radiations such as ultraviolet light or others.
The concept behind the scintillation detector of this invention is not restricted on]y to the detec-tion of gamma radiation. When other types of radiation are to be detected, appropriate scintillators respon-sive to the emitted radiation must be used for theconstruction of the detector.
Such variations of this invention are well within the reach and the knowledge of a man skilled in the art and for that reason they will not be explored in details here.
A plurality of detectors according to this invention are mounted together, and grouped together, preferably in one-dimensional or two-dimensional ar-rays. In a tomograph, each detector is formed ~y a plurality of scintillators and a photodetector, the photodetectors being mounted at the periphery of the ring. When a plurality of rings are used, they are mounted side by side so as to obtain images in a plurality of adjacent planes. Alternatively, two-dimensional arrays of detectors can be used, eachdetector being formed by a plurality of scintillators with the photodetectors mounted on top of the array.
.
~LZ4537S
The tomograph has the form of a cylindrical array of detectors.
Pxeferably, the scintillation detectors are optically isolated from each other in the array.
A tomograph according to this invention typically includes an array of scintillatlon detectors to which is connected a signal processing system, for analyzing the electric signals generated by the photo-detector so as to construct an image on a monitor or a representation in any other form of the organ or the body ~hich is scanned.
The present invention comprises in a most general aspect a scintillation detector for a tomo-graph, the scintillation detector being adapted for detecting radiation, the detector comprising:
- a first scintillator;
- a second scintillator optically coupled to the first scintillator, the scintillators having different scintillation characteristics; and - a 2hotodetector optically coupled to the second scintillator, a radiation interaction in the second scintillator generating a flash of light which is detected by the photodetector, a radiation inter-action in the first scintillator generating a flash of light which is transmitted through the second scintil-lator and detected by the photodetector.
The invention further comprehends an array of -124S;~S
scintillation detectors lor a tomograph, the scintilla-tion detectors being adapted for detecting radiation, each scintillation detector including:
- a first scintillator;
- a second scintillator optically coupled to the first scintillator, the scintillators having dif-ferent scintillation characteristics; and - a pho-todetector optically coupled -to the second scintillator, a radiation interaction in the second scintillator generating a flash of light which is detected by the photodetector, a ra~iation inter-action in the first scintillator generating a flash of light which is transmi.tted through the second scintil-lator and is detected by the photodetector.
The invention further comprehends a tomograph for obtaining information on a human ~ody or an animal, said tomograph comprising:
- radiation detecting rneans, which includes an array of scintillation detectors comprising a plurality of scintillation detectors, each de-tector including:
a) a first scintillator;
b) a second scintillator optically coupled -to the first scintillator, the scintillators having different scintillation characteristics; and c) a photodetector optically coupled to the second scintillator, a radiation interaction in the ....
i37S
second scintillator generating a flash of light which is detected by -the photodetec'cor which generates, in turn, an electric signal, a radiation interaction in the first scintillator generating a flash of light which is transmitted through the second scintillator and is detected by the photodetector which generates, in turn, an electric signal;
- processing means operatively connected to the photodetectors of the scintillation detectors of said array for processing the signal generated by the photodetectors to provide said information.
BRIEF DESCRIPTION OF THE DRAWINGS
Figure 1, is a perspective view of a prior art array of scintillation detectors forming a ring;
Figure 2, is an enlarged perspective view of a portion of the array shown in Figure l;
Figure 3, is a schematical view of a detector ring, illustrating the parallax error phenomena, Figure 4, is a schematical view of a scintil-lation detector according to the present invention cou-pled to a signal discrimination circuit;
Figure 4a ~ 4b: Figure 4a is a diagram of the signals generated by a photodetector in response to an interaction of a gamma ray in three different scintillators. It shows the difference in decay time of ~24S375 the scintillation light in each of the scintillators.
Figure 4`~ is the diagram of the corresponding signals at the output of an integrating amplifier; and Figure 5, is a diagram of aperture functions lllustrating the resolution improvement obtained with the scintillation detector of this invention.
DESCRIPTION OF A ~RIOR ART DEVICE
_ A typical detector array 10 for a tomograph is illustrated in Figure 1. Array 10 is constituted by a plurality of individual scintillation detectors 12 grouped in a ring 14. Ring 14 is sandwiched between two conventional lead shielding rings 16.
