JP2007101191A - Radiation detector - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation detector for dispensing with a scepter made of a heavy metal, simultaneously performing two-dimensional collection and three-dimensional collection, shortening measurement time, reducing burdens on both an inspected body and a radiological technician, and further, reducing an (RI) cost burden. <P>SOLUTION: A scintillator is divided into a plurality of levels in the depth direction. A scintillator 58 on the highest level close to the inspected body side is disposed on an upper part of a join of a next highest-level scintillator 51 to fulfil roles of a detector and of the scepter 71. In this event, the scintillators, having two or more kinds of different decay times, determine capturing places for γ-photons entering the scintillators by means of either or both of discrimination based on the differences in decay time and sharing of light based on a reflector structure, and therefore, it is possible to simultaneously perform two-dimensional collection and three-dimensional collection. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

The present invention detects a radiation (gamma ray) emitted from a radioisotope (RI) administered to a subject and accumulated in a site of interest, and obtains a tomographic image of the RI distribution of the site of interest, for example, PET The present invention relates to a radiation detector used in a medical diagnostic apparatus such as a (Positron Emission Tomography) apparatus or a SPECT (Single Photo Emission Computed Tomography) apparatus.

This type of radiation detector includes a scintillator that emits light upon incidence of gamma rays emitted from a subject, and a photomultiplier tube that converts light emitted from the scintillator into a pulsed electric signal. As for such a radiation detector, there has conventionally been a one-to-one correspondence between a scintillator and a photomultiplier tube. Are employed to determine the incident position of the gamma rays from the output ratio of these photomultiplier tubes to improve the resolution (see, for example, Patent Document 1).

FIG. 8 is a cross-sectional view (front view) in the X direction when the conventional radiation detector 10 is viewed from the Y direction. In the case of an isotropic voxel detector, a cross-sectional view (side view) in the Y direction when the radiation detector 10 is viewed from the X direction has the same shape as FIG. The radiation detector 10 is partitioned by appropriately sandwiching the light reflecting material 13, and a scintillator group 12 in which a total of 36 scintillators 11 of 6 in the X direction and 6 in the Y direction are arranged in close contact with each other, A light guide 14 that is optically coupled to the scintillator group 12 and is combined with a light reflecting material 15 and has a plurality of subsections defined therein and a light guide 14 that is optically coupled to the light guide 14 The photomultiplier tubes 101, 102, 103, and 104 are configured. In FIG. 8, only the photomultiplier tube 101 and the photomultiplier tube 102 are shown. Here, as the scintillator 11, for example, Bi ₄ Ge ₃ 0 ₁₂ (BGO), Gd ₂ SiO ₅ : Ce (GSO), Lu ₂ SiO ₅ : Ce (LSO), LuYSiO ₅ : Ce (LYSO), NaI, BaF ₂ , CsF, inorganic crystals such as PbWO ₄ used.

When gamma rays are incident on any one of the six scintillators 11 arranged in the X direction, the light is converted into visible light. This light is guided to the photomultiplier tubes 101 to 104 through the optically coupled light guide 14, and at this time, the photomultiplier tubes 101 (103) and the photomultiplier tubes 102 (104) arranged in the X direction. The position, length, and angle of each light reflecting material 15 in the light guide 14 are adjusted so that the output ratio of () changes at a constant rate.

More specifically, if the output of the photomultiplier tube 101 is P1, the output of the photomultiplier tube 102 is P2, the output of the photomultiplier tube 103 is P3, and the output of the photomultiplier tube 104 is P4, the X direction The position and length of the light reflecting material 14 are set so that the calculated value {(P1 + P3) − (P2 + P4)} / (P1 + P2 + P3 + P4) representing the position changes at a constant rate according to the position of each scintillator 11.

On the other hand, in the case of six scintillators arranged in the Y direction, light is similarly guided to the photomultiplier tubes 101 to 104 through the optically coupled light guide 14. That is, the position of each light reflecting material 15 in the light guide 14 so that the output ratio of the photomultiplier tubes 101 (102) and the photomultiplier tubes 103 (104) arranged in the Y direction changes at a constant rate. The length is set, and in the case of tilt, the angle is adjusted.

That is, the position and length of the light reflecting material 15 are set so that the calculated value {(P1 + P2) − (P3 + P4)} / (P1 + P2 + P3 + P4) representing the position in the Y direction changes at a constant rate according to the position of each scintillator. ing.

