JPS62206479A - Radiation detector - Google Patents

Abstract

PURPOSE:To reduce the number of photoelectric multipliers, by arranging sixteen scintillators in a matrix of 4X4 to connect photoelectric multipliers to scintillators, one per four as grouped in four zones divided by 2X2 areas. CONSTITUTION:A detector section 11 which detects radiation such as gamma rays radiated from an radioactive isotope dosed into a body as an object to be inspected is composed by combining sixteen scintillators 13 and four photoelectric multipliers 14. The scintillators 13 are arranged in a matrix of 4X4 and divided into a total of 9 blocks, four positioned at the center thereof, one per one of four corners and one with two scintillators each in the remaining perimeters. These blocks are optically separated from one another by binding boundary surfaces thereof with reflecting agent layers. Photoelectric multipliers are optically connected to sixteen scintillators 13, one per four as grouped in 4 zones divided by 2X2 areas.

Description

DETAILED DESCRIPTION OF THE INVENTION Field of Industrial Application The present invention relates to a radiation detector used in, for example, a positive 1 Heron CT apparatus that obtains tomographic images by detecting radiation emitted from a radioactive isotope administered into the body of a subject. In particular, the present invention relates to a radiation detector that can simultaneously improve the spatial resolution within the tomographic plane and the spatial resolution in the body axis direction without increasing the number of photomultiplier tubes in the detector section.

2. Description of the Related Art For example, a radiation detector 1 in a positron CT apparatus generally includes a light emitting surface of one scintillator 2 □ and a light receiving surface of one photomultiplier tube 3, as shown in FIGS. 4 and 5. It consists of a detector section 4 which is optically coupled to a surface using silicone grease or the like, and a signal pre-processing circuit 5 which processes the output signal of the photoelectron multiplier 93. Here, since the upper limit of the theoretical spatial resolution of the tomographic image obtained by the above-mentioned positron CT apparatus is 1/2 of the width W of the scintillator 2, the radiation detector 1 in the recent positron CT apparatus , scintillator 2 with a small width W□ has come to be used. However, with current technology, photomultiplier tube 3
The diameter of the scintillator 2 is at least about 1×2 inches, and is limited by the diameter of the photomultiplier tube 3, so the width W□ of the scintillator 2 cannot be smaller than 1×2 inches.

In addition, in a positron CT device, a large number of radiation detectors are used to efficiently detect gamma rays that are generated when positrons emitted from a radioactive isotope administered into a subject's body combine with negative electrons in the body and disappear. 1 are arranged in a lip shape around the subject. Here, in order to improve the spatial resolution of the tomographic image as described above, if the width W4 of the scintillators 2 is reduced and they are arranged in a ring shape, the photoelectron multiplier 4 is optically coupled to each scintillator 2. 'f! 3 would increase significantly, and the price of the ring-shaped radiation detector system would rise.

Therefore, as shown in Japanese Patent Application No. 54-124742, a method has been proposed to reduce the number of photomultiplier tubes 3 in the ring-shaped radiation detector system.

That is, as shown in FIG. 6, four narrow (W2<W□) scintillators 2a, 2b, 2c, and 2d are arranged in a straight line in the direction of the chain arrow, and the first scintillator 2a
and the second scintillator 2b, and between the third scintillator 2c and the fourth scintillator 2d, by the reflector layer 6, and the second scintillator 2b and the third scintillator 2c. are optically coupled by an optical coupling layer 7, and further provided with a first scintillator 2 and a second scintillator 2.
The first photomultiplier tube 3a was optically coupled to the scintillators a and 2b, and the second photomultiplier tube 3b was optically coupled to the third and fourth scintillators 2c and '2d. Then, by processing combinations of the presence/absence and strength of output signals from each of the photomultiplier tubes 3a and 3b in a signal preprocessing circuit (not shown), the gamma rays are incident on which scintillator among the four scintillators 28 to 2d. It was supposed to detect whether it was done.

In the radiation detector 1' having such a configuration, one photomultiplier tube 3 is combined for each two scintillators 2, so the photomultiplier is higher than the conventional one-to-one combination. The number of tubes 3 can be reduced. Also,
When the direction of the dashed line arrow A is taken as the direction of the tomographic plane, the width W2 of the scintillator 2 in this direction is narrower than that of the conventional scintillator, so that the spatial resolution within the tomographic plane can be improved.

