WO2010086899A1 - 放射線断層撮影装置 - Google Patents
放射線断層撮影装置 Download PDFInfo
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- WO2010086899A1 WO2010086899A1 PCT/JP2009/000362 JP2009000362W WO2010086899A1 WO 2010086899 A1 WO2010086899 A1 WO 2010086899A1 JP 2009000362 W JP2009000362 W JP 2009000362W WO 2010086899 A1 WO2010086899 A1 WO 2010086899A1
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- 230000005855 radiation Effects 0.000 title claims abstract description 179
- 238000003325 tomography Methods 0.000 title claims abstract description 71
- 238000001514 detection method Methods 0.000 claims abstract description 115
- 238000003860 storage Methods 0.000 claims description 24
- 239000013078 crystal Substances 0.000 abstract description 67
- 230000005251 gamma ray Effects 0.000 abstract description 16
- 238000003745 diagnosis Methods 0.000 abstract description 10
- 238000004364 calculation method Methods 0.000 description 20
- 238000010586 diagram Methods 0.000 description 10
- 238000009513 drug distribution Methods 0.000 description 8
- 230000000694 effects Effects 0.000 description 8
- 238000013500 data storage Methods 0.000 description 7
- 239000002131 composite material Substances 0.000 description 6
- 230000007246 mechanism Effects 0.000 description 6
- 238000003384 imaging method Methods 0.000 description 5
- 238000000034 method Methods 0.000 description 5
- 238000013507 mapping Methods 0.000 description 4
- 229940121896 radiopharmaceutical Drugs 0.000 description 4
- 239000012217 radiopharmaceutical Substances 0.000 description 4
- 230000002799 radiopharmaceutical effect Effects 0.000 description 4
- 230000009471 action Effects 0.000 description 3
- 230000035945 sensitivity Effects 0.000 description 3
- 230000008859 change Effects 0.000 description 2
- 210000000056 organ Anatomy 0.000 description 2
- 238000012545 processing Methods 0.000 description 2
- 238000002603 single-photon emission computed tomography Methods 0.000 description 2
- 230000005540 biological transmission Effects 0.000 description 1
- 238000002591 computed tomography Methods 0.000 description 1
- 238000013480 data collection Methods 0.000 description 1
- 238000009826 distribution Methods 0.000 description 1
- 238000007689 inspection Methods 0.000 description 1
- 238000004519 manufacturing process Methods 0.000 description 1
- 239000000463 material Substances 0.000 description 1
- 230000004048 modification Effects 0.000 description 1
- 238000012986 modification Methods 0.000 description 1
- 238000010248 power generation Methods 0.000 description 1
- 230000004044 response Effects 0.000 description 1
- 239000004065 semiconductor Substances 0.000 description 1
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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)
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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
- G01T1/16—Measuring radiation intensity
- G01T1/161—Applications in the field of nuclear medicine, e.g. in vivo counting
- G01T1/1611—Applications in the field of nuclear medicine, e.g. in vivo counting using both transmission and emission sources sequentially
Definitions
- the present invention relates to a radiation imaging apparatus for imaging radiation emitted from a subject, and more particularly to a radiation imaging apparatus having a wide field of view to the extent that a body portion of a subject can be imaged at once.
- a radiation tomography that obtains a tomographic image of a radiopharmaceutical distribution in a region of interest of a subject by detecting an annihilation radiation pair (for example, ⁇ -rays) released from a radiopharmaceutical that is administered to the subject and localized in the region of interest Used in photographic equipment (ECT: Emission-Computed Tomography).
- ECT mainly includes a PET (Positoron Emission Tomography) apparatus, a SPECT (Single Photon Emission Computed Tomography) apparatus, and the like.
- the PET apparatus has a detector ring in which block-shaped radiation detectors are arranged in a ring shape. This detector ring is provided to surround the subject and is configured to detect the radiation that has passed through the subject.
- the radiation detector deployed in the detector ring of such a PET apparatus has a configuration capable of position discrimination in the depth direction of the scintillator provided in the radiation detector for the purpose of increasing spatial resolution. Is often installed.
- a conventional PET apparatus 50 surrounds an introduction hole with a gantry 51 having an introduction hole for introducing a subject and a block-shaped radiation detector 52 for detecting radiation inside the gantry 51.
- the detector ring 53 formed in such a manner and a support member 54 provided so as to surround the detector ring 53 are provided.
- a bleeder unit 55 having a bleeder circuit is provided at a position where the support member 54 is interposed, and this connects the support member 54 and the radiation detector 52.
- a PET apparatus is described in Patent Document 1, for example.
- the PET device measures the annihilation radiation pair emitted from the radiopharmaceutical. That is, the annihilation radiation pair radiated from the inside of the subject M is a radiation pair whose traveling direction is opposite by 180 °. As shown in FIG. 11, the detector ring 53 has a detection element C for detecting an annihilation radiation pair stacked in the z direction. Thereby, the position of the annihilation radiation pair with respect to the detector ring 53 can be discriminated in the z direction.
- the annihilation radiation pair generated from the annihilation point P inside the subject is incident on two different detection elements C included in the detector ring 53.
- the two detection elements independently send two detection data D1, D2 to the incident time specifying unit 61.
- the time of incidence on the detector ring 53 is specified for each of the detection data D1 and D2.
- the detection data D1 and D2 are output to the simultaneous event certification unit 62, and it is certified whether the incidence of radiation indicated by the detection data D1 and D2 has occurred at the same time.
- the detection data D1 and D2 are paired, and each is determined to be caused by a single event of occurrence of an annihilation radiation pair. Is done.