Here, the light reflecting material 13 between the scintillators 11 and the light reflecting material 15 of the light guide 14 are mainly a multilayer film of silicon oxide and titanium oxide mainly using a polyester film, and its reflection efficiency is very high. Although it is used as a light reflecting element because it is high, strictly speaking, a transmission component is generated depending on the incident angle of light, and the shape and arrangement of the light reflecting material 13 and the light reflecting material 15 are also included in the calculation. It has been decided.

  The scintillator group 12 is bonded to the light guide 14 with a coupling adhesive to form a coupling adhesive layer 16, and the light guide 14 is also bonded to the photomultiplier tubes 101 to 104 with a coupling adhesive. An adhesive layer 17 is formed. Moreover, the outer peripheral surface where each scintillator 11 does not oppose is covered with the light reflection material except the optical coupling surface with the photomultiplier tubes 101-104 side. In this case, a PTFE tape is mainly used as the light reflecting material.

FIG. 9 is a block diagram showing the configuration of the position calculation circuit of the radiation detector. The position calculation circuit is composed of adders 21, 22, 23, 24 and position discrimination circuits 25, 26. As shown in FIG. 9, in order to detect the incident position of the gamma rays in the X direction, the output P1 of the photomultiplier tube 101 and the output P3 of the photomultiplier tube 103 are input to the adder 21 and the photomultiplier The output P2 of the double tube 102 and the output P4 of the photomultiplier tube 104 are input to the adder 22. The addition outputs (P1 + P3) and (P2 + P4) of both adders 21 and 22 are input to the position discriminating circuit 25, and the incident position of the gamma rays in the X direction is obtained based on both addition outputs.

Similarly, in order to detect the incident position of the gamma rays in the Y direction, the output P1 of the photomultiplier tube 101 and the output P2 of the photomultiplier tube 102 are input to the adder 23 and the photomultiplier tube 103 The output P3 and the output P4 of the photomultiplier tube 104 are input to the adder 24. The addition outputs (P1 + P2) and (P3 + P4) of both adders 23 and 24 are input to the position discrimination circuit 26, and the incident position of the gamma ray in the Y direction is obtained based on both addition outputs.

Further, the calculated value (P1 + P2 + P3 + P4) indicates energy for the event, and is displayed as an energy spectrum as shown in FIG.

The results calculated as described above are represented as a position coding map as shown in FIG. 11 according to the position of the gamma ray incident on the scintillator, and each position discrimination information is shown.

Further, image collection methods include a two-dimensional collection shown in FIG. 12 and a three-dimensional collection shown in FIG. In the two-dimensional collection, the scepter 31 is disposed between the detector rings 32, and has a structure capable of removing scattered radiation, random coincidence coefficient, and the like. Data acquisition of the direct slice 33 that takes the simultaneous coefficient in the same detector ring and the cross slice 34 that takes the simultaneous coefficient between the adjacent rings can be performed. In an apparatus having an n-layer detector ring, a 2n-1 slice PET image is obtained. I can image. The material of the scepter 31 is heavy metal such as tungsten or lead for gamma ray shielding. On the other hand, in the three-dimensional acquisition, since the simultaneous coefficient of the detector ring 42 is performed between all the detector rings, PET images in the direct slice 43 and all the slices 44 can be captured. In this case, the scepter is collected in a detached state. In other words, since the scepter must be arranged and removed depending on the collection conditions, it is a mechanism that can be taken in and out of the gantry.

On the other hand, scintillator arrays each made of a material with different emission decay times are stacked in multiple stages (see, for example, Non-Patent Document 1), and further, each scintillator array is arranged with a half-pitch shift (see, for example, Non-Patent Document 2). Various methods for improving the spatial resolution by realizing a block detector having DOI (depth of interaction) information have been proposed.

JP 2004-354343 A

S. Yamamoto and H. Ishibashi, A GSO depth of interaction detector for PET, IEEE Trans. Nucl.Sci., 45: 1078-1082, 1998.

H. Liu, T. Omura. M. Watanabe, et al., Development of a depth of interactiondetector for g-rays, Nucl. Instr. Meth., Physics Research A 459: 182-190, 2001.

However, the above-described conventional radiation detector has the following problems.