Problems to be Solved by the Invention However, in general, the half-life of a radioactive isotope administered into the body of a subject is short, so recent positron CT devices are difficult to measure because they measure a wide area of the subject at once. In addition, the ring-shaped radiation detector systems are now arranged in multiple rows along the body axis of the subject. in this case,
In the radiation detector 1' shown in Fig. 6, assuming that the diameter of the photomultiplier tubes 3a and 3b is approximately 1 x 2 inches, the width W in the body axis direction indicated by the solid arrow B cannot be made smaller than this. . Therefore, if a ring-shaped radiation detector system is constructed using the conventional radiation detector 1' shown in FIG. 6, and if this radiation detector system is arranged in multiple rows in the body axis direction, It was not possible to improve the accuracy of the position information in the body axis direction shown in B, and it was not possible to improve the spatial resolution. Therefore, the present invention aims to provide a radiation detector that can simultaneously improve the spatial resolution within the tomographic plane and the spatial resolution in the body axis direction without increasing the number of photomultiplier tubes in the detector section. purpose.

Means for Solving the Problems The means of the present invention for solving the above problems includes a plurality of scintillators and a plurality of photomultiplier tubes, and optically couples the light-emitting surface and light-receiving surface of each. and a signal preprocessing circuit for positioning the scintillator into which the radiation has entered and extracting timing information by comparing the peak values of the output signals of each photomultiplier tube. , 16 scintillators are arranged in a 4×4 matrix, and among these scintillators, 4 scintillators are located in the center and 1 scintillator is located in the four corners.
Each scintillator is optically separated from two scintillators located in the remaining peripheral area, and each scintillator is optically coupled within each separated block, so that the scintillators arranged in a matrix are This is accomplished by optically coupling one photomultiplier tube to each of four scintillators divided into four 2×2 areas.

Embodiments Hereinafter, embodiments of the present invention will be described in detail with reference to the accompanying drawings.

FIG. 1 is an explanatory diagram showing an embodiment of a radiation detector according to the present invention. This radiation detector 10 detects radiation such as gamma rays emitted from a radioisotope administered into the body of a subject in a positron CT apparatus, for example, and is composed of a detector section 11 and a signal preprocessing circuit 12. . The detector unit 11 is configured to receive radiation emitted from the body of the subject and send out an electrical signal, and includes a plurality of scintillators 13 that cause scintillation and emit light when the radiation is incident. , a plurality of photomultiplier tubes 14 that receive the light emitted from these scintillators 13 and convert it into an electric signal, and the light receiving surface of the photomultiplier tube 14 is optically coupled to the joint surface of the scintillator 13. There is. Further, the signal preprocessing circuit 12 performs positioning of the scintillator 13 into which radiation is incident and extracts timing information by comparing the peak values of the output signals of the respective photomultiplier tubes 14. It is connected to the signal lead line 15.

Here, in the present invention, the detector section 1] has 16
It is made up of a combination of four scintillators 13 and four photomultiplier tubes 14. That is, as shown in FIG. 2, first, 16 scintillators C□□, C□21C131...
・C4, are arranged in a 4×4 matrix. These scintillators C□□ to C44 are composed of four scintillators C22, C23, and C32 located in the center.
, C33, and one scintillator C1□, C14, C4□ located at each of the four corners.

C44 and two scintillators C□2. located at the remaining periphery. C□3; C21, C3□; C24, C34
; Divided into a total of nine blocks, C4□ and C43,
The boundary surfaces of each block are joined by a reflector layer 16 made of barium sulfate or the like, and are optically separated. Also,
Inside each block separated by the reflector layer 16, each scintillator is joined by an optical coupling layer 17 made of silicone grease or the like and optically coupled. That is, in the block located in the center, the scintillators C2□+C231C3□, C33 are optically coupled to each other by the optical coupling layer 17, and in the four blocks located in the peripheral part, the scintillators C1□, C13, C
2, and C3□, C24 and C34, and C4□ and C43 are optically coupled to each other by an optical coupling layer 17, respectively. moreover,
The 16 scintillators C to C44 in the 4×4 71-iron arrangement are divided into four groups of 4 scintillators each in a 2×2 area, and the scintillators in each group are One photomultiplier tube P1. P2. P, , P4 are arranged and combined.

That is, the first group of scintillators C1□, c,
2. First photomultiplier tube I for C2 engineering and C2□]
□ is the scintillator C3□, C32, C of the second loop
41, for C4□, the second photomultiplier tube P2 is the scintillator C of the third group.