- the detection data D1 and D2 are sent to the detection intensity specifying unit 63 and the LOR specifying unit 64. In the detection intensity specifying unit 63, the intensity of the incident radiation is calculated from the detection data D1 and D2.
- position information of the detection data D1 and D2 is determined.
- the vector data N associated with the calculated incident time, position information, and detection intensity is stored in the simultaneous event number storage unit 65 and used for generating a tomographic image of the subject.
- the conventional radiation tomography apparatus has the following problems. That is, if the detector ring 53 is long in the z direction, there is a problem that the computation is greatly complicated. In recent years, radiation tomography apparatuses in which the detector ring 53 is wide enough to cover the whole body of the subject are being developed. In such a configuration, the detector ring 53 has a larger number of detection elements than before. Accordingly, the number of combinations of two different detection elements in the detector ring 53 is so large that it cannot be considered in the conventional apparatus. If a tomographic image is to be taken in consideration of the number of simultaneous events for all of these combinations, the calculation load on each unit 61, 62, 63, 64, 65 becomes very high. If this state is left as it is, an expensive arithmetic device is required to acquire a tomographic image of the subject M, and the time required for generating the tomographic image becomes long.
- the present invention has been made in view of such circumstances, and an object thereof is to provide a radiation tomography apparatus capable of reducing the load of calculation of detection data even when the detector ring is wide. There is to do.
- the radiation tomography apparatus is a ring unit configured by arranging unit detection rings in which radiation detection elements for detecting radiation are arranged in an annular shape so as to share their central axes.
- a direct coincidence means for counting the number of simultaneous events, which is the number of times that two different radiation detection elements belonging to the ring unit have simultaneously detected radiation, wherein a plurality of ring units have their central axes
- the detector rings are configured by sharing and are connected to both of the first ring unit and the second ring unit adjacent to each other, and belong to each of the first ring unit and the second ring unit.
- Counts the number of simultaneous events which is the number of times that two radiation detection elements have simultaneously detected radiation. It is characterized in further comprising a counting means.
- the configuration of the present invention has a plurality of ring units. Each ring unit is provided with direct coincidence means. The direct coincidence means counts the number of simultaneous events for the ring unit. In addition to this, the present invention is provided with cross coincidence means connected to both ring units adjacent to each other. The cross coincidence means counts the number of simultaneous events only when two radiation detection elements belonging to the first ring unit and the second ring unit adjacent to each other simultaneously detect radiation.
- the detector ring in the present invention can be configured by connecting ring units having the same configuration as that of conventionally used radiographic apparatuses. Then, even if the detector ring becomes wider and the number of radiation detection elements increases, the simultaneous counting in the radiation tomography apparatus according to the present invention is shared by a plurality of direct simultaneous counting means for each ring unit. Therefore, no matter how many ring units are added, the computation load per one direct coincidence means does not fluctuate.
- a cross coincidence means is provided in the configuration of the present invention. If the ring units are connected, annihilation radiation may be incident on each of the adjacent ring units. According to the present invention, such counting of annihilation radiation pairs is performed by the cross coincidence means. As a result, the number of annihilation radiation pairs used for generating a tomographic image is large.
- the distance in the central axis direction between the radiation detection elements for which simultaneous counting is performed is limited to the thickness of the ring unit in the central axis direction or less. Therefore, the configuration of the present invention does not take into account annihilation radiation pairs that are not suitable for generating a tomographic image, so that a tomographic image suitable for diagnosis is generated while reducing the calculation load.
- the radiation tomography apparatus which can be provided can be provided.
- the above-described detector ring is more desirable if it can be mechanically disassembled in units of ring units.
- the detector ring can be disassembled and transported into a plurality of ring units, and when the radiation tomography apparatus fails, the detector ring can be disassembled for each ring unit. So you can easily check the inside of the detector ring. In addition, the detection ring can be repaired by exchanging the ring unit.
- the above-mentioned cross coincidence means counts the number of simultaneous cross events only when the distance between the two radiation detection elements in the central axis direction is equal to or less than a predetermined length.
- the calculation load of the radiation tomography apparatus can be further reduced. That is, the cross coincidence unit does not perform coincidence counting even if two radiation detection elements that simultaneously detect radiation belong to each of the adjacent ring units unless their distance is equal to or less than a predetermined length. It is. By doing so, it is possible to provide a radiation tomography apparatus in which the calculation load is further reduced.
- predetermined length storage means for storing the predetermined length and an input means for inputting the predetermined length are provided, and the predetermined length can be changed according to the input of the input means.
- a list storage unit that stores a combination list in which combinations of two radiation detection elements are listed, and a count that instructs the cross coincidence unit to execute the count of the number of cross simultaneous events.
- the count instruction means includes a combination list of combinations of the two detection elements that have detected the radiation. It is more desirable to instruct the cross coincidence means to execute counting only when registered in.
- the above-described configuration shows a specific aspect of the tomography apparatus of the present invention. That is, according to the above-described configuration, the cross coincidence unit is instructed to perform counting using a combination list in which combinations of two radiation detection elements are listed. By configuring in this way, the cross coincidence unit can perform simultaneous counting only when the combination of two radiation detecting elements that simultaneously detect radiation is in the combination list. Since the combination list can list only the radiation detection element pairs whose distance is equal to or less than the predetermined length, the distance between the two radiation detection elements in the central axis direction can be easily set to the predetermined length. Or less.
- the above-mentioned radiation tomography apparatus further includes a top plate extending in the central axis direction and inserted into the detector ring.