The two-dimensional collection shown in FIG. 12 has advantages such as low noise due to the effect of removing scattered rays by the scepter, high quantitativeness, and uniform sensitivity distribution in the body axis direction. It is extremely low, about 1/5 to 1/6. On the other hand, the three-dimensional acquisition shown in FIG. 13 has an advantage that the sensitivity is about 5 to 6 times higher than that in the case of the two-dimensional acquisition. In order to acquire two data of two-dimensional collection and three-dimensional collection from the subject, it is necessary to perform measurement twice, but there is a problem that the measurement time is long and the burden on both the subject and the radiographer is large. Furthermore, since the gamma rays emitted from the radioisotope (RI) decay with time, it is necessary to inject a large amount of RI corresponding to the gamma rays in advance, which increases the cost.

Such a problem cannot be solved only by a structure in which scintillator arrays are simply stacked in multiple stages as described in the prior art, or a structure in which each scintillator array is arranged with a half-pitch shift. In other words, even if the scintillator array is stacked, if some radiation that has passed through the uppermost scintillator array is incident on the subsequent scintillator array, the image quality by scattered radiation and accidental coincidence counting will be provided unless a separate scepter is provided. Deterioration occurs. On the other hand, when a scepter is provided, three-dimensional collection is not possible.

The radiation detector provided by the present invention is optically coupled to a scintillator group in which a plurality of scintillator arrays in which a plurality of scintillators are arranged two-dimensionally are stacked, and the scintillator array at the lowest stage with the subject side facing up. Each scintillator constituting the uppermost scintillator array is provided with each of the scintillators constituting the next scintillator array. By being placed at the top of the joint, it acts as a scepter. In that case, the scintillator has two or more different attenuation times, and the capture location of the γ photons incident on the scintillator by either one or both of discrimination by the difference in attenuation time and light sharing by the reflector structure. In order to specify, two-dimensional collection and three-dimensional collection can be performed simultaneously.

In order to solve the above problems, the following configuration is adopted. In other words, the radiation detector according to claim 1 is optically applied to a scintillator group in which a plurality of scintillator arrays in which a plurality of scintillators are two-dimensionally connected and arranged are stacked, and the scintillator array at the lowest stage with the subject side up. Each scintillator constituting the uppermost scintillator array constitutes the next scintillator array, and a light guide optically coupled to the light guide and a photomultiplier tube optically coupled to the light guide. The scintillator is arranged above the joint of each scintillator and functions as a scepter.

Based on such characteristics, the invention described in claim 1 operates as follows. That is, the radiation incident on each scintillator constituting the uppermost scintillator array is converted into light by the scintillator, and the light is amplified by a photomultiplier tube and counted to obtain position discrimination information. On the other hand, radiation that does not enter each scintillator constituting the uppermost scintillator array is converted into light by each scintillator constituting the next scintillator array, and the light is amplified and counted by a photomultiplier tube. Get position discrimination information. At this time, the uppermost scintillator array is made of a material that does not transmit incident radiation, and serves as a scepter.

Due to this action, the invention described in claim 1 can perform two-dimensional collection with little degradation in image quality due to scattered radiation and coincidence coincidence with the next scintillator array while performing three-dimensional collection with the uppermost scintillator array. There is an effect.

Further, the radiation detector according to claim 2 is the radiation detector according to claim 1, wherein each scintillator configuring the uppermost scintillator array includes each scintillator configuring the scintillator array subsequent to the next stage. It is characterized by being smaller than the width of.

Based on such characteristics, the invention described in claim 2 operates as follows. That is, each scintillator that constitutes the uppermost scintillator array is smaller than the width of each scintillator that constitutes the scintillator array in the subsequent stage and is located above the joint of each scintillator that constitutes the scintillator array in the next stage. Therefore, there is always a space for transmitting radiation between the scintillators constituting the uppermost scintillator array. Radiation reaches the scintillator array in the subsequent stages through this space.

Due to this action, the invention described in claim 2 is formed between the scintillators constituting the uppermost scintillator array even if the scintillator arrays in the subsequent stages are arranged so that the scintillators are closely arranged. There is an effect that the radiation can reach the scintillator arrays in the subsequent stages through the space.

The radiation detector according to claim 3 is the radiation detector according to claim 1, wherein the capture location of γ photons incident on the scintillator group is specified by sharing light by a reflector structure. Features.