A third photomultiplier tube P3 is connected to C1, , C23, , C, 4, and a fourth group of scintillators c33 . C,,
.

A fourth photomultiplier tube P4 is arranged for each of C43 and C44, and the light-receiving surface of each photomultiplier tube P to P4 is optically coated with silicone grease or the like to the joint surface of each scintillator C11 to C44. is combined with

As shown in FIG. 2, the outer peripheral surfaces of the 16 scintillators C□, ~C44 arranged in a matrix are all coated with a reflector layer 16, and in addition, in FIG. A reflector layer 16 is applied to all of the upper surfaces of the scintillators 13 that are not connected to the photomultiplier tubes P1 to P4 and the bottom surfaces of the plurality of scintillators 13. This prevents the light emitted by the scintillators C□□ to C44 from leaking to the outside.

Next, in the detector section configured as described above]1, radiation such as gamma rays is transmitted through 16 scintillators C□1 to C4.
4, for example, the first group of scintillators C□□～
Table 1 shows the magnitude relationship of the output signals from each of the photomultiplier tubes P1 to P4 when incident on C2□.

Table 1 In Table 1, ◎ indicates that the peak value of the output signal is large, O indicates medium, Δ indicates small, and X indicates that there is no output signal. First, when radiation enters the scintillator C11 and the scintillator C1 emits light, as shown in FIG.
16 by other scintillators C1゜l C21,,C
2□, the above-mentioned light emission enters only the first photomultiplier tube P□ and enters the other photomultiplier tube P2.
~It does not enter P4. Therefore, an output signal with a large peak value appears only in the -th photomultiplier tube P, and there are no output signals in the other photomultiplier tubes P2 to P4. Next, when radiation enters the scintillator C+-z and this scintillator C1□ emits light, as shown in FIG. Since the emitted light is optically coupled to the scintillator c13, a part of the emitted light is also distributed to the scintillator c13, and the emitted light is incident on the first photomultiplier tube P, and the distributed light is also transmitted to the third photomultiplier tube P. It also enters the tube P. Therefore, an output signal with a medium peak value appears in the first photomultiplier tube (■), and an output signal with a small peak value also appears in the third photomultiplier tube P3. Double tube P2. There is no output signal on P4. Also, the radiation enters the scintillator C2□ and this scintillator C
21 emits light, as shown in FIG. 2, since the scintillator C21 is optically coupled to the adjacent scintillator C31 by the optical coupling layer 17 within the block, a part of the emitted light is transmitted to the scintillator C3. , and the above-mentioned emitted light is incident on the first photomultiplier tube P, and the above-mentioned distributed light is also incident on the second photomultiplier tube P2. Therefore, an output signal with a medium peak value appears in the first photomultiplier tube P□, and an output signal with a small peak value also appears in the second photomultiplier tube P2, and other photomultipliers Pipe P3. There is no output signal on P4. Furthermore, the radiation enters the scintillator C2□ and this scintillator C
2, when the scintillator C2□ emits light, as shown in FIG. C3□
Since the other scintillator C33 adjacent to the scintillator C33 is also optically coupled by the optical coupling layer "-7, a part of the emitted light is distributed to the scintillators C23 and C3° as well as to the other scintillator C33, The above-mentioned light emission enters the first photomultiplier tube P, and each of the above-mentioned distributed lights also enters the second to fourth photomultiplier tubes P2 to P4. Therefore, an output signal with a medium peak value appears in the first photomultiplier tube P, and output signals with small peak values also appear in the second to fourth photomultiplier tubes P2 to P4. From this, as shown in Table 1, each photomultiplier tube P1
~ By comparing the magnitude relationship of the peak values of the output signals from P4, the scintillator C1 at the position where the radiation is incident,
〜C2□ can be known.

In addition, in the following explanation and Table 1, the case where radiation is incident on the scintillators Ct, ~C2° of the first group is described, but the scintillators C of the other groups are
3□ to C4z; C13 to C24; C33 to C44, the magnitude relationship between the incident position of radiation and the peak value of the output signal of each photomultiplier tube P□ to P4 is exactly the same as in Table 1 above. You can get the following. Therefore, for the other groups of scintillators mentioned above, in the same way, the peak value of the output signal from each photomultiplier tube P1 to P4 is
By comparing the JX relationship, it is possible to know the scintillator at the position where the radiation is incident.