Detector ring 10 is of a size to accomodate a human body 17 which has previously been injected with a substance producing an emission of gamma rays in oppo-site directions, at 180 rrom each other. The gamma rays are coincidently detected by two opposed scintil-lation detectors 12 to determine the trajectory of the gamma rays.
Suitable electronic detection and processing circuitrY is used to construct an image of the organ in which the radioactive substance is accumulated, in the plane of the detector ring 14, from t'ne signals generated by the detectors of array 10.
Referring to Figure 2, illustrating a group of three adjacent scintillation detectors 12a, 12b and 12c, the detectors comprising scintillators 18a, 18b and 18c, respectively, known in the art. When a gamma ray passes through scintillator 18b, it interacts therein and produces a flash of light detected by a photodetector 20 (u~ually a photomultiplier tube), mounted on top of scintillator 18b. Tungsten septas 22 may be inserted between the detectors so as to prevent the passage of gamma rays from one detector to another.
Referring to Figure 3, when gamma rays are emitted from the human body 17, near the periphery of the detector ring 14, they penetrate the scintillator 18b at an incident angle which increases as the point of emission of the gamma rays is near the periphery of ring 14. In the example given in Figure 3, the incident angle is of 30, but the following also holds true for other values of incident angles.
When a gamma ray penetrates scintillator 18b the position of interaction in the scintillator is a function of probability. In extreme cases, the gamma ray may pass through scintillator 18b without inter-acting, penetrate crystal 18a, in the absence of Tungsten septas, and interact in scintillator 18a.
Similarly, a gamma ray may pass through scintillator 18c and interact in scintillator 18b. Therefore, when a detector generates an output signal, a gamma ray which has interacted therein, may have been emitted anywhere i ~A
,,,~,;
' - 12 - ~ ~ ~53~
within the zone identified by the reference letter A.
Zone A has a width D which corresponds to the uncer-tainty on the position of the source for an incidence angle of 30. This incertitude which has previously been defined as the parallax error, is obviously unde-sirable and increases as the position of emission of the gamma ray is near the periphery of ring 14.
DESCRIPTION OF A PREFERRED EMBODIMENT
Figure 4 illustrates schematically a detector 23 according to the present invention which comprises three scintillators 24, 26 and 28 respectively, opti-cally coupled to each other through optical contacts 30. Each scin-tillator has different scintillation characteristics. A photodetector 32 such as an avalanche photodiode is mounted to scintillator 28.
As an example, avalanche photodiodes manufac-tured by RCA (trademark) and sold under the part numberC30994E, may be used for the construction of scintilla-tion detectors according to this invention.
The assembly of scintillators 24, 26 and 28 defines a light guide which transmits the flash of light generated in response to an interaction of a gamma ray, in any one of the scintillators to the - 13 - ~ 37~
photodetector 32.
Scintillation detector 23 is connected to an amplification and signal discrimination circuit 34 comprising an in-tegrating amplifier 36 connected to photodetector 32. A pulse shape analyzer 38 is con-nected to amplifier 36.
Since -the sci.ntillators 24, 26 and 28 have different scintillation characteristics, when a gamma interacts in detector 23, it suffices to observe the decay time of the output signal generated by photode-tector 32 or the rise time of the integrated signal at the output of integrated amplifier 36 to determine in which scintillator the interaction has occured. Figures 4a and 4b are diagrams of the output signals from photodetector 32 and from integrating amplifier 36, respectively, produced in response to an interaction in each scintillator of detector 23. The decay and the rise times of the signals associated with each scin-tillator are di.fferent which allows to determine in which scintillator the gamma ray has interacted.
Electronic circuit 34 for discriminating signals having different rise or decay times is well known in the art and, for that reason, it will not be described i.n details here.
.. ~, .
- 13a ~ S3~
~ y constructing each detector of a plurality of individual scintillators, results, for all practical purposes, in a reduction of the depth of the detector without a substantial reduction in the efficiency thereof.