Based on this feature, the invention according to claim 3 operates as follows. That is, light generated by radiation incidence in the scintillator is guided to the photomultiplier tube by the reflector structure. From the output of each photomultiplier tube, the capture location of γ photons is specified by light sharing.

Furthermore, the radiation detector according to claim 4 is the radiation detector according to claim 1, wherein the scintillator group is composed of at least two types of scintillators having different decay times. It is characterized in that the capture location of γ photons incident on the scintillator is specified by the difference and the light sharing by the reflector structure.

Based on this feature, the invention according to claim 4 operates as follows. That is, when radiation is incident on a scintillator having a certain decay time and a scintillator different from the decay time, a signal in which light generated by each scintillator is mixed is obtained.

Based on such an action, the invention according to claim 4 separates the signal including the light generated by the different scintillators by the waveform selection using the difference in the decay time, and outputs the light by the above-described reflector structure. By this method of sharing, it is possible to specify the capture location of γ photons incident on the scintillator. Furthermore, by configuring each scintillator array with a material having a different attenuation time, it is possible to obtain information in the depth direction based on the difference in the attenuation time, and the spatial resolution can be improved.

(First embodiment)
Hereinafter, the configuration of the first embodiment of the radiation detector of the present invention is shown in the drawings and will be described in detail according to the embodiment. FIG. 1 is a cross-sectional view (front view) in the X direction of the radiation detector 50 of the present invention viewed from the Y direction. FIG. 2 is a sectional view (side view) in the Y direction of the radiation detector 50 of the present invention as viewed from the X direction. FIG. 3 is a sectional view (top view) of the upper scintillator array 61 when the radiation detector 50 of the present invention is viewed from above. The scintillator group 52 of the radiation detector 50 includes two parts, an upper scintillator array 61 and a lower scintillator array 62. In the cross-sectional view in the X direction of FIG. 1, six scintillators 58 are arranged in the X direction in the upper scintillator array 61, and six scintillators 51 are arranged in the X direction in the lower scintillator array 62. Here, the center pitch of the scintillator 58 in the X direction coincides with the center pitch of the scintillator 51. On the other hand, in the cross-sectional view in the Y direction of FIG. 2, the scintillator 58 in the Y direction in the upper scintillator array 61, the six light transmitting materials 59 in the Y direction, and the scintillator 51 in the Y direction in the lower scintillator array 62. 6 are in a line. Here, the center pitch of the scintillator 58 in the Y direction is arranged so as to deviate from the center pitch of the scintillator 51 by ½. The scintillators 58 at both ends have half the width w with respect to the other scintillators 58. The light transmitting material 59 is disposed between the scintillators 58 in the Y direction, and is selected from materials having good light transmittance and good gamma ray energy, for example, transparent acrylic.

3 shows a portion of the upper scintillator array 61 as viewed from above, and the scintillator 58 divides a portion corresponding to the ring when the radiation detector 50 is incorporated into the PET gantry into each ring. Has been placed. The radiation detector 50 is partitioned by appropriately sandwiching the light reflecting material 53. In the upper scintillator array 61, there are six scintillators 58 in the X direction, seven in the Y direction, and six light transmitting materials 59 in the Y direction only. In the lower scintillator array 62, a scintillator group 52 in which six scintillators 51 are arranged two-dimensionally in close contact with the X direction and six in the Y direction are optically coupled to the scintillator group 52 and a light reflecting material 55 is combined. A light guide 54 in which a lattice frame is embedded and a plurality of small sections are defined, and four photomultiplier tubes 201, 202, 203, 204 optically coupled to the light guide 54. ing. Here, as the scintillator, for example, Bi ₄ Ge ₃ 0 ₁₂ (BGO), Gd ₂ SiO ₅ : Ce (GSO), Lu ₂ SiO ₅ : Ce (LSO), LuYSiO ₅ : Ce (LYSO), NaI, BaF ₂ , CsF, inorganic crystals such as PbWO ₄ used. In particular, the scintillator 58 of the upper stage scintillator array 61 has a particularly high gamma ray shielding ability and a large effective atomic number, Bi ₄ Ge ₃ 0 ₁₂ (BGO), Lu ₂ SiO ₅ : Ce (LSO), LuYSiO ₅ : Ce (LYSO). PbWO ₄ is preferred.