FIG. 3 is a block diagram showing the signal preprocessing circuit 12 in the present invention. The operation of the signal preprocessing circuit 12 will be explained below based on this block diagram. First, the operation of creating a timing signal indicating the timing at which radiation enters each scintillator C3, to C44 will be described. First, the anode signals An output from the first to fourth photomultiplier tubes P1 to P4 are added and input to the first amplifier 18a, and are amplified. It is input to the discriminator 19. The timing discriminator 19 identifies the timing of incidence, and its output signal is input to the delay circuit 20 . This delay circuit 20 takes a delay time in order to match the timing with an output signal from an energy discriminator 22, which will be described later. On the other hand, the dynode signals Dy outputted from the first to fourth photoelectron amplification tubes P to P4, respectively, are
After being input to and amplified by the fifth amplifiers 18b to 18e,
The signals are respectively input to an adding circuit 21 equipped with an integrating circuit and added. The output signal from this adder circuit 21 is
It is input to an energy discriminator 22 that discriminates between annihilation radiation and scattered radiation. Next, the output signal S1 whose radiation energy has been discriminated by the energy discriminator 22 and the output signal S2 whose timing has been matched by the delay circuit 20 are input to the AND gates 1 to G1, and the AND is performed. It will be done. The output from this AND gate G1 is input to the mono multivibrator 23, and one pulse is output. In this way, the timing signal t is output from the mono-multivibrator 23 only when each of the scintillators 011 to C44 detects annihilation radiation.

Next, the scintillators C11 to C4 that detected the radiation
The operation of creating a position signal indicating the position No. 4 will be explained. First, the dynode signals Dy output from the first to fourth photomultiplier tubes P1 to P4, respectively, are sent to the corresponding second to fifth amplifiers 18, respectively.

b to 18a and are amplified. Here, the output signals resulting from the amplification by these second to fifth amplifiers 18b to 18e are referred to as A, B, C, and D, respectively. Next, each of the above output signals A, B, C, and D is used to create a position signal.
It is input to each analog calculation circuit 24a-24d, 25a-25d and 26. Each of the above analog calculation circuits 24a to 2
4d, 25a to 25d, and 26 perform the calculations shown in the corresponding frames in FIG. 3 between the signals A, B, C, and D, and output the results as analog signals Z□ to Zg, respectively. For example, the first group of analog calculation circuits 24
a to 24d are two photomultiplier tubes P1. P2; P
3. P4; P engineering, P3; P2. The difference between the amplified dynode signals of P4 is calculated and analog signals 2□ to z4 are output. In addition, the analog calculation circuit 25a of the second loop
~25d are two photomultiplier tubes p,, p2; p3.
p4; p□, P, ; P2. P.

Of the amplified dynode signals, the difference between the signal with a large amplitude multiplied by 2 and the sum of both amplified dynode signals is calculated to output analog signals Z5 to Zll. Further, the third analog arithmetic circuit 26 performs the same arithmetic operation as the analog arithmetic circuits 25a to 25d of the second loop on the amplified dynode signals of the four photomultiplier tubes P1 to P4, and generates an analog signal Z, Output. However, the value of the proportionality constant ' is different from the value of the proportionality constant in the analog calculation circuits 25a to 25d of the second loop.

In such a circuit, the correspondence relationship between the position of each scintillator 011 to C44 that detected radiation and the polarity of the analog signal z1 to Z9 output from each analog calculation circuit 24a to 24d, 25a to 25d, and 26 in that case. is shown in Table 2.

In Table 2, the cross mark indicates that the analog signal has positive polarity, the - mark indicates negative polarity, and the X mark indicates that there is no analog signal. First, when the scintillator C1□ detects radiation, a signal is output only to the first photomultiplier tube P1 as shown in Table 1 above, so among the outputs of each analog calculation circuit 24a to 26, Analog signal z1. z
3. zS, z,, z, have positive polarity, and z2. z
4. Z6. No signal appears at z and l. Next, when the scintillator C1□ detects radiation, as shown in Table 1, a signal with a medium peak value is output to the first photomultiplier tube P□, and a signal with a medium peak value is output to the third photomultiplier tube P□. Since a signal with a small peak value is also output to the multiplier tube P3, among the outputs of the analog calculation circuits 24a to 26, the analog signals Z1 to 23
．． 25 and 26 have positive polarity, and the analog signal 2. , 2g
becomes negative polarity, and no signal appears at Z4 and ZB. Further, when the scintillator C2 detects radiation, as shown in Table 1, a signal with a medium peak value is output to the first photomultiplier tube P1, and a signal with a medium peak value is output to the second photomultiplier tube P2.
Since a signal with a small peak value is outputted to
tz3. 'z, , z, , z become positive polarity, and the analog signal 25.2. becomes negative polarity, and no signal appears at Z2 and ZG.