- the top plate can be rotated around the central axis.
- a radiation source (B) a radiation detection means rotatable around the central axis with respect to the top plate, (C) a support means for supporting the radiation source and the radiation detection means, and (D) a rotation for rotating the support means.
- a radiation tomography apparatus capable of acquiring both the internal structure of the subject and the drug distribution can be provided.
- a PET device can generally obtain information relating to drug distribution.
- a composite image suitable for diagnosis can be generated.
- the burden on the coincidence counting means can be reduced. That is, the configuration of the present invention is provided with direct coincidence means for each ring unit.
- the simultaneous counting in the radiation tomography apparatus according to the present invention is performed by a plurality of direct coincidence means for each ring unit, and no matter how many ring units are added, the calculation per direct coincidence means is performed.
- the load does not fluctuate.
- a cross coincidence means is provided in the configuration of the present invention. As a result, the number of annihilation radiation pairs used for generating a tomographic image is large.
- simultaneous counting is performed unless two radiation detection elements that simultaneously detect radiation are (A) belong to the same ring unit or (B) belong to each of adjacent ring units. It will never be. Therefore, the distance in the central axis direction between the radiation detection elements for which coincidence is performed is limited to the thickness of the ring unit in the central axis direction or less. Therefore, the configuration of the present invention does not take into account annihilation radiation pairs that are not suitable for generating a tomographic image, so that a tomographic image suitable for diagnosis is generated while reducing the calculation load.
- the radiation tomography apparatus which can be provided can be provided.
- FIG. 1 is a functional block diagram illustrating a configuration of a radiation tomography apparatus according to Embodiment 1.
- FIG. 1 is a perspective view illustrating a configuration of a radiation detector according to Embodiment 1.
- FIG. It is a figure explaining the structure of the detector ring which concerns on Example 1.
- FIG. It is a figure explaining the structure of the ring unit which concerns on Example 1.
- FIG. FIG. 6 is a schematic diagram for explaining the operation of the filter according to the first embodiment.
- FIG. 6 is a schematic diagram for explaining the operation of the filter according to the first embodiment.
- 6 is a functional block diagram illustrating a configuration of a radiation tomography apparatus according to Embodiment 2.
- FIG. 10 is a schematic diagram illustrating the operation of a filter according to a second embodiment.
- FIG. 10 is a schematic diagram illustrating the operation of a filter according to a second embodiment.
- FIG. 9 is a functional block diagram illustrating a configuration of a PET / CT apparatus according to a third embodiment. It is a figure explaining the structure of the PET apparatus of a conventional structure. It is a functional block diagram explaining the structure of the PET apparatus of a conventional structure.
- the present invention is applied to a PET apparatus, and in the third embodiment, the present invention is applied to a PET / CT apparatus.
- FIG. 1 is a functional block diagram illustrating the configuration of the radiation tomography apparatus according to the first embodiment.
- the radiation tomography apparatus 9 according to the first embodiment includes a top plate 10 that lies on the subject M and a gantry 11 having a through hole that surrounds the subject M.
- the top plate 10 is provided so as to pass through the opening of the gantry 11 and is movable back and forth along the direction in which the opening of the gantry 11 extends (z direction).
- Such sliding of the top plate 10 is realized by the top plate moving mechanism 15.
- the top plate moving mechanism 15 is controlled by the top plate movement control unit 16.
- a detector ring 12 for detecting an annihilation gamma ray pair emitted from the subject M is provided inside the gantry 11, a detector ring 12 for detecting an annihilation gamma ray pair emitted from the subject M is provided.
- the detector ring 12 has a cylindrical shape extending in the body axis direction z (corresponding to the extending direction of the present invention) of the subject M, and its length is 1 m to 1.8 m. That is, the detector ring 12 extends to such an extent that at least the body portion of the subject M can be completely covered.
- the detector ring 12 is configured by arranging block-shaped radiation detectors 1 in a ring shape. Assuming that the width per one in the radiation detector 1 is about 5 cm, about 20 to 36 radiation detectors 1 are arranged on the detector ring 12 in the z direction. The configuration of the radiation detector 1 will be briefly described.
- FIG. 2 is a perspective view illustrating the configuration of the radiation detector according to the first embodiment. As shown in FIG. 2, the radiation detector 1 includes a scintillator 2 that converts radiation into fluorescence, and a photodetector 3 that detects fluorescence. A light guide 4 for transmitting and receiving fluorescence is provided at a position where the scintillator 2 and the photodetector 3 are interposed.
- the scintillator 2 is configured by scintillator crystals arranged three-dimensionally.
- the scintillator crystal is composed of Lu 2 (1-X) Y 2X SiO 5 (hereinafter referred to as LYSO ) in which Ce is diffused.
- the photodetector 3 can specify the fluorescence generation position indicating which scintillator crystal emits fluorescence, and can also specify the intensity of fluorescence and the time when the fluorescence is generated. it can.
- the scintillator crystal corresponds to the radiation detection element of the present invention.
- FIG. 3 is a diagram illustrating the configuration of the detector ring according to the first embodiment.
- the radiation detectors 1 are arranged along a virtual circle (more precisely, a regular n-gon) in the detector ring 12.
- the scintillator crystals are also arranged along a virtual circle (precisely, a regular n-gon), and constitute a unit detection ring 12b as shown in FIG.
- the unit detection ring 12b is formed of scintillator crystals C (radiation detection elements) arranged at the same position in the z direction and arranged along an annular ring.