~LX~53~;
Figure 5 shows the resolution improvement which may be obtained with the detector according to the present invention. It may be observed that, for a prior art detector formed by a single scintillator having a depth of 20 mm irradiated at an angle of 30, the resolution is of 5.1 mm. However, when a detector according to the present invention, formed by 4 scin-tillator crystals having each a depth of 5 mm is used, the resolution is of 2.2 mm, a significant improvement.
However, the overall depth of the detector is still 20 mm which implies that there is little or no drop in tne efficiency.
A small loss of efficiency may be expected in the multiscintillator system according to this inven-tion resulting from the use of scintillation crystalswhich have less ability to stop gamma rays than the BGO
crystal being one of the most efficient. For example, with a two scintillators detector, BGO/GSO (Gd2SiO5), a drop of efficiency of about 5% may be expected, which is tolerable.
A plurality of scintillation detectors are mounted side by side without interacting optically with each other. This may be achieved by optically isolating the detectors from each other. The scintillation detec-tors form a one-dimensional or a two-dimensional array.
In each detector of the array, the photodetector is mounted on one end of the scintillator assembly and ~,..
. .
r - "
aligned with the scintillators. The array may have the shape of a ring, or any other shape surrounding the body to be examined, with the photodetectors extending on the periphery of the ring. With such an arrangement, a plurality of adjacent rings may be placed side by side along the same axis to obtain at the same time images in a plurality of adjacent planes. Alternative-ly, a two dimensional array may have the shape of a cylinder or any other shape surrounding the body to be examined with the photodetectors extending on the periphery of the cylinder, therefore allowing the three-dimensional image of a complete volume to be obtained simultaneously.
A tomograph according to this invention comprises one or more detector rings, or a cylinder or any other shape surrounding the body to be examined, formed by a two-dimensional array of detectors, to which is connected a signal processing and analysis circuitry, generally ~nown in the art. This circuitry permits to analyze the signals generated by the photo-detectors so as to construct an image on a monitor or in any other form of the specimen under observation.
It should be understood that the scope of the present invention is not intended to be limited to the specific preferred embodiment illustrated in the draw-ings and described above.
,~
,.
. ~, .
Claims (19)
1. A scintillation detector for a tomograph for detecting radiation, said detector comprising:
- a first scintillator;
- a second scintillator optically coupled to said first scintillator, said scintillators having different scintillation characteristics; and - a photodetector optically coupled to said second scintillator, an interaction of radiation in said second scintillator generating a flash of light which is detected by said photodetector, an interaction of radiation in said first scintillator generating a flash of light which is transmitted through said second scintillator and detected by said photodetector.
- a first scintillator;
- a second scintillator optically coupled to said first scintillator, said scintillators having different scintillation characteristics; and - a photodetector optically coupled to said second scintillator, an interaction of radiation in said second scintillator generating a flash of light which is detected by said photodetector, an interaction of radiation in said first scintillator generating a flash of light which is transmitted through said second scintillator and detected by said photodetector.
2. A scintillation detector as defined in claim 1, wherein said first scintillator, said second scintillator and said photodetector are aligned.
3. A scintillation detector as defined in claim 1, wherein said radiation is gamma radiation.
4. A scintillation detector as defined in claim 1, wherein said photodetector is a photodiode.
5. A scintillation detector as defined in claim 4, wherein said photodiode is an avalanche photodiode.
6. An array of scintillation detectors for a tomograph, said scintillation detectors detecting radiation, each scintillation detector including:
- a first scintillator;
- a second scintillator optically coupled to said first scintillator, said scintillators having different scintillation characteristics; and - a photodetector optically coupled to said second scintillator, an interaction of radiation in said second scintillator generating a flash of light which is detected by said photodetector, an interaction of radiation in said first scintillator generating a flash of light which is transmitted through said second scintillator and is detected by said photodetector.
- a first scintillator;
- a second scintillator optically coupled to said first scintillator, said scintillators having different scintillation characteristics; and - a photodetector optically coupled to said second scintillator, an interaction of radiation in said second scintillator generating a flash of light which is detected by said photodetector, an interaction of radiation in said first scintillator generating a flash of light which is transmitted through said second scintillator and is detected by said photodetector.
7. An array as defined in claim 6, wherein said first scintillator, said second scintillator and said photodetector are aligned.