Here, the analog calculation circuits 258 to 25 of the second loop
The output signal of d is transmitted through the 16 scintillators C shown in FIG.
Scintillators C12 and C13, C21, C31, and C2 in four blocks located at the periphery among □□ to C44
The value of the proportionality constant is set so that when any of C4 and C34, C4□ and C43 detects radiation, the polarity becomes negative. As described above, the optimum value of k is determined by the distribution ratio of scintillation light between adjacent scintillators within the block.

Further, when the scintillator C22 detects radiation, as shown in Table 1, a signal with a medium peak value is output to the first photomultiplier tube P1, and a signal with a medium peak value is output to the second to fourth photomultiplier tubes P1. Since signals with small peak values are also output to the multipliers P2 to P4, among the outputs of the respective analog calculation circuits 24a to 26, the analog signals Z□ to Z4 have positive polarity, and the analog signal Z has negative polarity. 75 to Z8 have positive polarity or negative polarity depending on the value of the proportionality constant, etc.

Here, the output signal of the third analog calculation circuit 26 is the scintillator C2□, c in the block located in the center among the 16 scintillators 011 to C44 shown in FIG.
, 3. A value of ' is set in the proportionality constant so that when any one of c, □, and C33 detects radiation, the polarity becomes negative. However, 4 in the block located in the center above
Since the distribution ratio of scintillation light is different between the individual scintillators and the two scintillators in each of the four blocks located at the periphery, the above proportionality constant ′ is
This is different from the value of the proportionality constant described above. Further, when the scintillator C2□ detects radiation, analog signals 2. ~Z8 depends on the value of the proportionality constant described above and the distribution ratio of scintillation light among the four scintillators in the block located in the center, so its polarity is not uniquely determined. However, among the four scintillators in the block located in the center, the optical conditions between the adjacent scintillators are the same vertically and horizontally.
In this case, the analog calculation circuits 25a to 2 of the second loop
Analog signals Z5 to Z, l output from 5d are signals with the same polarity.

In addition, in the above explanation, the case where radiation was incident on the scintillators 01□ to C22 of the first group was described, but the scintillators C81 to C42; C□3 of the other groups were described.
~C74; By considering the case where the light is incident on C33 to C44 in the same manner, the correspondence table shown in Table 2 can be obtained.

Next, as described above, each analog calculation circuit 248 to 2
Analog signals 2□~Z output from 6.

are input to the corresponding comparators 27a to 27j, respectively. Each of the comparators 27a to 27j determines the polarity of the input analog signals 2□ to Z9, and outputs a digital signal if the input signal has positive polarity. and second and second comparators 27a,
The output from 27b is input to OR gate G2, the output from third and fourth comparators 27c and 27d is input to OR gate 1-G3, and fifth and sixth comparator 27
Outputs from e and 27f are input to OR gate G4, and outputs from seventh and seventh comparators 27g and 27h are input to O
Input to R gate GS. Also, the ninth comparator 2
The output from 7i is AND gate G□. input to one of the input terminals.

In this state, the outputs of the OR gates G4 and G5 are input to one input terminal of the AND gate G11 or G9, respectively, and are also input to Ex and the OR gate b a s. The outputs of Ex and OR gate GG are input to the other input terminals of the AND gates 1 to GI1 and G9, and after being inverted in logic by the inverter 28, they are input to the AND gate G□.

is input to the other input terminal. Next, the outputs from the AND gates G8 and Gg are respectively OR gate G1
The output from the AND gate G1o is input to the other input terminal of the OR gates G11 and G□2. With the logic circuit having such a configuration, when the output logics of the OR gates G4 and 05 do not match, the outputs of the OR gates G4 and G5 are set to 0, respectively.
It is output from the R gates G,..., G1□. Also, OR
When the output logics of gates G4 and 05 match, the output from the ninth comparator 271 is output from OR gate G□□
and G□2.

The outputs from the OR gates G2 and G3 are directly applied to the AND gates G13. input to one input terminal of G14, and also input to one input terminal of the OR gates G11 and G1.
The outputs from □ are input to one input terminal of AND gates G15 and G16, respectively. In this state, the timing signal t output from the mono-multivibrator 23 is applied to the AND gate G13. G□,, G15
．． Input each to the other input terminal of G□6 and AND
is taken. As a result, these AND gates 613~
The signal output from G is a position signal indicating the position of the scintillator that detected radiation.
3t P4.