- the unit detection ring 12b is a concept in which scintillator crystals are arranged in a line and is independent of the radiation detectors 1 arranged along a virtual circle. As shown in FIG. 3B, the unit detection ring 12b is connected in the z direction to form the detector ring 12. In other words, the unit detection rings 12b are connected to share a central axis along the z direction.
- the unit detection ring 12b has a through-hole at the center, and the unit detection ring 12b is arranged so that the through-holes of the unit detection ring 12b are connected, and the detector ring 12 is configured. You can also.
- the detector ring 12 is formed by arranging around 100 radiation detectors 1 in an annular shape. Therefore, when the through hole 12a is viewed from the z direction, the through hole 12a is, for example, It is a regular 100-gon. In this case, the plurality of unit detection rings 12b are connected so as to share the respective central axes, and the through hole 12a has a shape of a 100 prism.
- the detector ring 12 is configured by connecting each of a plurality of ring units 121, 122, and 123 as shown in FIG.
- This ring unit is configured by connecting the above-described unit detection rings 12b in the z direction.
- the ring unit 121 includes about 12 radiation detectors 1 arranged in the z direction. In other words, the ring units are connected to share a central axis along the z direction.
- the detector ring 12 is configured by mechanically connecting ring units 121, 122, and 123 manufactured separately. Therefore, the detector ring 12 can be mechanically disassembled in units of ring units.
- the detector ring 12 is constituted by three ring units 121, 122, and 123, but the number of ring units constituting the detector ring 12 is not limited to this.
- the radiation tomography apparatus 9 is further provided with various units for acquiring a tomographic image of the subject M as shown in FIG.
- the radiation tomography apparatus 9 includes a C coincidence unit 20c and a D coincidence unit 20d that perform coincidence counting of annihilation ⁇ rays based on detection data detected by the detector ring 12, and detection data.
- a data storage unit 17 that stores data
- a mapping unit 18 that forms a tomographic image of the subject M
- a calibration unit 19 that applies calibration to the tomographic image of the subject M are provided.
- the calibration unit 19 is provided for the purpose of removing a false image reflected in the tomographic image, and superimposes predetermined calibration data on the tomographic image of the subject M.
- the MRD storage unit 37 stores MRD described later.
- the C coincidence unit corresponds to the cross coincidence unit of the present invention, and the D coincidence unit corresponds to the direct coincidence unit of the present invention.
- the MRD storage unit corresponds to the predetermined length storage means of the present invention.
- the radiation tomography apparatus 9 includes a main control unit 35 that performs overall control of each unit and a display unit 36 that displays a radiation tomographic image.
- the main control unit 35 is constituted by a CPU, and realizes the mapping unit 18 and the calibration unit 19 by executing various programs.
- each above-mentioned part may be divided
- the input unit 38 is for inputting a surgeon's instruction, and for example, accepts a change in the MRD setting by the surgeon.
- the input unit corresponds to the input unit of the present invention.
- the C coincidence unit 20c includes a C filter 21c
- the D coincidence unit 20d includes a D filter 21d.
- the D coincidence unit 20d is provided in each of the ring units 121, 122, and 123.
- two scintillator crystals belonging to each of the ring units in charge detect an annihilation ⁇ -ray pair.
- the C coincidence counting unit 20c is connected to the ring units 121 and 122 adjacent to each other.
- the C coincidence unit 20c counts the number of simultaneous events when two scintillator crystals belonging to different ring units detect an annihilation gamma ray pair.
- the C filter 21c is provided across adjacent ring units, and another C filter 21c is also provided in the ring units 122 and 123 adjacent to each other.
- the number of coincidence counters is as follows. That is, if the number of ring units constituting the detector ring 12 is n, the number of D coincidence units 20d provided in the radiation tomography apparatus 9 is n, and the number of C coincidence units 20c is n ⁇ . 1.
- the clock 23 outputs time information to the D filter 21d and the C filter 21c.
- the clock 23 is shown as if it is connected only to a single D filter 21 d, but in reality, the clock 23 is connected to all the filters provided in the radiation tomography apparatus 9. Will be.
- an LOR specifying unit 22 and a fluorescence intensity calculating unit 24 are provided downstream of the D filter 21d.
- the C coincidence unit 20c is also provided with an LOR specifying unit 22 and a fluorescence intensity calculation unit 24 downstream of the C filter 21c.
- the output of the fluorescence intensity calculation unit 24 is output to the data storage unit 17.
- the data output from the fluorescence intensity calculation unit 24 is data in which LOR, detection time, and detection intensity are related.
- the LOR specifying unit 22 and the fluorescence intensity calculating unit 24 are not shown for a part of the coincidence counting unit for the sake of brevity.
- the data processing in the radiation tomography apparatus 9 will be briefly described. Assume that a pair of annihilation ⁇ rays are incident on two different locations of the ring unit 121 of the detector ring 12 from the vanishing point P in FIG. Then, the ring unit 121 outputs two detection signals to the D filter 21d provided therein, and recognizes the fact that two annihilation ⁇ rays are incident on the ring unit 121 at the same time. This determination of simultaneity is performed using the clock 23. That is, time data is added to the detection signal output from the clock 23 to the detection signal.
- the D filter 21d recognizes that the two detection signals are paired and are due to the annihilation gamma ray pair.
- the two detection signals recognized as annihilation gamma rays pass through the D filter 21 d and are sent to the LOR specifying unit 22.
- LOR is an abbreviation for Line of response, and is a line segment connecting two scintillator crystals on which annihilation ⁇ rays are incident. This LOR can also be expressed as position information in annihilation ⁇ -rays.
- the fluorescence intensity calculation unit 24 specifies the fluorescence intensity of each scintillator crystal.
- the D filter 21d allows each detection signal to pass only when two ⁇ rays are simultaneously incident on one of the ring units that it is responsible for, and discards the detection data of the ⁇ rays that could not be paired. To do.
- the role of the D filter 21d will be described more specifically. As shown in FIG. 5A, it is assumed that ten scintillator crystals C are arranged in the z direction in the ring unit 121. Consider the scintillator crystal Cr belonging to the ring unit 121. Since ten unit detection rings 12b are arranged in the ring unit 121, among the LORs of the scintillator crystal Cr, the LORs through which the D filter 21d passes detection signals are LOR1d to LOR10d.
- the D filter 21d discards the detection signal output from the scintillator crystal Cr as long as there is no scintillator crystal that can be paired in the ring unit 121.
- the role of the C filter 21c is similar to that of the D filter 21d. However, unlike the D filter 21d, the C filter 21c is included in the scintillator crystal Cr and the LOR through which the C filter 21c passes the detection signal is either the scintillator crystal Cr or one of the ring units 122a and 122b adjacent to the ring unit 121. There are 20 types of LOR1c to LOR20c between the scintillator crystals belonging to (see FIG. 5B).
- the C filter 21c provided between the ring units 121 and 122a has a scintillator crystal as long as there is no scintillator crystal that can be paired inside the ring unit 122a even when the scintillator crystal Cr senses ⁇ rays. Discard the detection signal for Cr.
- the C filter 21c provided between the ring units 121 and 122b has a scintillator as long as there is no scintillator crystal that can be paired inside the ring unit 122b even if the scintillator crystal Cr senses ⁇ rays. The detection signal related to the crystal Cr is discarded.
- the LOR through which the C filter 21c provided across the ring units 122a and 121 passes the detection signal is LOR1c to LOR10c, and the C filter 21c provided across the ring units 121 and 122b detects the detection signal.
- LORs that pass are LOR11c to LOR20c.
- the ring unit 121 corresponds to the first ring unit of the present invention
- the ring units 122a and 122b correspond to the second ring unit of the present invention.
- LOR1d to LOR10d there are 30 types of LOR used for coincidence counting: LOR1d to LOR10d, and LOR1c to LOR20c.
- the data storage unit 17 in FIG. 1 stores how often an annihilation ⁇ -ray pair is detected for each LOR.
- the LOR is selected only by the action of the C filter 21c and the D filter 21d and is suitable for acquisition of radiation tomographic images. Thereby, the burden concerning the LOR specific
- the C coincidence counting unit 20c and the D coincidence counting unit 20d sequentially transmit detection data to the data storage unit 17.
- the data storage unit 17 stores the number of times the coincidence counting unit outputs detection data, and records the number of annihilation gamma ray pairs (number of simultaneous events) for each LOR. . That is, the data storage unit 17 stores LOR and the number of counts of annihilation ⁇ -ray pairs in association with each other.
- the mapping unit 18 assembles this related data and acquires a tomographic image (PET image).
- PET image tomographic image of the subject M generated in this way is output to the calibration unit 19 and data processing for removing the false image superimposed on the tomographic image of the subject M is performed.
- the completed image obtained in this way is displayed on the display unit 36. With this, the inspection using the radiation tomography apparatus 9 according to the configuration of the first embodiment is completed.
- the configuration of the first embodiment includes a plurality of ring units.
- Each of the ring units is provided with a D coincidence counting unit 20d.
- the D coincidence counting unit 20d counts the number of simultaneous events for the ring unit.
- a C coincidence unit 20c connected to both of the ring units adjacent to each other is provided.
- the C coincidence counter 20c counts the number of simultaneous events only when two scintillator crystals belonging to each of the adjacent ring units detect radiation simultaneously.
- the D coincidence counting unit 20d for each ring unit, the calculation does not become complicated even if the width of the detector ring 12 is long. That is, the ring unit has the same configuration as that of a conventionally used radiographic apparatus, and they are connected. Then, even if there are many scintillator crystals provided in the detector ring 12, the simultaneous counting in the radiation tomography apparatus 9 according to the first embodiment is shared by the plurality of D coincidence units 20d for each ring unit. Therefore, no matter how many ring units are added, the calculation load per D coincidence unit 20d does not change.
- the configuration of the first embodiment is provided with a C coincidence unit 20c. If the ring units are connected, annihilation radiation may be incident on each of the adjacent ring units. According to the first embodiment, the C coincidence counting unit 20c performs such counting of annihilation ⁇ -ray pairs. As a result, the number of annihilation ⁇ -ray pairs used for generating a tomographic image becomes large.
- the distance in the z direction between the scintillator crystals is only the thickness of the ring unit in the z direction.
- the configuration of the first embodiment does not take into account annihilation ⁇ -ray pairs that are not suitable for generating a tomographic image, so that a tomography suitable for diagnosis while reducing the calculation load.
- a radiation tomography apparatus 9 capable of generating an image can be provided.
- the above-described detector ring 12 can be mechanically disassembled in units of ring units.
- the detector ring 12 can be disassembled and transported into a plurality of ring units, and when the radiation tomography apparatus 9 breaks down, the detector ring 12 is provided for each ring unit. Since it can be disassembled, the inside of the detector ring 12 can be easily inspected. In addition, the detection ring can be repaired by exchanging the ring unit.
- each of the C filters 21c is different in that a C list reference unit 25c and a C list storage unit 26c are provided.
- each D filter 21d is provided with a D list reference unit 25d and a D list storage unit 26d.
- the C list storage unit corresponds to the list storage unit of the present invention
- the C list reference unit corresponds to the counting instruction unit of the present invention.
- the configuration of the second embodiment can discard the detection signal more than the first embodiment. That is, the detection signal discarded in the first embodiment is discarded as it is. In addition to this, the configuration of the second embodiment additionally discards the detection signal according to the list stored in the list storage unit.
- FIG. 8 is a schematic diagram for explaining how the detection signal of the radiation tomography apparatus according to the configuration of the second embodiment is discarded.
- a region R in FIG. 8 is a region having a width corresponding to ten scintillator crystals before and after in the z direction from the scintillator crystal Cr.
- the region R includes (A) a unit detection ring 12b to which the scintillator crystal Cr belongs, (B) ten unit detection rings 12b positioned in front of the scintillator crystal Cr in the z direction, and (C) in the z direction.
- MRD the maximum number of different detector rings with the number of detector rings considering LOR.
- MRD 10. Therefore, when the number of unit detection rings considering LOR is 1, MRD is 0.
- the radiation tomography apparatus 9 according to the second embodiment can read the MRD stored in the MRD storage unit 37.
- Several types of lists are stored in the C list storage unit 26c and the D list storage unit 26d, and a list suitable for reference is sent to the C list reference unit 25c and the D list reference unit 25d according to the MRD. .
- the MRD can be changed, and the operator can adjust the width of the region R by resetting the MRD stored in the MRD storage unit 37 through the input unit 38. It has become.
- the C filter 21c discards the detection signal when the LOR is not within the range of the region R in FIG. That is, in the first embodiment, the LOR between the ring unit 122a and the scintillator crystal Cr is 10 types of LOR1c to LOR10c as shown in FIG. 5, of which LOR1c to LOR8c are related. Since the detection signal is an LOR that protrudes from the region R, it is discarded.
- the LOR drawn between the ring unit 122b and the scintillator crystal Cr was 10 types of LOR11c to LOR20c. Among them, detection signals related to the LOR20c protruded from the region R. Since it is LOR, it is discarded.
- the LOR used for the coincidence counting is 21 types (2 ⁇ MRD + 1 types) of LOR1d to LOR10d (see FIG. 5A) and LOR9c to LOR19c (see FIG. 8).
- the data storage unit 17 in FIG. 1 stores how often an annihilation ⁇ -ray pair is detected for each LOR.
- the selection of the LOR corresponding to the MRD is performed by the C list reference unit 25c. That is, the C list reference unit 25c refers to the C list stored in the C list storage unit 26c and instructs the C filter 21c to discard the detection signals related to LOR1c to LOR8c and LOR20c.
- the C list combinations of scintillator crystals corresponding to a predetermined MRD are listed. Specifically, when the MRD is 10, the C list lists scintillator crystal combinations related to LOR 9c to LOR 19c for the scintillator crystal Cr. The scintillator crystal combinations for LOR1c to LOR8c and LOR20c are not listed.
- the C filter 21c additionally discards the detection data described above in accordance with an instruction from the C list reference unit 25c.
- the configurations of the D list reference unit 25d, the D list storage unit 26d, and the D list stored in the D list reference unit 25d are the same as those of the C list reference unit 25c, the C list storage unit 26c, and the C list. To do. In the ring unit 121 to which the scintillator crystal Cr belongs, the detection data considering MRD is discarded.
- the list generation unit 27 generates a C list and a D list for each MRD.
- the D list is generated by listing, from a combination of two scintillator crystals belonging to a single ring unit, only those whose distances in the z direction of the two scintillator crystals are within the range specified by the MRD.
- the C list is generated by listing, from a combination of two scintillator crystals belonging to each of the ring units adjacent to each other, only those whose z-direction distance between the two scintillator crystals is within the range specified by the MRD. .
- the list generation unit corresponds to the list generation means of the present invention.
- the calculation load of the radiation tomography apparatus 9 can be further reduced. That is, even if two scintillator crystals that simultaneously detect radiation belong to each of adjacent ring units, the C coincidence unit 20c performs coincidence if the distance between them is not less than the length specified by the MRD. It is not done. By doing so, it is possible to provide the radiation tomography apparatus 9 in which the calculation load is further reduced. That is, using the C list that lists the combinations of two scintillator crystals, the C coincidence counting unit 20c is instructed to perform counting. With this configuration, the C coincidence unit 20c can perform simultaneous counting only when the combination of two scintillator crystals that simultaneously detect radiation is in the C list.
- the distance between the two scintillator crystals in the z direction can be easily specified by the MRD. It can be less than or equal to the length.
- the PET / CT apparatus has a configuration including the radiation tomography apparatus (PET apparatus) 9 described in the first and second embodiments and a CT apparatus that generates a tomographic image using X-rays.
- PET apparatus radiation tomography apparatus
- This is a medical device that can generate a composite image obtained by superimposing the tomographic images.
- the configuration of the PET / CT apparatus according to the third embodiment will be described.
- the radiation tomography apparatus (PET apparatus) 9 described in the first or second embodiment can be used. Therefore, a CT apparatus which is a characteristic part in the third embodiment will be described.
- the CT apparatus 8 has a gantry 45.
- the gantry 45 is provided with an opening extending in the z direction, and the top plate 10 is inserted into the opening.
- An gantry 45 supports an X-ray tube 43 that irradiates the subject with X-rays, an FPD (flat panel detector) 44 that has passed through the subject, and the X-ray tube 43 and the FPD 44.
- a support 47 is provided.
- the support 47 has a ring shape and is rotatable around the z axis.
- the rotation of the support 47 is performed by a rotation mechanism 39 including a power generation unit such as a motor and a power transmission unit such as a gear.
- the rotation control unit 40 controls the rotation mechanism 39.
- the X-ray tube corresponds to the radiation source of the present invention.
- the FPD corresponds to the radiation detection means of the present invention, and the support corresponds to the support means of the present invention.
- the rotation mechanism corresponds to the rotation means of the present invention, and the rotation control unit corresponds to the rotation control means of the present invention.
- the CT image generation unit 41 generates an X-ray tomographic image of the subject M based on the X-ray detection data output from the FPD 44.
- the superimposing unit 42 generates a superposition image by superimposing the PET image indicating the drug distribution in the subject output from the radiation tomography apparatus (PET apparatus) 9 and the above-described X-ray tomographic image. It has a configuration.
- the main control unit 35 executes various programs, so that, in addition to the mapping unit 18 and the calibration unit 19 according to the first and second embodiments, the rotation control unit 40, the CT image generation unit 41, the superposition unit 42, And the X-ray tube control unit 46 is realized.
- each above-mentioned part may be divided
- a method for obtaining a fluoroscopic image will be described.
- the X-ray tube 43 and the FPD 44 rotate around the z axis while maintaining their relative positions.
- the X-ray tube 43 intermittently irradiates the subject M with X-rays, and each time the CT image generation unit 41 generates an X-ray fluoroscopic image.
- the plurality of fluoroscopic images are assembled into a single tomographic image using the existing back projection method in the CT image generation unit 41, for example.
- a method for generating a composite image In order to acquire a composite image with the PET / CT apparatus, a region of interest of the subject M is introduced into the CT apparatus, and an X-ray tomographic image thereof is acquired. In addition to this, a region of interest of the subject M is introduced into a radiation tomography apparatus (PET apparatus) 9 to acquire a PET image. Both images are superimposed by the overlapping unit 42, and the completed composite image is displayed on the display unit 36. Thereby, since the drug distribution and the internal structure of the subject can be recognized simultaneously, a tomographic image suitable for diagnosis can be provided.
- PET apparatus radiation tomography apparatus
- the radiation tomography apparatus 9 that can acquire both the internal structure of the subject M and the drug distribution can be provided.
- a PET device can generally obtain information relating to drug distribution.
- a composite image suitable for diagnosis can be generated by superimposing both images.
- the present invention is not limited to the above-described configuration, and can be modified as follows.
- the scintillator crystal referred to in each of the above embodiments is composed of LYSO.
- the scintillator crystal is composed of other materials such as GSO (Gd 2 SiO 5 ) instead. Also good. According to this modification, it is possible to provide a method of manufacturing a radiation detector that can provide a cheaper radiation detector.
- the fluorescence detector is composed of a photomultiplier tube, but the present invention is not limited to this. Instead of the photomultiplier tube, a photodiode, an avalanche photodiode, a semiconductor detector, or the like may be used.
- the top plate is slidable.
- the present invention is not limited to this.