8. An array as defined in claim 6, wherein said radiation is gamma radiation.
9. An array as defined in claim 6, wherein said photodetector is a photodiode.
10. An array as defined in claim 9, wherein said photodiode is an avalanche photodiode.
11. An array as defined in claim 6, wherein the detectors of said array do not interact optically with each other.
12. A tomograph, for obtaining information on a human body, or an animal emitting radiation, said tomograph comprising:
- gamma rays detecting means which includes an array of scintillation detectors said array comprising a plurality of detectors, each detector including:
a) a first scintillator;
b) a second scintillator optically coupled to said first scintillator, said scintillators having different scintillation characteristics; and c) a photodetector optically coupled to said second scintillator, an interaction of a gamma ray in said second scintillator generating a flash of light which is detected by said photodetectors which generates in turn an electric signal, an interaction of a gamma ray in said first scintillator generating a flash of light which is transmitted through said second scintillator and is detected by said photodetector which generates in turn an electric signal, - processing means operatively connected to the photodetectors of the scintillation detectors of said array for processing the signals generated by the photodetectors to provide said information.
- gamma rays detecting means which includes an array of scintillation detectors said array comprising a plurality of detectors, each detector including:
a) a first scintillator;
b) a second scintillator optically coupled to said first scintillator, said scintillators having different scintillation characteristics; and c) a photodetector optically coupled to said second scintillator, an interaction of a gamma ray in said second scintillator generating a flash of light which is detected by said photodetectors which generates in turn an electric signal, an interaction of a gamma ray in said first scintillator generating a flash of light which is transmitted through said second scintillator and is detected by said photodetector which generates in turn an electric signal, - processing means operatively connected to the photodetectors of the scintillation detectors of said array for processing the signals generated by the photodetectors to provide said information.
13. A tomograph as defined in claim 12, wherein the first scintillator, the second scintillator and the photodetector of a scintillation detector are aligned along an axis.
14. A tomograph as defined in claim 12, wherein the scintillation detectors of said array do not interact optically with each other.
15. A tomograph as defined in claim 12, wherein the scintillation detectors of said array surrounds at least partially said human body or said animal emitting radiation.
16. A tomograph as defined in claim 12, wherein the scintillation detector of said array are grouped in plurality of adjacent axially aligned rings.
17. A tomograph as defined in claim 12, wherein the photodetectors are photodiodes.
18. A tomograph as defined in claim 17, wherein said photodiodes are avalanche photodiodes.
19. A scintillation detector for a tomograph for detecting radiation, said detector comprising:
- at least two scintillators having different scintillation charateristics and being optically coupled to each other so as to define a light guide;
and - a photodetector optically coupled to only one of said scintillators, an interaction of a radiation in either one of said scintillators generating a flash of light characteristic of the scintillator in which the interaction has occurred, said flash of light being transmitted by said light guide defined by the scintillators to said photodetector which generates in turn an output signal.
- at least two scintillators having different scintillation charateristics and being optically coupled to each other so as to define a light guide;
and - a photodetector optically coupled to only one of said scintillators, an interaction of a radiation in either one of said scintillators generating a flash of light characteristic of the scintillator in which the interaction has occurred, said flash of light being transmitted by said light guide defined by the scintillators to said photodetector which generates in turn an output signal.