Further, the image acquisition method includes the two-dimensional acquisition and the three-dimensional acquisition described above. In the case of the radiation detector of the present invention, a PET image can be taken without distinction. FIG. 4 shows a sectional view in the body axis direction when the radiation detector of the present invention is incorporated in a gantry. Here, the scepter 71 is disposed between the detector rings 72 and has a structure capable of removing scattered radiation, random coincidence coefficient, etc., but the scepter 71 is the scintillator 58 itself of the upper scintillator array 61, and the detector The ring 72 corresponds to the scintillator 51 of the lower scintillator array 62. FIG. 4 further shows a part of the locus by gamma rays, but data collection of the direct slice 73 taking the simultaneous coefficient in the same detector ring and the cross slice 74 taking the simultaneous coefficient between the adjacent rings can be performed. Here, since the light transmitting material 59 has good permeability to the energy of gamma rays, the gamma rays can be transmitted without losing energy. As described above, the material of the scepter 71, that is, the scintillator 58, is a scintillator having a large effective atomic number for gamma ray shielding. That is, if the detection information of only the scintillator 51 of the lower scintillator array 62 is used, a PET image based on the two-dimensional collection data can be captured. On the other hand, since the scintillator 58 of the upper scintillator array 61 also performs gamma ray detection, data collection of the slice 75 is performed. That is, if the detection information of the scintillator 51 of the lower scintillator array 62 and the scintillator 58 of the upper scintillator array 61 is used, a PET image based on the three-dimensional collection data can be captured.

Here, as shown in FIG. 1 and FIG. 2, the radiation detector of the present invention is converted into visible light when gamma rays are incident on any one of the scintillator groups 52, and photoelectron amplification is performed through a light guide 54 that is optically coupled. Light is guided to the double tubes 201 to 204, and at this time, the output ratio of the photomultiplier tubes 201 (203) and the photomultiplier tubes 202 (204) arranged in the X direction is changed at a constant rate. The position, length, and angle of each light reflecting member 55 in the light guide 54 are adjusted. The result of calculating the output is represented as a position coding map as shown in FIG. 5 according to the position of the gamma ray incident on the scintillator group 52, and each position discrimination information is shown.

（例２）

（例３）

出力による位置計算と減衰時間により弁別された結果はシンチレータ群５２に入射したガンマ線の位置に従って、上段シンチレータアレイ６１は図７に示すような位置コーディングマップとして表され、下段シンチレータアレイ６２は図６に示すような位置コーディングマップとして表される。上段シンチレータアレイ６１と下段シンチレータアレイ６２は位置コーディングマップ上で完全に分離されているため、両者の重なり（クロストーク）はなく、その結果ＰＥＴ像の空間分解能が劣化するということはない。 (Second embodiment)
In the first embodiment described above, since the incident position of the gamma ray is discriminated only by light sharing by the reflector structure, the map of the upper scintillator array 61 and the map of the lower scintillator array 62 are both plotted in the position coding map. As a result, the overlap (crosstalk) between the two becomes large, and as a result, the spatial resolution of the PET image may be deteriorated. Therefore, in the second embodiment, the upper scintillator array 61 and the lower scintillator array 62 have different attenuation times, and the capture location of γ photons incident on the scintillator due to the difference in attenuation time and the sharing of light by the reflector structure. It has a structure that identifies. For example, the following combinations are possible.
(Example 1)

(Example 2)

(Example 3)

According to the position calculation by the output and the discrimination result by the decay time, the upper scintillator array 61 is represented as a position coding map as shown in FIG. 7 according to the position of the gamma rays incident on the scintillator group 52, and the lower scintillator array 62 is shown in FIG. It is represented as a position coding map as shown. Since the upper scintillator array 61 and the lower scintillator array 62 are completely separated on the position coding map, there is no overlap (crosstalk) between them, and as a result, the spatial resolution of the PET image does not deteriorate.

Further, as a modified example, the upper scintillator array 61 or the lower scintillator array 62 is divided into a plurality of stages in the depth direction and has different attenuation times for each of the stages. By this, the capture location of γ photons incident on the scintillator is specified. In this case, a higher density spatial resolution can be obtained.