- the top plate may be fixed and the gantry 11 may slide.
- the present invention is suitable for a medical radiation tomography apparatus.
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Abstract
Description
すなわち、検出器リング53をz方向に長いものとすると、飛躍的に演算が煩雑化するという問題点がある。近年において、被検体の全身を覆う程度に検出器リング53が幅広となっている放射線断層撮影装置が開発されつつある。この様な構成においては、検出器リング53には、従来より多数の検出素子が搭載される。したがって、検出器リング53における異なる2つの検出素子の組み合わせは、従来装置では考えられないほど多いものとなる。これらの組み合わせの全てをについて同時イベント数を考慮して断層画像を撮影しようとすると、各部61,62,63,64,65にかかる演算の負荷は、非常に高いものとなる。この状態を放置すれば、被検体Mの断層画像を取得するのに高価な演算装置が必要であるとともに、断層画像の生成に必要な時間は、長いものとなる。
すなわち、本発明に係る放射線断層撮影装置は、放射線を検出する放射線検出素子が環状に配列された単位検出リングの各々が、それらの中心軸を共有して配列されることにより構成されたリングユニットと、リングユニットに属する2つの異なる放射線検出素子が同時に放射線を検出した回数である同時イベント数を計数するダイレクト同時計数手段とを備えた放射線断層撮影装置において、複数のリングユニットがそれらの中心軸を共有して配列されることにより検出器リングが構成され、互いに隣接する第1リングユニット、第2リングユニットの両方に接続されるとともに、第1リングユニット、第2リングユニットのそれぞれに属する2つの放射線検出素子が同時に放射線を検出した回数である同時イベント数を計数するクロス同時計数手段とを備えることを特徴とするものである。
9 放射線断層撮影装置
10 天板
12b 単位検出リング
20c C同時計数部(クロス同時計数手段)
20d D同時計数部(ダイレクト同時計数手段)
25c Cリスト参照部(計数指示手段)
26c Cリスト記憶部(リスト記憶手段)
27 リスト生成部(リスト生成手段)
37 MRD記憶部(所定長記憶手段)
38 入力部(入力手段)
39 回転機構(回転手段)
40 回転制御部(回転制御手段)
43 X線管(放射線源)
44 FPD(放射線検出手段)
47 支持体(支持手段)
121 リングユニット(第1リングユニット)
122a,122b リングユニット(第2リングユニット)
以下、本発明に係る放射線断層撮影装置の各実施例を図面を参照しながら説明する。図1は、実施例1に係る放射線断層撮影装置の構成を説明する機能ブロック図である。実施例1に係る放射線断層撮影装置9は、図1に示すように、被検体Mを仰臥させる天板10と、被検体Mを包囲する貫通穴を有するガントリ11を有している。天板10は、ガントリ11の開口を貫通するように備えられているとともに、ガントリ11の開口の伸びる方向(z方向)に沿って進退自在となっている。この様な天板10の摺動は、天板移動機構15によって実現される。天板移動機構15は、天板移動制御部16によって制御される。
次に、本発明における最も特徴的な構成である、C同時計数部20c,およびD同時計数部20dの構成について説明する。図4に示すように、実施例1に係るC同時計数部20cは、Cフィルタ21cを備え、D同時計数部20dは、Dフィルタ21dを備えている。D同時計数部20dは、各リングユニット121,122,123の各々に設けられており、D同時計数部20dは、各々が担当するリングユニットに属する2つのシンチレータ結晶が消滅γ線対を検出したとき、同時イベント数の計数を行うものである。一方、C同時計数部20cは、互いに隣接するリングユニット121,122に接続されている。このC同時計数部20cは、異なるリングユニットに属する2つのシンチレータ結晶が消滅γ線対を検出したとき、同時イベント数の計数を行うものである。Cフィルタ21cは、隣接するリングユニットに跨って設けられるものであり、別のCフィルタ21cが互いに隣接するリングユニット122,123にも設けられている。
次に、実施例1に係る放射線断層撮影装置9の動作について説明する。実施例1に係る放射線断層撮影装置9で検査を行うには、まず、放射線薬剤を予め注射投薬された被検体Mを天板10に載置する。そして、天板10を摺動させ、被検体Mをガントリ11の開口に導入する。この時点より、被検体Mから発せられる消滅γ線対が検出される。このとき被検体Mにおける撮影部位は、全てガントリ11の内部に納まる格好となり、放射線検出の期間中、天板10は移動しない。
Claims (7)
- 放射線を検出する放射線検出素子が環状に配列された単位検出リングの各々が、それらの中心軸を共有して配列されることにより構成されたリングユニットと、
前記リングユニットに属する2つの異なる放射線検出素子が同時に放射線を検出した回数である同時イベント数を計数するダイレクト同時計数手段とを備えた放射線断層撮影装置において、
複数のリングユニットがそれらの中心軸を共有して配列されることにより検出器リングが構成され、
互いに隣接する第1リングユニット、第2リングユニットの両方に接続されるとともに、前記第1リングユニット、前記第2リングユニットのそれぞれに属する2つの放射線検出素子が同時に放射線を検出した回数である同時イベント数を計数するクロス同時計数手段とを備えることを特徴とする放射線断層撮影装置。 - 請求項1に記載の放射線断層撮影装置において、
前記検出器リングは、リングユニット単位で機械的に分解可能となっていることを特徴とする放射線断層撮影装置。 - 請求項1または請求項2に記載の放射線断層撮影装置において、
前記クロス同時計数手段は、前記中心軸方向における2つの放射線検出素子間の距離が所定の長さ以下となっているときのみクロス同時イベント数を計数することを特徴とする放射線断層撮影装置。 - 請求項3に記載の放射線断層撮影装置において、
前記所定長さを記憶する所定長記憶手段と、
前記所定長さを入力させる入力手段とを備え、
前記所定長さは、前記入力手段の入力にしたがって、変更可能となっていることを特徴とする放射線断層撮影装置。 - 請求項3または請求項4に記載の放射線断層撮影装置において、
2つの放射線検出素子の組み合わせをリスト化した組み合わせリストを記憶するリスト記憶手段と、
前記クロス同時計数手段に対してクロス同時イベント数の計数の実行を指示する計数指示手段とを備え、
前記第1リングユニット、前記第2リングユニットのそれぞれに属する2つの放射線検出素子が同時に放射線を検出したとき、前記計数指示手段は、放射線を検出した2つの検出素子の組み合わせが前記組み合わせリストに登録されている場合のみ、前記クロス同時計数手段に対して計数の実行を指示することを特徴とする放射線断層撮影装置。 - 請求項5に記載の放射線断層撮影装置において、
前記所定長さを基に前記組み合わせリストを生成するリスト生成手段を備えることを特徴とする放射線断層撮影装置。 - 請求項1ないし請求項6のいずれかに記載の放射線断層撮影装置において、
前記中心軸方向に伸びるとともに前記検出器リングの内部に挿入された天板を更に備えるとともに、これに加えて、
(A)前記天板に対し前記中心軸周りに回転可能な放射線源と、
(B)前記天板に対し前記中心軸周りに回転可能な放射線検出手段と、
(C)前記放射線源と前記放射線検出手段とを支持する支持手段と、
(D)前記支持手段を回転させる回転手段と、
(E)前記回転手段を制御する回転制御手段を備えた画像生成装置を更に備えることを特徴とする放射線断層撮影装置。
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PCT/JP2009/000362 WO2010086899A1 (ja) | 2009-01-30 | 2009-01-30 | 放射線断層撮影装置 |
EP09839094.1A EP2392946B1 (en) | 2009-01-30 | 2009-01-30 | Radiation tomography apparatus |
JP2010548254A JP4983984B2 (ja) | 2009-01-30 | 2009-01-30 | 放射線断層撮影装置 |
US13/145,412 US8519341B2 (en) | 2009-01-30 | 2009-01-30 | Radiation tomography apparatus |
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