Priority Applications (2)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CA000510983A CA1245375A (en) | 1986-06-06 | 1986-06-06 | Scintillation detector for tomographs |
| US07/058,363 US4843245A (en) | 1986-06-06 | 1987-06-04 | Scintillation detector for tomographs |
Applications Claiming Priority (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CA000510983A CA1245375A (en) | 1986-06-06 | 1986-06-06 | Scintillation detector for tomographs |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CA1245375A true CA1245375A (en) | 1988-11-22 |
Family
ID=4133302
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA000510983A Expired CA1245375A (en) | 1986-06-06 | 1986-06-06 | Scintillation detector for tomographs |
Country Status (2)
| Country | Link |
|---|---|
| US (1) | US4843245A (en) |
| CA (1) | CA1245375A (en) |
Cited By (2)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| EP0819406A1 (en) * | 1996-06-20 | 1998-01-21 | Siemens Aktiengesellschaft | Computer tomography apparatus |
| US7132664B1 (en) | 2002-11-09 | 2006-11-07 | Crosetto Dario B | Method and apparatus for improving PET detectors |
Families Citing this family (20)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| CA1303256C (en) * | 1988-12-14 | 1992-06-09 | Royal Institution For The Advancement Of Learning (The) | Scintillation crystals for positron emission tomography having a non reflecting band |
| US5514870A (en) * | 1994-03-11 | 1996-05-07 | James R. Langenbrunner | Fast CsI-phoswich detector |
| YU29896A (en) * | 1996-05-17 | 1998-11-05 | Vladimir Bošnjaković | Visualising detector for universal nucleo-medical visualising system |
| EP0958508B1 (en) * | 1997-02-10 | 2007-03-28 | THE UNIVERSITY OF ALBERTA, SIMON FRASER UNIVERSITY, THE UNIV. OF VICTORIA,THE UNIV. OF BRITISH COLUMBIA, carrying on as TRIUMF | Segmented scintillation detector for photon interaction coordinates |
| US6194728B1 (en) | 1997-05-05 | 2001-02-27 | Adac Laboratories | Imaging detector for universal nuclear medicine imager |
| US6288399B1 (en) | 1997-11-12 | 2001-09-11 | Cti Pet Systems, Inc. | Depth of interaction detector block for high resolution positron emission tomography |
| US6362479B1 (en) | 1998-03-25 | 2002-03-26 | Cti Pet Systems, Inc. | Scintillation detector array for encoding the energy, position, and time coordinates of gamma ray interactions |
| CA2252993C (en) | 1998-11-06 | 2011-04-19 | Universite De Sherbrooke | Detector assembly for multi-modality scanners |
| US6749761B1 (en) * | 2000-10-10 | 2004-06-15 | Cti Pet Systems, Inc. | Method for producing a high resolution detector array |
| GB2378112A (en) * | 2001-06-26 | 2003-01-29 | Europ Org For Nuclear Research | A PET scanner with LuAP or LuYAP scintillators |
| US7102135B2 (en) * | 2001-06-26 | 2006-09-05 | European Organization For Nuclear Research | PET scanner |
| JP4093013B2 (en) * | 2002-10-23 | 2008-05-28 | 株式会社日立製作所 | Radiation inspection equipment |
| US20050109943A1 (en) * | 2003-09-18 | 2005-05-26 | Vaquero Juan J. | Tomography scanner with axially discontinuous detector array |
| JP3841358B2 (en) * | 2005-02-04 | 2006-11-01 | 株式会社日立製作所 | Radiation inspection apparatus and radiation inspection method |
| US7635848B2 (en) * | 2005-04-01 | 2009-12-22 | San Diego State University Research Foundation | Edge-on SAR scintillator devices and systems for enhanced SPECT, PET, and compton gamma cameras |
| WO2007019708A1 (en) * | 2005-08-18 | 2007-02-22 | Societe De Commercialisation Des Produits De La Recherche Appliquee - Socpra | Digital identification and vector quantization methods and systems for detector crystal recognition in radiation detection machines |
| US7414247B2 (en) * | 2006-07-17 | 2008-08-19 | General Electric Company | Methods and apparatus for geophysical logging |
| ITRM20100082A1 (en) * | 2010-03-02 | 2011-09-03 | Univ Roma La Sapienza | GONIOMETRIC SCANNING PROBE |
| WO2012034220A1 (en) | 2010-09-14 | 2012-03-22 | Abdelmounaime Faouzi Zerrouk | Depth-of-interaction scintillation detectors |
| US8993971B2 (en) | 2011-06-15 | 2015-03-31 | The Board Of Trustees Of The Leland Stanford Junior University | High resolution positron emission tomography |
Family Cites Families (13)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US2910592A (en) * | 1953-09-09 | 1959-10-27 | Texaco Development Corp | Scintillation type detector |