Table 3 shows the correspondence between the positions of the scintillators C□1 to C44 that detected radiation through such circuit operation and the position signals P and p4 output in that case.

Table 3 In Table 3, if the scintillator C□□ detects radiation, each analog calculation circuit 24. Among the analog signals output from a to G26, 2□t 231 Z51 Z71Z has positive polarity, so the OR gate G2. G3-G4. The output of G5 becomes "1". Also, since the output logics of OR gates G4 and G5 match, only AND gate G1o is opened, and ○R gate G□□.

The output "1" of the ninth comparator 27j is output from G1□. Therefore, at this time, AND gate 013
~Position signal P s +P2 output from G□6
1 Pa+P4 means all bits “
1". Next, when the scintillator C12 detects radiation, as shown in Table 2, among the analog signals output from each analog calculation circuit 24a to 26, zl
, z,, z3. z5. Since OR gate G2.z has positive polarity, OR gate G2. G3. The output of G4 becomes "1".

Also, since analog signal Z7 has negative polarity and zl+ is absent, the output of OR gate G5 becomes 1'0'', and Ex
, OR gate G becomes II I II, and the output logic from OR gates G□□ and G□2 is outputted from OR gate G4 . Therefore, at this time, the position signals p□-P4 output from the AND gates G-G16 are "1" as shown in Table 3. ,1,1. The bit pattern is "O". By considering the case where other scintillators detect radiation in the same way, the correspondence table shown in Table 3 can be obtained.

Effects of the Invention Since the present invention is configured as described above, the 16 scintillators 0□1 to C44 are arranged in a 4 x 4 matrix, and each 2 x 2 area is divided into four sections. One photomultiplier tube P1 for each scintillator. P
2. P3. By positionally coupling P4, the number of photomultiplier tubes can be reduced compared to one-to-one coupling to a conventional scintillator. Furthermore, as is clear from a comparison between FIG. 6 showing the prior art and FIG. 2 showing the present invention, even if the diameters of the photomultiplier tubes P1 to P4 are reduced to about 1/2 inch, The scintillators C1□ to C44 according to the present invention have a width W in the tomographic direction indicated by the chain line arrow A and a width W in the body axis direction indicated by the solid line arrow B at the same time, which is larger than the diameter of the photomultiplier tubes P1 to P4. It can be made small. Therefore, the radiation detector 10 of the present invention
By configuring a ring-shaped radiation detector system using , and further arranging this radiation detector system in one or more rows in the body axis direction, it is possible to improve the spatial resolution in the tomographic direction shown by arrow A, and at the same time, It is possible to improve the accuracy of positional information in the body axis direction indicated by arrow B and improve its spatial resolution. This makes it possible to provide more accurate diagnostic information by making finer tomographic images obtained by, for example, a positron CT device.

[Brief explanation of drawings]

Fig. 1 is an explanatory diagram showing an embodiment of the radiation detector according to the present invention, Fig. 2 is an explanatory plan view showing the optical coupling state of each scintillator and photomultiplier tube in the detector section, and Fig. 3 is an explanatory diagram showing a signal FIG. 4 is a plot diagram showing the preprocessing circuit, FIG. 4 is an explanatory diagram showing the first conventional example, FIG. 5 is a plan view showing the detector section thereof, and FIG. 6 is the detector section of the second conventional example. FIG. ] C0... Radiation detector 11... Detector section 12... Signal pre-processing circuit 13... Multiple scintillators 14... Multiple photomultiplier tubes 15... Signal leader line 16... -Reflector layer 17...optical coupling layer

Claims (1)

[Claims] Radiation is detected by comparing the peak values of the output signals of each photomultiplier tube with a detector section that has multiple scintillators and multiple photomultiplier tubes and optically couples the light emitting surface and light receiving surface of each. In a radiation detector comprising a signal preprocessing circuit that positions the incident scintillator and extracts timing information, the detector section has 16 scintillators arranged in a 4×4 matrix, and among these scintillators, The four scintillators located in the center, one scintillator each located in the four corners, and the remaining two scintillators located in the periphery are optically separated, and within each separated block, Each scintillator is optically coupled, and the scintillators arranged in a matrix are divided into four regions of 2 x 2, and one photomultiplier tube is arranged and optically coupled to each of the four scintillators. A radiation detector characterized by: