JP2005106807A - Semiconductor radiation detector, positron emission tomograph, semiconductor radiation detecting device, detector unit and nuclear medicine diagnosis system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To improve sensitivity of a semiconductor radiation detecting element used in a nuclear medicine diagnosis system. <P>SOLUTION: A semiconductor radiation detector 21 is made up so that a semiconductor portion of CdTe which mutually interacts with γ-rays and generates charges, is sandwiched by an anode electrode A and a cathode electrode C disposed opposite to each other. The thickness of the semiconductor portion sandwiched by the opposite anode electrode A and cathode electrode C is 0.2-2 mm. Respective elements are mounted (or arranged) on substrates 20 (20a, 20b) so that the semiconductor radiation detector 21 (or its conductor) is in close proximity to an analog ASIC 24 (or its conductor) which processes a signal detected by the detector 21. Furthermore, the substrates 20 (20a, 20b) on which the detector 21 is mounted, are contained in a chassis, thereby unitizing them (as a detector unit). <P>COPYRIGHT: (C)2005,JPO&NCIPI

Description

The present invention relates to a nuclear medicine diagnostic apparatus, and more particularly to a positron emission tomography (Positron) which is a kind of a nuclear medicine diagnostic apparatus using a semiconductor radiation detector, a semiconductor radiation detection apparatus, and a detector unit.
Emission Tomography) apparatus (hereinafter referred to as PET apparatus).

Conventionally, a radiation detector using NaI scintillator is known as a radiation detector for detecting radiation such as γ rays. In a gamma camera (a type of nuclear medicine diagnostic apparatus) equipped with a NaI scintillator, radiation (γ rays) is incident on the scintillator at an angle limited by a large number of collimators and interacts with the crystal of NaI to produce scintillation light. To emit. This light sandwiches the light guide and reaches the photomultiplier tube to become an electric signal. The electrical signal is shaped by a measurement circuit attached to the measurement circuit fixing board and sent from the output connector to an external data collection system. Note that the scintillator, light guide, photomultiplier tube, measurement circuit, measurement circuit fixing board, etc. are entirely housed in a light shielding shield case to block electromagnetic waves other than external radiation.

In general, a gamma camera using a scintillator has a structure in which a large photomultiplier tube (also referred to as a photomultiplier) is placed after one large crystal such as NaI.
It stays at the level of about mm. In addition, since the scintillator performs detection through multi-step conversion from radiation to visible light and from visible light to electrons, it has a problem that the energy resolution is very poor.
Incidentally, some PET devices (positron emission tomography) have a position resolution of 5 to 6 mm and a high-end PET device of about 4 mm. However, since the photomultiplier tube uses a vacuum tube, the position resolution is further increased. It is difficult to do well.

As a radiation detector that detects radiation based on a principle different from such a scintillator, Cd
There is a semiconductor radiation detector provided with a semiconductor radiation detection element using a semiconductor material such as Te (cadmium telluride), TlBr (thallium bromide), and GaAs (gallium arsenide).
In this semiconductor radiation detector, the semiconductor radiation detection element converts the electric charge generated by the interaction between radiation and semiconductor material into an electrical signal. Therefore, the semiconductor radiation detector has a higher conversion efficiency to an electrical signal than a scintillator and can be downsized. Therefore, it is attracting attention.
JP 2003-79614 A (paragraph number 0016) Japanese Patent Laying-Open No. 2003-167058 (paragraph numbers 0020 and 0023)

By the way, in a semiconductor radiation detector, when a semiconductor material such as Tl constituting the semiconductor radiation detection element interacts with radiation, positively charged holes and negatively charged electrons are generated. The mobility of electrons is relatively large, but the mobility of holes is relatively small. That is, electrons are relatively easy to move but holes are difficult to move. For this reason,
In order for the holes to reach the electrode, it takes more time than the electrons reach the electrode. In addition, the holes may disappear on the way to the electrodes. Thereby, there exists a problem that the detection sensitivity of a radiation worsens. Therefore, it is necessary to solve such a problem.
An object of this invention is to provide the semiconductor radiation detector which can improve a detection sensitivity.

In order to solve the above problems, in the first invention, the detection sensitivity is increased by narrowing (shortening) the distance between the electrodes for collecting charges in the semiconductor radiation detector. That is, the distance between the anode electrode and the cathode electrode, or the thickness of the semiconductor region sandwiched between the anode electrode and the cathode electrode is 0.2 to 2 mm. In this configuration, the distance from the position of the electrons and holes generated by the interaction between the semiconductor material and the radiation to the electrode is shortened.
The time until they reach the electrode is shortened. Also, since the distance to the electrode is shortened,
The degree to which the hole disappears in the middle is reduced.
A second invention is a nuclear medicine diagnostic apparatus having a plurality of unit substrates including a plurality of semiconductor radiation detectors that receive radiation and an integrated circuit that processes a radiation detection signal output from each of the plurality of semiconductor radiation detectors. . Since the semiconductor radiation detector and the integrated circuit for processing the output thereof can be arranged close to each other, when the weak output signal of the semiconductor radiation detector is transmitted to the integrated circuit, the influence of noise on the weak output signal can be reduced.

Preferably, semiconductor radiation detector, analog LSI (Large Scale Integration), A
These are arranged on the unit substrate so as to be in the order of the D converter and the digital LSI, the signal detected by the semiconductor radiation detector is processed by the analog LSI, the signal processed by the analog LSI is processed by the AD converter, Each element is connected by wiring so that a signal processed by the AD converter is processed by a digital LSI. In this configuration, the wiring distance between the semiconductor radiation detector and the analog LSI can be shortened by making the distance between the semiconductor radiation detector and the analog LSI close to each other. Noise that is superimposed on the wiring until the processed signal reaches the analog LSI can be reduced. Note that an LSI (integrated circuit) corresponds to an ASIC in an embodiment of the invention described later. Further, in the embodiment of the invention to be described later, the semiconductor radiation detection apparatus is a combined substrate (detector substrate + A
Corresponds to a SIC substrate).
In the second invention, the detection signal when the radiation is detected by the semiconductor radiation detector is ASIC (Applic
New problem discovered by the inventors that it is susceptible to noise because the detection signal output from the semiconductor radiation detector is weak. It was made to solve. The reduction in noise leads to a substantial increase in detection sensitivity (count number, peak value, time detection accuracy) by the semiconductor radiation detector.

Preferably, the substrate on which the semiconductor radiation detector is mounted and the substrate on which the LSI is mounted are different substrates. And, normally, it is used as a combined board (integrated board) that integrates the two boards. If a problem occurs, only the defective board can be replaced, making maintenance and inspection easier. I tried to do it.
In the third invention, a plurality of unit substrates including a semiconductor radiation detector and an integrated circuit are detachably housed in a frame. Since a frame containing a plurality of unit substrates may be attached to the nuclear medicine diagnostic apparatus, a plurality of semiconductor radiation detectors can be attached to the nuclear medicine diagnostic apparatus at a time. For this reason, the time required to attach the semiconductor radiation detector to the nuclear medicine diagnostic apparatus can be remarkably shortened.
Preferably, each of the unit substrates can be removed from the frame body, and the entire frame body can be removed from the nuclear medicine diagnosis apparatus, specifically, the camera, thereby further facilitating maintenance and inspection.
By the way, nuclear medicine diagnostic equipment (radiological diagnostic equipment) such as PET, SPECT, gamma camera
A number of semiconductor radiation detectors are used. For example, with PET, 100,000 to hundreds of thousands (
A large number of semiconductor radiation detectors called “ch” are used, and it is desired to reduce the time required for attaching these many semiconductor radiation detectors to the nuclear medicine diagnostic apparatus. The fourth invention is made to answer such a demand. It is also desirable to facilitate maintenance and inspection of semiconductor radiation detectors.

ADVANTAGE OF THE INVENTION According to this invention, it can prevent thru | or reduce that the detection sensitivity of the radiation in a semiconductor radiation detector worsens. Further, it is possible to prevent or reduce the deterioration of the signal detected by the semiconductor radiation detector. Thereby, for example, a clear image can be obtained in a nuclear medicine diagnostic apparatus.

Embodiment 1
Next, a nuclear medicine diagnosis apparatus which is a preferred embodiment of the present invention will be described in detail with reference to the drawings as appropriate. In the following, the description of the nuclear medicine diagnostic apparatus of the present embodiment and the arrangement of each element such as the distance between electrodes of the semiconductor radiation detector, the analog ASIC, etc. on the substrate (
The elements applied to the present embodiment, such as layout) and substrate unitization, will be described. The analog ASIC is an ASIC (Ap that is an application specific IC that processes analog signals).
replication specific integrated circuit), which is a kind of LSI.

≪Nuclear medicine diagnostic equipment≫
First, the nuclear medicine diagnostic apparatus (radiological medicine diagnostic apparatus) of this embodiment will be described. FIG.
As shown in FIG. 1, a PET apparatus 1 as a nuclear medicine diagnosis apparatus includes a camera (imaging apparatus) 11, a data processing apparatus 12, a display apparatus 13, and the like. The subject is placed on the bed 14 and photographed by the camera 11. The camera 11 incorporates a number of semiconductor radiation detectors 21 (see FIGS. 3, 7, and 10), and γ-rays emitted from the body of the subject are detected by a semiconductor radiation detector (hereinafter simply referred to as a detector). This is detected at 21). The camera 11 is provided with an integrated circuit (ASIC) for measuring the crest value and detection time of the γ-ray, and measures the crest value and detection time of the detected radiation (γ-ray). . Data processor 12
Includes a storage device, a simultaneous measurement device 12A (see FIG. 2), and a tomogram information creation device 12B (see FIG. 2). The data processing device 12 captures packet data including the detected peak value of γ-rays, detection time data, and detector (channel) ID. The simultaneous measuring device 12A performs simultaneous measurement based on the packet data, particularly the detection time data and the detector ID.
The detection position of the KeV gamma ray is specified and stored in the storage device. The tomographic image information creation device 12B creates a functional image based on the specified position and displays it on the display device 13.

As shown in FIG. 2, the inside of the camera 11 detects γ rays emitted from the subject.
A large number of detector units 2 (refer to FIG. 10 for details) each including a plurality of combined substrates 20 (refer to FIG. 7 for details) each having a large number of detectors 21 are arranged circumferentially. The subject lies on the bed 14 and is positioned at the center of the camera 11. At this time, each detector surrounds the bed 14. From the detector unit 2, the γ-ray peak value information and γ-ray detection time information obtained based on the detection signal when the detector 21 interacts with the γ-ray, and the detector 21.
Address information (detector ID) is output for each detector 21 included in the detector unit 2. The configurations of the detector 21, the coupling substrate 20, and the detector unit 2 will be described in detail later. By the way, the subject is a radiopharmaceutical, for example, a half-life of 18 / 110F.
Fluorodioxyglucose (FDG) containing From the body of the subject, γ rays (annihilation γ rays) are emitted when the positrons emitted from the FDG disappear.
Hereafter, the characteristic part of this embodiment is demonstrated.

≪Semiconductor radiation detector; distance between electrodes≫
First, the detector 21 applied to this embodiment will be described. As shown in FIG.
Is a semiconductor radiation detection element (hereinafter referred to as a detection element) 211 made of a plate-shaped semiconductor material S.
Are covered with thin plate-like (film-like) electrodes (anode A, cathode C) (minimum configuration). Among these, the semiconductor material S includes the above-described CdTe (cadmium telluride), Tl.
It is composed of any single crystal such as Br (thallium bromide) or GaAs (gallium arsenide). The electrodes (anode A, cathode C) are made of any material such as Pt (platinum), Au (gold), and In (indium). In the following description, the semiconductor material S is Cd.
Assume that a single crystal of Te is sliced. Further, it is assumed that the radiation to be detected is 511 KeV γ rays used in the PET apparatus.

An outline of the principle of detection of γ rays by the detector 21 will be described with reference to FIG. When γ rays are incident on the detector 21 and an interaction occurs between the γ rays and the semiconductor material S constituting the detector 21, they are schematically indicated by “+” and “−” in the drawing. Holes and electrons are generated in pairs in an amount proportional to the energy of the gamma rays. Here, a voltage for collecting charges is applied between the anode (anode) A and the cathode (cathode) C of the detector 21 (for example, 300 V). For this reason, holes are attracted and moved to the cathode C, and electrons are attracted and moved to the anode A. Comparing holes and electrons, as described in the above section “Disclosure of the invention”, since electrons are relatively larger in mobility, electrons are relatively shorter. Reach the anode in time. On the other hand, since holes are relatively less likely to move, the holes reach the cathode over a relatively long time. Incidentally, the holes may disappear before reaching the electrodes.

“Time−” when the semiconductor material S (detection element 211) of the detector 21 is thick and thin.
As shown in FIG. 4 in which the “crest value curve” is compared, the crest value rises (rises) faster and the maximum crest value is higher in the semiconductor material S having a smaller thickness t. The rapid increase of the peak value contributes to improving the accuracy of simultaneous measurement in PET, for example. In addition, a high peak value contributes to an increase in energy resolution. As the thickness t becomes smaller in this way, the speed at which the crest value rises becomes faster and the crest value becomes higher (the charge collection efficiency is improved) because electrons and holes are used for electrodes (anode A, cathode). C) Time to reach (charge collection time)
Because it becomes shorter. In addition, holes that had previously disappeared in the middle are short distances,
This is because the electrode (cathode C) can be reached without disappearing. By the way, thickness t is facing (
It can also be expressed as the distance between the anode A and the cathode C.

The thickness (interelectrode distance) t of the detection element 211 is preferably 0.2 mm to 2 mm. This is because when the thickness t is 2 mm or more, the rising speed of the crest value becomes slow and the maximum value of the crest value becomes low. On the other hand, when the thickness t is 0.2 mm or less, the thickness (volume) of the electrodes (anode, cathode) is relatively increased, and when installed on the substrate, the ratio of the essential semiconductor material S that interacts with radiation is increased. Because it will decrease. That is, when the thickness t of the semiconductor material S is reduced, the thickness of the electrode that does not interact with γ-rays increases relatively, while the ratio of the semiconductor material S that interacts with γ-rays decreases relatively. As a result, the sensitivity of detecting γ rays is lowered (γ rays pass through). In addition, if the thickness t is small, a large leakage current is generated, which may prevent a high voltage from being applied for charge collection.
For the same purpose, the thickness t of the semiconductor material S is more preferably 0.5 mm to 1.5 mm. With this thickness t, γ rays can be detected more reliably. ,
Measurement of the peak value can be performed more correctly.

Incidentally, in the case of the PET apparatus 1, since simultaneous measurement is performed, it is a problem to correctly measure the detection time of γ rays. For example, in FIG. 3, there is a difference in detection time between the position where the γ-ray and the semiconductor material S interact with each other near the cathode C and near the anode A. That is, since the hole moving speed is slow, the detection time when the interaction occurs near the anode A is relatively late, and the detection time when the interaction occurs near the cathode C is relatively early (true Close to the time). That is, even when γ rays interact with the semiconductor material S in the same detection element 211, there arises a problem that the detection time varies depending on the position where the interaction occurs. Specifically, when the thickness t is thick, the difference in detection time depending on the position where the interaction occurs is increased. Such an event is not a big problem in other fields, but becomes a big problem in the case of the PET apparatus 1 that performs simultaneous measurement (simultaneous counting) on the order of nsec (nanosecond). Therefore, also from this meaning, the detection time can be appropriately determined within the above-described thickness range. The detection time in PET is determined by the LET method or the CFD method.

As shown in FIG. 5 schematically showing the relationship between the thickness t of the semiconductor material S of the detector 21 and the crest value (maximum value), the maximum value of the crest value increases as the thickness t of the semiconductor material S increases. Get smaller. The reason why the peak value becomes small is that, for example, holes disappear before reaching the electrode.
When the thickness t is 2 mm, the detected wave height is below a threshold value that can be distinguished as 511 KeV γ-rays. Therefore, as described above, the thickness t of the semiconductor material S is increased from 2 mm. Is not preferred.

As shown in FIG. 6, the detector 21 includes a semiconductor material S (detection element 211) sandwiched between a cathode C and an anode A and stacked in five layers. Each layer of the semiconductor material S is a single-layer detector 21 having the above-described thickness t (0.2 to 2 mm (more preferably 0.5 to 1.5 mm)). The thickness of anode A and cathode C are each about 20 microns. Incidentally, the detector 21 of the laminated structure shown in FIG.
Since the cathodes C are connected in common, each layer is not configured to detect radiation independently of the other layers. In other words, when the γ-ray and the semiconductor material S interact with each other, it is not determined whether it occurred in the uppermost layer or the lowermost layer. Of course, it can also be set as the structure detected for every layer. Incidentally, the five-layer structure as described above is that when the thickness t of the semiconductor material S is reduced, the rising speed of the crest value and the maximum value of the crest value may be increased, but the thickness t is thin. Because there will be more gamma rays that pass through,
This is to increase the interaction between the semiconductor material S and the γ rays by increasing the charge collection efficiency and reducing the amount of γ rays that pass through (to increase the number of counts).
By adopting the configuration of the detector 21 having such a laminated structure, a better rising speed of the crest value (
Rise) and a more accurate peak value, and interaction with the semiconductor material S
The number of lines (count number) can be increased (sensitivity can be increased).

The area s of the electrodes (anode A, cathode C) is preferably 4 to 120 square mm. Area s
This increases the capacitance of the detector 21 (stray capacitance), and noise increases easily due to the increase in the stray capacitance. Therefore, the electrode area s should be as small as possible. Further, since the charge generated at the time of γ-ray detection is partially stored in the stray capacitance, the analog ASIC 24 increases when the stray capacitance increases.
This causes a problem that the amount of charge stored in the charge amplifier 24b and thus the output voltage (crest value) decreases. When CdTe is used as the detector 21, the relative dielectric constant is 11, and when the area s of the detector 21 is 120 square mm and the thickness t is 1 mm, the capacitance is 12 pF, and the stray capacitance such as the connector of the circuit is reduced. Considering that it is several pF, it cannot be ignored. Therefore, the electrode area s is preferably 120 mm 2 or less.
Further, the lower limit value of the electrode area s is determined from the position resolution of the PET apparatus. The position resolution of the PET apparatus is determined by the range of the detector 21 (arrangement pitch) as well as the range of positrons,
Since the range of the 18F positron is 2 mm, it is meaningless to set the size of the detector 21 to 2 mm or less. The mounting method with the smallest electrode area is when the electrode surface is placed perpendicular to the radial direction of the camera 11, and the lower limit of one side of the electrode is 2 mm from the above consideration, and the electrode area s
The lower limit of is 4 square mm.

In the above description, the semiconductor material S that interacts with γ-rays is CdTe, but it goes without saying that the semiconductor material S may be TlBr, GaAs, or the like. In addition, the terms “laminated structure” and “upper layer / lower layer” are used. These terms are based on FIG.
In the case of 0 ° sideways, the laminated structure can be read as, for example, a parallel structure, and the top and bottom are, for example, left and right. Further, the direction in which the γ-rays are incident may be upward / downward, leftward / rightward in FIG. In other words, the detector 21 has a structure in which a plurality (for example, five) of semiconductor materials S are arranged in parallel with the cathode C and the anode A being alternately sandwiched therebetween.

<< Coupling board; Detector board and ASIC board >>
A detailed structure of the coupling substrate (unit substrate) 20 installed in the detector unit 2 (FIG. 10) will be described with reference to FIG. The coupling substrate (semiconductor radiation detection apparatus) 20 includes a plurality of detectors 2.
1, a detector board (first board) 20A, a capacitor 22, a resistor 23, an analog ASIC 24, an analog / digital converter (hereinafter referred to as ADC) 25, and a digital ASIC 26 (second board). Substrate) 20B.

(Detector board)
With reference to FIG. 7, the detector board | substrate 20A which installed the detector 21 is demonstrated. As shown in FIG. 7A, the detector substrate 20A has a plurality of detectors 21 arranged (installed) in a grid on one surface of the substrate body 20a (16 detections in a row). 4 rows of vessel 21 = 16 horizontal x 4 vertical
64 in total). In the radial direction of the camera 11, the detectors 21 are arranged in four rows on the substrate body 20a. The 16 horizontal detectors 21 are arranged in the axial direction of the camera 11, that is, in the longitudinal direction of the bed 14. Further, as shown in FIG. 7B, the detector substrate 20
Since the semiconductor radiation detectors 21 are installed on both sides of A, a total of 128 detectors 21 are installed on one detector substrate 20A. Here, the detector 21 to be installed
As the number of γ increases, it becomes easier to detect γ-rays, and the positional accuracy during γ-ray detection can be increased. For this reason, the detector 21 is installed on the detector substrate 20A as closely as possible. Incidentally, in FIG. 7A, when γ rays emitted from the subject on the bed 14 travel from the lower side of the drawing upward (in the direction of the arrow 32, that is, the radial direction of the camera 11), the detector Board 2
It is preferable to densely arrange the detectors 21 in the left-right direction at 0A because the number of passing γ rays (the number of γ rays passing through the gap between the detectors 21) can be reduced. This increases the detection efficiency of γ rays, and can increase the spatial resolution of the obtained image.
In the detector substrate 20A of the present embodiment, as shown in FIG. 7B, the detector 21 is installed on both sides of the substrate body 20a. 20a can be shared by mounting on both sides. For this reason, the number of substrate bodies 20a can be halved, and the detectors 21 can be arranged more densely in the circumferential direction. In addition, as described above, the number of detector substrates 20A (combined substrates 20) can be reduced by half, so that it is possible to save time and labor for mounting the combined substrates 20 on a case 30 (see FIG. 10) described later. There are also benefits.
In the above description, the 16 horizontal detectors 21 are arranged in the axial direction of the camera 11, but the present invention is not limited to this. For example, 16 horizontal detectors 21 may be arranged in the circumferential direction of the camera 11.

As shown in FIG. 7C, each detector 21 has a laminated structure in which single crystals of the thin plate-like semiconductor material S (detecting element 211) are laminated. The configuration and operation of this point are as already described with reference to FIG. 6, but will be supplementarily described here. The detector 21 is provided with the anode A and the cathode C as described above, and a potential difference (voltage) of, for example, 300 V is set between the anode A and the cathode C in order to collect charges. This voltage is ASI
It is supplied from the C substrate 20B side to the detector substrate 20A side via the connector C1 (FIG. 7A). The signal detected by each detector 21 is supplied to the ASIC board 20B side via the connector C1. For this reason, in the board main body 20a of the detector board 20A, a board wiring (for charge collection / signal exchange) (not shown) for connecting the connector C1 and each detector 21 is provided. The in-substrate wiring has a multilayer structure. In the present embodiment, each detection element 211 of the detector 21 is arranged in parallel to the substrate body 20a. However, each detection element 211 is connected to the substrate body 2.
The detector 21 may be provided so as to be perpendicular to 0a.

(ASIC substrate)
Next, the ASIC substrate 20B on which the ASIC is mounted will be described. As shown in FIG.
The ASIC board 20B is provided with two analog ASICs 24 and one digital ASIC 26 on one side of the board body 20b. In addition, as shown in FIG.
Since C24 is installed on both sides of the substrate body 20b, one ASIC substrate 20B has a total of four analog ASICs 24. Further, the ASIC substrate 20B has eight ADCs 25 (= 4 × 2) on one side of the substrate main body 20b and 16 on both sides. Also, 1
The number of capacitors 22 and resistors 23 corresponding to the number of detectors 21 are provided on both surfaces of one substrate body 20b. These capacitors 22, resistors 23, analog ASI
In order to electrically connect the C24, the ADC 25, and the digital ASIC 26, the ASIC board 20
In the B (substrate body 20b), an in-substrate wiring (not shown) is provided in the same manner as the detector substrate 20A. This intra-substrate wiring also has a laminated structure.
The arrangement of these elements 22, 23, 24, 25, 26 (in-substrate wiring)
The signal supplied from 0A is the capacitor 22, resistor 23, analog ASIC 24, ADC
25 and digital ASIC 26 are supplied in this order.
The ASIC board 20B is a connector (spiral contact) C1 that is connected to the in-board wiring connected to each capacitor 22 to make electrical connection with the detector board 20A.
And a board connector C2 for electrical connection to the data processing device side (unit integrated FPGA side described later). Incidentally, the above-described detector board 20 </ b> A also has a connector C <b> 1 connected to the in-board wiring connected to each detector 21.

(Connection structure of detector board and ASIC board)
A connection structure between the detector substrate 20A and the ASIC substrate 20B will be described.
The detector substrate 20A and the ASIC substrate 20B are not connected by abutting the end surfaces (end portions) thereof, but as shown in FIG. The connectors C1 existing in the overlap portion are connected to each other. This connection is
It is detachable (separate / connectable) with a fastening screw or the like. This connection is made for the following reason. That is, when the combined substrate 20 in which the detector substrate 20A and the ASIC substrate 20B are connected (coupled) is supported in one end (cantilever support) or both ends in the horizontal direction, the central portion (connection portion) of the combined substrate 20 is provided. A force that bends or bends the bonding substrate 20 downward acts. Here, it is not preferable that the connection portion is formed by abutting the end faces because the connection portion is easily bent or bent.

In view of this point, in the present embodiment, the detector substrate 20A and the ASIC substrate 20B are not connected with the end surfaces attached to each other, but as described above, the overlap portion is formed so as to overlap the vicinity of the end portion. Provided and connected. For this reason, since the toughness with respect to a bending | flexion and a bending improves compared with connecting face-to-face, it is preferable. In addition, when the toughness with respect to a bending | flexion and a bending of a coupling board | substrate improves, the position shift of the detector 21 will be suppressed, for example, and the fall of the precision which pinpoints the generation | occurrence | production position of a gamma ray will be prevented. Incidentally, as shown in FIG. 2, the camera 11 of the PET apparatus 1 has a large number of detector units 2 (FIG. 10) including the coupling substrate 20 shown in FIG. 3 o'clock direction corresponding to the horizontal direction
The coupling substrate 20 located in the direction of time is easily bent or bent. For this reason, the toughness with respect to the bending and bending of the coupling substrate 20 is important.

The detector substrate 20A and the ASIC substrate 20B are electrically connected using the overlap portion described above. For this reason, the detector substrate 20A and the ASIC substrate 20 shown in FIG.
Each overlap portion of B is provided with a connector (connector) C1 (FIG. 7A) for electrically connecting the in-board wirings of both the boards 20A and 20B. For example, a spiral contact (registered trademark) is used as the connector C1 in order to improve electrical connection. A spiral contact (registered trademark) has a characteristic that a ball-shaped connection terminal contacts a spiral contactor over a wide area to achieve good electrical connection. When the ball-shaped connection terminal is provided on the ASIC substrate 20B side, the spiral contact is provided on the detector substrate 20A side, and when the ball-shaped connection terminal is provided on the detector substrate 20A side, The spiral contact is provided on the ASIC substrate 20B side.
By using such an electrical connection structure between the detector board 20A and the ASIC board 20B, a signal can be transmitted from the detector board 20A to the ASIC board 20B with low loss. Incidentally, when the loss is reduced, for example, the energy resolution of the detector 21 is improved.

Further, as described above, the connection between the detector substrate 20A and the ASIC substrate 20B is a detachable connection using screws or the like. For this reason, for example, the semiconductor radiation detector 21 and the ASI
Even if a defect occurs in C24, 26, only the defective part needs to be replaced. Therefore, it is possible to eliminate waste such as replacement of the entire combined substrate 20 due to some defects. The electrical connection between the detector substrate 20A and the ASIC substrate 20B is as follows:
Since it is performed by the connector C1 such as the spiral contactor (registered trademark) described above, it is easy to connect / release the connection (connection / release of connection) between the substrates.

In the above configuration, one detector board 20A is connected to the ASIC board 20B, but the detector board may be divided into a plurality of parts. For example, eight detectors 21 in the horizontal direction and four detectors 21 in the vertical direction are 1
A configuration may be adopted in which two detector boards are connected to the ASIC board. In this configuration, when one detector 21 fails, it is only necessary to replace the detector substrate on which two of the failed detectors are mounted, thereby reducing waste during maintenance (cost reduction). .

(Element layout)
Next, the layout of the elements such as the detector 21 and the ASICs 24 and 26 on the combined substrate 20 will be described with reference to FIGS.

As shown in FIG. 8, the detector 21 is connected to the analog ASIC 24 via a connector C1, a capacitor 22 and a resistor 23 by electrical wiring (not shown), and a detection signal of γ rays detected by the detector 21 is Through the electric wiring, it passes through the capacitor 22 and the resistor 23 and is processed by the analog ASIC 24. A signal processed by the analog ASIC 24 is processed by the ADC 25 and the digital ASIC 26.

Here, it is preferable that the length of the circuit and the length (distance) of the wiring are short because there is less influence of noise and signal attenuation during the process. Moreover, when performing simultaneous measurement processing with the PET apparatus 1, it is preferable that the length of the circuit or wiring is short because the time delay is small (preferably because the accuracy of the detection time is not impaired). For this reason, in this embodiment, as shown in FIG. 7A from the central axis of the camera 11 toward the outside in the radial direction of the camera 11, the detector 21 and the capacitor 2
2, the resistor 23, the analog ASIC 24, the ADC 25, and the digital ASIC 26 are arranged (laid out) in the order of the elements 21, 22, 23, 24, 25, and 26, respectively. This order is the same as the signal processing order by the elements 21, 22, 23, 24, 25, and 26 (FIGS. 8 and 9).
reference). That is, from the central axis of the camera 11 toward the outside, “the detector, the analog integrated circuit, the AD converter, and the digital integrated circuit are arranged on the substrate in this order and wired in this order”. For this reason, a weak signal detected by the detector 21 can be transmitted to the analog ASIC 24 with the wiring length (distance) shortened.
Since processing such as signal amplification is performed by the analog ASIC 24, even if the wiring length after the analog ASIC 24 is long, it is not easily affected by noise. In other words, in view of noise, there is no problem even if the wiring length after the analog ASIC 24 is long. However, since the signal transmission is delayed when the length of the wiring is long, the accuracy of the detection time may be impaired as described above.

Incidentally, in this embodiment, since not only the detector 21 but also the analog ASIC 24 and the digital ASIC 26 are included in one coupling substrate 20, the detector 21, the analog ASIC
24 and the digital ASIC 26 can be arranged in the longitudinal direction of the bed 14, that is, in the direction orthogonal to the body axis of the subject to be inspected, so that the length of the camera (imaging device) 11 in the longitudinal direction of the bed is longer than necessary. You don't have to. Analog ASIC 24 and digital ASI
Although it is conceivable to arrange the C26 in the annular direction in the longitudinal direction of the bed 14 on the outer side in the radial direction, the length of the camera 11 in the longitudinal direction of the bed becomes longer than necessary. Further, a semiconductor radiation detector is used as the detector 21 and an analog AS is used as the signal processing device.
Since the IC 24 and the digital ASIC 26 are used, the length of the coupling substrate 20 in the longitudinal direction is shortened, and the length of the camera 11 in the orthogonal direction can be significantly reduced as compared with the case where a scintillator is used. Furthermore, the coupling substrate 20 has a detector 21 and an analog A in the longitudinal direction.
Since the SIC 24 and the digital ASIC 26 are sequentially arranged, the length of the wiring connecting them can be shortened, and the wiring on the substrate is simplified.

Here, in the present embodiment, one analog ASIC 24 includes 32 detectors 2 each.
1 and processes the signal obtained from the detector 21. As shown in FIG. 8 and FIG. 9, one analog ASIC 24 includes 32 sets of analog signal processing circuits (analog signal processing devices) 33 having a slow system and a fast system. The analog signal processing circuit 33 is provided for each detector 21, and one analog signal processing circuit 33 is connected to one detector 21. Here, the first system has a timing pick-off circuit 24a that outputs a timing signal for specifying the detection time of γ rays. The slow system has a polarity amplifier (linear amplifier) 24c, a bandpass filter (waveform shaping device) 24d, and a peak hold circuit (peak value holding device) 24e for the purpose of obtaining the detected peak value of γ-rays. Connected in order. Incidentally, the slow system is named “slow” because it takes a certain amount of processing time to obtain the peak value. In addition, the code | symbol 24b is
It is a charge amplifier (preamplifier). Incidentally, the γ-ray detection signal output from the detector 21 and passed through the capacitor 22 and the resistor 23 is a charge amplifier 24b, a polarity amplifier 24c.
It is amplified by. The amplified γ-ray detection signal is input to the peak hold circuit 24e through the band pass filter 24d. The peak hold circuit 24e holds the maximum value of the detection signal, that is, the peak value of the γ-ray detection signal proportional to the detected γ-ray energy. 1 analog AS
The IC 24 is an LSI formed from 32 analog signal processing circuits 33.
Although the capacitor 22 and the resistor 23 can be provided inside the analog ASIC 24, the present embodiment is used for obtaining an appropriate capacitor capacity and an appropriate resistance value, and for reducing the size of the analog ASIC 24. In this case, the capacitor 22 and the resistor 23 are arranged outside the analog ASIC 24. Incidentally, it is said that the capacitor 22 and the resistor 23 are less provided with less variation in individual capacitor capacity and resistance value when provided externally.

In the present embodiment, the analog ASIC 24 shown in FIG.
The output of the slow system 4 is supplied to an ADC (analog / digital converter) 25. Furthermore, the first output of the analog ASIC 24 is the digital ASIC 26.
To be supplied.

The analog ASIC 24 and each ADC 25 are connected to each other by a single wiring that collectively transmits slow signals for 8 channels. Each analog ASIC 24 and digital ASI
C26 is connected by 32 wires that transmit 32ch fast signals one by one. That is, one digital ASIC 26 is combined with four analog ASICs 24 in total.
They are connected by 8 wires.

The slow output signal output from the analog ASIC 24 is an analog peak value (the maximum value in the graph shown in FIG. 4). Also, from analog ASIC 24 to digital A
The fast output signal output to the SIC is a timing signal indicating the timing corresponding to the detection time. Among these, the peak value, which is the output of the slow system, is input to the ADC 25 via a wiring (wiring with the above-mentioned 8ch as one) connecting the analog ASIC 24 and the ADC 25, and converted into a digital signal by the ADC 25. In the ADC 25, for example, the peak value is converted into a digital peak value of 8 bits (0 to 255) (ex. 511 KeV → 255).
The timing signal, which is a slow output, is supplied to the digital ASIC 26 through the wiring connecting the analog ASIC 24 and the digital ASIC 26.

The ADC 25 digitizes and transmits 8-bit peak value information to the digital ASIC 26. For this reason, each ADC 25 and the digital ASIC 26 are connected by wiring. Incidentally, since the digital ASIC 26 has 16 ADCs 25 on both sides, it is connected to the ADC 25 by a total of 16 wires. One ADC 25 processes signals for 8 channels (signals for 8 detection elements). The ADC 25 is connected to the digital ASIC 26 by one wiring for ADC control signal transmission and one wiring for peak value information transmission.

As shown in FIG. 9, the digital ASIC 26 includes a plurality of packet data generation devices 34 including eight time determination circuits (time information generation devices) 35 and one ADC control circuit (ADC control device) 36, and data transfer. Circuit (data transmission device) 37, which includes LSI
It has become. All the digital ASICs 26 provided in the PET apparatus 1 receive clock signals from a 500 MHz clock generator (crystal oscillator) (not shown) and operate in synchronization. The clock signal input to each digital ASIC 26 is input to each time determination circuit 35 in all packet data generation devices 34. The time determination circuit 35 is provided for each detector 21 and receives a timing signal from the timing pick-off circuit 24 a of the corresponding analog signal processing circuit 33. The time determination circuit 35 determines the detection time of the γ ray based on the clock signal when the timing signal is input. Since the timing signal is based on the first signal of the analog ASIC 24, a time close to the true detection time can be set as the detection time (time information). The ADC control circuit 36 includes a time determination circuit 35.
The detector ID is identified by receiving the timing signal from which the γ-ray is detected. That is, AD
The C control circuit 36 stores a detector ID for each time determination circuit 35 connected to the ADC control circuit 36. When time information is input from a certain time determination circuit 35, the C control circuit 36 corresponds to the time determination circuit 35. The detector ID can be specified. This is possible because the time determination circuit 35 is provided for each detector 21. Further, after inputting the time information, the ADC control circuit 36 outputs an ADC control signal including the detector ID information to the ADC 25. The ADC 25 converts the peak value information output from the peak hold circuit 24e of the analog signal processing circuit 33 corresponding to the detector ID into a digital signal and outputs the digital signal. This peak value information is input to the ADC control circuit 36. The ADC control circuit 36 adds the peak value information to the time information and the detector ID to generate packet data. The ADC control circuit 36 is an AD that controls the ADC 25.
It has functions of a C control device and an information integration device that integrates detector ID information (detector position information), the time information, and peak value information. The information integration device outputs integrated information (packet information) which is digital information including these three pieces of information. AD of each packet data generator 34
The packet data (including detector ID, time information, and peak value information) output from the C control circuit 36 is input to the data transfer circuit 38.
For example, the data transfer circuit 38 and the packet data, which is digital information output from the ADC control circuit 36 of each packet data generator 34, are periodically detected, for example, in the detector unit 2 (FIG. 10) containing 12 coupled substrates 20. 11) is transmitted to a unit integration integrated circuit (unit integrated FPGA (Field Programmable Gate allay)) 31 provided in one case 30 of FIG. The unit integrated FPGA (hereinafter referred to as “FPGA”) 31 transmits the digital information to the data processing device 12 via the information transmission wiring connected to the connector 38.
In the present embodiment, the ADC 25 converts the peak value information output from the peak hold circuit 24e corresponding to the detector ID information included in the control signal output from the ADC control circuit 36 into a digital signal. One ADC 25 is provided for a plurality of analog signal processing circuits 33 in the analog ASIC 24. Therefore, one analog signal processing circuit 33
Therefore, it is not necessary to provide the ADCs 25 one by one, and the circuit configuration of the ASIC substrate 20B can be greatly simplified. One information integration device for generating integrated information may be provided for a plurality of analog signal processing circuits 33 in one analog ASIC 24, and the circuit configuration of the digital ASIC 26 can be simplified. In addition, the ADC controller for specifying the detector ID is also one analog A
One is sufficient for a plurality of analog signal processing circuits 33 in the SIC 24, and the digital ASIC is used.
26 circuit configurations can be simplified.

Thus, the packet data including (1) peak value information, (2) determined time information, and (3) a detector ID that uniquely identifies each of the detectors 21 output from the digital ASIC 26 is as follows. The data is transmitted to the subsequent data processing device 12 (see FIG. 1) via the information transmission wiring. Based on the packet data transmitted from the digital ASIC 26, the simultaneous measurement device 12A of the data processing device 12 performs simultaneous measurement processing (when two γ rays of a predetermined energy are detected in a time window of a set time, these γ rays Of the two detectors 21 that detect the pair of γ-rays by counting the pair of γ-rays measured at the same time as one. The position is specified by their detector ID. The data processing device 12 has three or more γ-ray detection signals detected within the time window (the detector 21 that detects γ-rays).
If there are three or more detectors), the two detectors 21 to which γ rays are incident first among the three or more detectors 21 are specified using the peak value information of the γ ray detection signals. The specified pair of detectors 21 are simultaneously measured to generate one count value. Further, the tomogram information creation device 12B of the data processing device 12 uses the count value obtained by the simultaneous measurement and the position information of the detector 21 to
The tomographic image information of the subject at the radiopharmaceutical accumulation position, that is, the position of the malignant tumor is created. This tomographic image information is displayed on the display device 13. Information such as the digital information, the count value obtained by simultaneous measurement, the position information of the detector 21, and the tomographic image information is stored in the storage device of the data processing device 12.