| US3851177A (en) * | 1973-02-23 | 1974-11-26 | Searle & Co | Scintillation camera with improved linearity |
| US3899675A (en) * | 1974-03-04 | 1975-08-12 | Cleon Corp | Whole body imaging |
| US3919556A (en) * | 1974-05-15 | 1975-11-11 | Gen Electric | Gamma camera |
| US3955088A (en) * | 1974-10-02 | 1976-05-04 | G. D. Searle & Co. | Positron imaging device with plural coincidence channels and graded radiation absorption |
| US4095107A (en) * | 1976-04-15 | 1978-06-13 | Sebastian Genna | Transaxial radionuclide emission camera apparatus and method |
| US4037105A (en) * | 1976-06-01 | 1977-07-19 | Laurer Gerard R | Radiation detector with array of different scintillators |
| US4075483A (en) * | 1976-07-12 | 1978-02-21 | Raytheon Company | Multiple masking imaging system |
| US4124804A (en) * | 1976-12-17 | 1978-11-07 | Stuart Mirell | Compton scatter scintillation camera system |
| US4323778A (en) * | 1978-05-31 | 1982-04-06 | Coal Industry (Patents) Limited | Radiation detectors supported by resilient optical coupler |
| US4398092A (en) * | 1979-08-08 | 1983-08-09 | Technicare Corporation | Shaped detector |
| US4675526A (en) * | 1985-04-15 | 1987-06-23 | Rogers Joel G | Method and apparatus for 3-D encoding |
| US4677299A (en) * | 1985-05-13 | 1987-06-30 | Clayton Foundation For Research | Multiple layer positron emission tomography camera |
-
1986
- 1986-06-06 CA CA000510983A patent/CA1245375A/en not_active Expired
-
1987
- 1987-06-04 US US07/058,363 patent/US4843245A/en not_active Expired - Lifetime
Cited By (2)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| EP0819406A1 (en) * | 1996-06-20 | 1998-01-21 | Siemens Aktiengesellschaft | Computer tomography apparatus |
| US7132664B1 (en) | 2002-11-09 | 2006-11-07 | Crosetto Dario B | Method and apparatus for improving PET detectors |
Also Published As
| Publication number | Publication date |
|---|---|
| US4843245A (en) | 1989-06-27 |
Similar Documents
| Publication | Publication Date | Title |
|---|---|---|
| CA1245375A (en) | Scintillation detector for tomographs | |
| US4672207A (en) | Readout system for multi-crystal gamma cameras | |
| CA1244561A (en) | Two dimensional photon counting position encoder system and process | |
| CA2252993C (en) | Detector assembly for multi-modality scanners | |
| EP0500798B1 (en) | High resolution gamma ray detectors for positron emission tomography (pet) and single photon emission computed tomography (spect) | |
| US4750972A (en) | Process for fabricating tuned light guide for photoelectrons | |
| US5144141A (en) | Photodetector scintillator radiation imager | |
| CA1058333A (en) | Transverse section radionuclide scanning system | |
| US5122667A (en) | Means for measuring the depth interaction of gamma-rays in scintillation crystals in order to improve the spatial resolution of positron imaging systems | |
| US5753917A (en) | Dual crystal scintillation camera | |
| US20020195565A1 (en) | PET scanner | |
| GB2198620A (en) | A scintillation detector and a positron ct apparatus incorporating it | |
| US7381956B2 (en) | Detector element for spatially resolved detection of gamma radiation | |
| US6175116B1 (en) | Hybrid collimation and coincidence imager for simultaneous positron and single photon imaging | |
| US20040200966A1 (en) | Gamma-ray detection apparatus and method for positron emission tomography | |
| KR102182318B1 (en) | A device that fuses a dual particle image based on the coded-aperture | |
| US3329814A (en) | Stereo positron camera for determining the spatial distribution of radioactive material in a test body | |
| EP0829022B1 (en) | An apparatus for scintigraphic analysis, particularly a mammograph, with sub-millimetric spatial resolution | |
| JP2002090458A (en) | Ect apparatus | |
| US5015861A (en) | Lead carbonate scintillator materials | |
| JPS61132888A (en) | Positron ct device | |
| US6329657B1 (en) | Coincidence transmission source | |
| EP0181322A1 (en) | Improvement in positron cameras | |
| US3591807A (en) | Photon detector utilizing a well-type scintillation crystal | |
| Burnham et al. | A low-Z PET detector |
Legal Events
| Date | Code | Title | Description |
|---|---|---|---|
| MKEX | Expiry |