(Third embodiment)
In the third embodiment, the light transmitting material 89 is not present, and the other portions are completely the same as the first and second embodiments described above. Due to the absence of the light transmitting material 89, there is no attenuation during transmission of gamma rays, and detection sensitivity is advantageous.

  In the embodiment described above, as a desirable form, the width of the upper scintillator 58 in the Y direction is made smaller than the width of the lower scintillator 51 in the Y direction, but these may be the same width. The upper scintillator 58 may be wider. In that case, the upper scintillators 58 may be arranged at intervals, and the lower scintillators 51 may be arranged at intervals greater than or equal to the intervals. That is, as long as gamma rays can reach the lower scintillator 51 from the gaps between the upper scintillators 58, various widths of the scintillators can be selected.

As described above, the radiation detector of the present invention does not require a scepter made of heavy metal as in the prior art, and can perform two-dimensional collection and three-dimensional collection simultaneously. Therefore, it is not necessary to perform two measurements to acquire two data of two-dimensional collection and three-dimensional collection from the subject as in the past, and the measurement time can be shortened, and both the subject and the radiation technician are burdened. Can be reduced. Furthermore, although gamma rays emitted from the radioisotope (RI) are attenuated in time, RI can be saved because the measurement time can be shortened, and the cost burden can be reduced.

As described above, the present invention is suitable for medical and industrial radiation imaging apparatuses.

Sectional drawing of the X direction of the radiation detector of this invention is shown. Sectional drawing of the Y direction of the radiation detector of this invention is shown. Sectional drawing seen from the upper surface of the radiation detector of this invention is shown. The body axis direction sectional drawing of the radiation detector of this invention is shown. 1 shows a position coding map of a first embodiment of the radiation detector of the present invention. 6 shows a position coding map of a second embodiment of the radiation detector of the present invention. 6 shows a position coding map of a third embodiment of the radiation detector of the present invention. Sectional drawing of the conventional radiation detector is shown. An example of the position calculation circuit of the radiation detector of this invention and the conventional radiation detector is shown. The energy spectrum of the radiation detector of this invention and the conventional radiation detector is shown. 2 shows a position coding map of a conventional radiation detector. The body axis direction sectional drawing in the two-dimensional collection of the conventional radiation detector is shown. The body axis direction sectional drawing in the three-dimensional collection of the conventional radiation detector is shown.

Explanation of symbols

DESCRIPTION OF SYMBOLS 10 ... Conventional radiation detector 11 ... Scintillator 12 ... Scintillator group 13 ... Light reflecting material 14 ... Light guide
DESCRIPTION OF SYMBOLS 15 ... Light reflecting material 16 ... Coupling adhesive layer 17 ... Coupling adhesive layer 21, 22, 23, 24 ... Adder 25, 26 ... Position discrimination circuit 31 ... Scepter 32 ... Detector ring 33 ... Direct slice 34 ... Cross slice 43 ... direct slice 44 ... all slices 50 ... radiation detector 51 according to the first embodiment of the present invention ... scintillator 52 ... scintillator group 53 ... light reflector 54 ... light guide 55 ... light reflector 56 ... coupling adhesion Agent layer 57 ... Coupling adhesive layer 58 ... Scintillator 59 ... Light transmitting material 61 ... Upper scintillator array 62 ... Lower scintillator array 71 ... Scepter 72 ... Detector ring 73 ... Direct slice 74 ... Cross slice 75 ... Slices 101, 102, 103, 104 ... photomultiplier tubes 201, 2 2,203,204 ... photomultiplier tube

Claims (4)

A scintillator group in which a plurality of scintillator arrays in which a plurality of scintillators are two-dimensionally stacked are stacked, a light guide optically coupled to the lowermost scintillator array with the subject side up, and the light guide Each scintillator constituting the uppermost scintillator array is disposed above the joint of each scintillator constituting the next scintillator array, and includes a scepter. A radiation detector characterized by functioning as: 2. The radiation detector according to claim 1, wherein each scintillator constituting the uppermost scintillator array is smaller than a width of each scintillator constituting the next scintillator array. 2. The radiation detector according to claim 1, wherein a capture location of γ photons incident on the scintillator group is specified by light sharing using a reflector structure. The scintillator group is composed of at least two types of scintillators having different decay times, and specifies the capture location of γ photons incident on the scintillator by the difference in decay times and the sharing of light by the reflector structure. The radiation detector according to claim 1.

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