In the above description, the detector board 20A includes the detector 21, and the ASIC board 20B includes the capacitor 22, the resistor 23, the analog ASIC 24, the ADC 25, and the digital ASIC 2.
6. However, the detector board (first board) 20A includes the detector 21, the capacitor 22, the resistor 23, the analog ASIC 24, and the like, and the ASIC board (second board) 20B includes the ADC 25, the digital ASIC 26, and the like. May be. The detector substrate 20A includes the detector 21 and the analog ASIC 24, so that the detector 21 and the analog ASIC 24 are included.
(Distance of wiring) can be further shortened. For this reason, it becomes possible to further reduce the influence of noise.
Further, the combined substrate 20 includes three substrates (detector substrate 20A, analog ASIC substrate, digital A
The SIC substrate may be configured to be detachably connected via connectors.
In this case, among the three boards, the detector board 20A has a detector 21, the analog ASIC board has a capacitor 22, a resistor 23, and an analog ASIC 24, and the digital ASIC board has an ADC 25 and a digital ASIC 26. To do. In this configuration, by separating the board on which the analog circuit is mounted from the board on which the digital circuit is mounted, noise on the digital circuit side is prevented from entering the analog circuit. In this configuration, the analog ASIC and digital ASIC mounting boards are separated and connected by a detachable connector.
For example, even if only the digital ASIC fails, only the digital ASIC board needs to be replaced. Therefore, this configuration can further eliminate waste.

Incidentally, in the above description, the substrate body 20a (detector substrate 20) on which the detector 21 is installed.
A) and the substrate body 20b (ASIC substrate 20B) on which the ASICs 24 and 26 are installed are different substrates. For this reason, for example, when both ASICs are soldered to the substrate via BGA (Ball Grid Allay) by reflow, only the ASIC substrate can be soldered, which is preferable because the semiconductor radiation detector 21 does not need to be exposed to a high temperature. . Of course, all elements 21
It is also possible to avoid using the connector C1 by arranging ~ 26 on the same substrate.

≪Detector unit; unitization by storing the combined substrate≫
Next, unitization by storing the combined substrate 20 in the housing 30 will be described. In the present embodiment, twelve coupled substrates 20 are accommodated in a housing (frame body) 30 and a detector unit (12
Substrate unit) 2 is configured. Incidentally, the camera 11 of the PET apparatus 1 has a configuration in which 60 to 70 detector units 2 are detachably arranged in the circumferential direction (FIG. 12B).
(See Fig. 2).

(Storing in a housing)
As shown in FIG. 10, the detector unit 2 includes twelve combined substrates 20,
A high-voltage power supply PS for supplying a voltage for collecting charges to a single coupling substrate 20, the aforementioned FPGA 31,
A signal connector for exchanging signals with the outside, a housing 30 for housing and holding a power connector for receiving power supply from the outside, and the like are provided.

As shown in FIGS. 10 and 11, the combined substrate 20 has a depth direction (longitudinal direction of the bed 14).
3 rows so that they do not overlap with each other, and 4 rows are arranged in the frontage direction (circumferential direction of the camera 11).
It is contained within 0. That is, 12 coupling substrates 20 are accommodated in one housing 30. In order to store in this way, one guide groove (guide rail) G1 extending in the depth direction
Are arranged in the housing 30 and are attached to the upper end of the housing (lid) 30. The guide member 39 has an opening 40 at a position facing each connector C3 of the top plate 30a in each guide groove G1. Further, four guide members 41 having one guide groove (guide rail) G2 extending in the depth direction are provided on the bottom plate 30b of the housing 30 so as to be appropriately separated in the circumferential direction (see FIG. 11). Guide grooves G1, G2
Has a depth enough to accommodate three (three) combined substrates 20. ASI of bonded substrate 20
The end portion on the C substrate 20B side is accommodated in the guide groove G1, and the end portion on the detector substrate 20A side of the coupling substrate 20 is accommodated in the guide groove G2. Three combined substrates 20 are held in the depth direction of the guide grooves G1, G2. Incidentally, the coupling substrate 20 is configured such that the end portion on the ASIC substrate 20B side and the end portion on the detector substrate 20A side slide in the guide grooves G1, G2.
The coupling substrate 20 can be easily positioned at a predetermined position by sliding the inside of the guide grooves G1 and G2 with a finger or the like. At this time, each board connector C2 is located in the opening 40 portion. After the predetermined number of coupled substrates 20 are arranged in the housing 30, the top plate 30a is detachably attached to the upper end of the housing 30 with screws or the like. Each connector C3 provided on the top plate 30a is inserted into the corresponding opening 40 and connected to the corresponding board connector C2. The housing 30
The upper part and the lower part are when the casing 30 is taken out from the camera 11, and as shown in FIG. 2, when the casing 30 is provided in the camera 11, the top and bottom are inverted, 9 up and down
Rotate 0 ° and turn left or right or tilt.

As shown in FIG. 11, the top plate 30a of the housing 30 is provided with an FPGA 31 and a connector 38 in addition to the above-described four rows of guide grooves G1. Connector 38 is FPGA3
1 is connected. The FPGA 31 can create a program on site. This is different from the ASIC that cannot be programmed. Therefore, as in this embodiment, the FPG
In A31, for example, even when the number or type of the combined substrates 20 to be stored is changed, it is possible to appropriately cope with the change in the number of pieces by programming at the site.

Since the detector 21 using CdTe as the semiconductor material S used in the present embodiment generates charges in response to light, the housing 30 is made of a light-shielding material such as aluminum or an aluminum alloy. In addition, there is no gap for light to enter. That is, the housing 30 has a light shielding property. Incidentally, when the light shielding property is ensured by other means, the housing 30 does not need to have the light shielding property itself, and may be a frame (frame body) that detachably holds the detector 21 (for example, for light shielding). The face material (panel) etc. is unnecessary).

As shown in FIG. 12A, the detector unit 2 is mounted via a unit support member 2A. As shown in FIG. 12B, the detector unit 2 is mounted on the camera 11 with one end supported by the unit support member 2A. The unit support member 2A has a hollow disk shape (donut shape), and includes a large number of windows in the circumferential direction of the camera 11 (as many as the number of detector units 2 to be mounted). Thus, in order to support the detector unit 2 at one end, a flange portion serving as a stopper is provided on the front side of the housing 30 of the detector unit 2 in the body axis direction. Incidentally, when trying to arrange the detector units 2 as closely as possible in the circumferential direction, the flange portion on the inner side in the circumferential direction becomes an obstacle. In view of this, the flange portion, which is a hindrance, may be eliminated from the housing 30 and the flange portion on the outer side in the circumferential direction may be left.
Another unit support member 2A may be installed, and both ends of the detector unit 2 may be held by both unit support members 2A.
In the present embodiment, since the detector unit 2 is attached to the unit support member 2A, a large number of detectors 21 can be attached to the camera 11 at a time. For this reason, the attachment time to the camera 11 of the detector 21 can be remarkably shortened. Further, the packet data (all packet data for all detectors 21 of the combined substrate 20) output from the data transfer device 38 of all the combined substrates 20 in the detection unit 2 is data from the unit integrated FPGA 31 provided in the detection unit 2. It is sent to the processing device 12. For this reason, the number of wires for transmitting packet data to the data processing device 12 in the present embodiment is remarkably reduced as compared with the case of transmitting packet data from each data transfer device 38 of the combined substrate 20 to the data processing device 12. Is done.

When the detector unit 2 is attached to the camera 11, the lid 11a is removed to expose the unit support member 2A, and the detector unit 2 is inserted and attached until the flange portion hits. . By inserting and attaching, the connector of the camera 11 and the detector unit 2 is connected, and the signal and power supply of the camera 11 and the detector unit 2 are connected.

(Power supply)
Next, the high-voltage power supply device PS that supplies a voltage for collecting charges will be described. As shown in FIG. 10, the detector unit 2 is provided with a high-voltage power supply device PS that supplies a voltage for collecting charges to each detector 21 in a space formed in the housing 30 on the back surface side of the FPGA 31. . The high-voltage power supply device PS is supplied with a low-voltage power supply, boosts the voltage to 300 V by a DC-DC converter (means for boosting the voltage) (not shown), and supplies the boosted voltage to each detector 21. Incidentally, 64 detectors 21 on one side and 128 on both sides are provided for one coupling substrate 20 (= detector substrate 20A). And this coupling board | substrate 20 is one housing | casing 30 (
That is, 12 detectors are stored in one detector unit 2). Therefore, a voltage is supplied to 128 × 12 = 1536 detectors 21 from the high-voltage power supply device PS.

By the way, in the past, 300V with very little fluctuations from a precision power supply located far away.
(1) When the distance from the precision power supply is increased, a wider insulation structure for high-voltage wiring is required (the insulation distance becomes longer).
(2) Since the voltage fluctuates due to temperature fluctuations of the detector 21, even if a precise voltage is supplied from the precision power supply device, the target detector 21 is as accurate as expected. There were problems such as not being voltage.
In order to facilitate maintenance and inspection, for example, the detector unit 2 as in this embodiment is provided with a power connector (not shown) so that the high-voltage power line extending from the precision power device can be removed at the power connector portion. Can be considered. That is, in the present embodiment, it is also conceivable to supply high voltage power to the detector unit 2 from the outside of the unit 2 via the power connector. However, in the case of a high voltage of 300 V, there is a problem that the power supply connector becomes large in addition to the insulation problem.

In the present embodiment, the high-voltage power supply device PS built in the detector unit 2 is connected to the external low voltage (5 by the power supply wiring via the power supply connector 42 and the connector 38 provided on the top board 30a.
To 15 VDC). The high voltage side terminal of the high voltage power supply device PS is the top plate 30
A connector 43 provided on a is connected to twelve connectors C3 provided on the top plate 30a, respectively. Connector C2 of each coupling substrate 20, power supply wiring (not shown) in the substrate body 20b, connector C1 And the substrate body 20 via a power supply wiring (not shown) in the substrate body 20a.
A is connected to the electrode C of each detector 21 provided in a. The connectors C1 and C2 include a connector for power supply wiring in addition to the connector for transmitting the output signal of the detector 21. Since the high voltage power supply PS boosts the low voltage applied from the direct current power source to 300 V by the DC-DC converter, the high voltage portion can be reduced. As a result, the insulation distance could be shortened. That is, it is no longer necessary to use high voltage wiring from the connector 42 to the DC power source. In addition, maintenance and the like have become easier. In addition, with respect to the problem of voltage fluctuation, in the present embodiment, not the precision power supply apparatus, but the high-voltage power supply apparatus P having an accuracy corresponding to the voltage fluctuation due to temperature.
S was provided. This eliminates the need for a highly accurate power supply. Further, since it is a low voltage that is supplied from an external power source, a small power connector can be used as the power connector provided in the connector 38. And since the power connector is small, the degree of freedom in layout is increased. Further, since the high-voltage power supply device PS is disposed in the space formed in the housing 30 on the back surface side of the FPGA 31, the detector unit 2 is increased in size by the placement of the high-voltage power supply device PS in the housing 30. There will be no compactness. Note that the high-voltage power supply PS may be directly connected to the power supply wiring provided on the substrate body 20a via a connector, not via the top board 30a. Further, the power connector may be arranged separately from the output signal connector of the detector 21. Thereby, it is possible to prevent noise from being applied to the signal wiring from the power supply system.

In addition, as described above, by reducing the power supply to the detector unit 2, for example, each ASIC
Similar to the supply of power to 24 and 26, the high-voltage power supply PS can be supplied with a low voltage via the unit integrated FPGA 31.
Further, by supplying power using the high-voltage power supply device PS, insulation from the housing (GND) becomes unnecessary.

By the way, the voltage supplied from the FPGA 31 to the high voltage power supply PS is the high voltage power supply PS.
The voltage is boosted to 300 V by a DC-DC converter (not shown), and after being boosted, it passes through the top plate 30a of the housing 30 and is supplied from the ASIC substrate 20B to the detector substrate 20A to each detector 21 together with the combined substrate 20. Is done. That is, the housing 30 (top plate 30a) includes a voltage supply wiring (not shown) that supplies a voltage from the high-voltage power supply device PS to each coupling substrate 20. In addition, each coupling substrate 2
0 represents the voltage supplied from the high-voltage power supply device PS via the board connector C2 to each detector 21.
Voltage supply wiring to be provided.

[Embodiment 2]
A nuclear medicine diagnostic apparatus according to another embodiment will be described. The nuclear medicine diagnostic apparatus of this embodiment is a single photon emission tomography apparatus (SPECT (Single Photon Emission Computer Tomography) apparatus).

The SPECT apparatus 51 will be described with reference to FIGS. The SPECT device 51
A pair of radiation detection blocks 52, a rotation support base (rotating body) 57, a data processing device 12A, and a display device 13 are provided. These radiation detection blocks 52 are arranged on the rotation support base 57 at positions shifted by 180 ° in the circumferential direction. Specifically, each unit support member 56 of each radiation detection block 52 is attached to the rotation support table 57 at a position 180 degrees apart in the circumferential direction. A plurality of detector units 2 </ b> A including 12 coupling substrates 53 are detachably attached to the respective unit support members 56. Therefore, the detector 21 is held by the unit support member. The configuration of each detector unit 2 </ b> A is the same as the configuration of the detector unit 2 in Embodiment 1 except for the configuration of the coupling substrate 53.

The coupling substrate 53 is similar to the above-described coupling substrate 20 in the detector substrate 20A and the ASIC substrate 5.
3B (FIG. 14). The detector 21 located at the tip of each detector substrate 20A is the bed 1
Located on the 4th side. Collimator 5 made of radiation shielding material (eg lead, tungsten, etc.)
5 is provided in each radiation detection block 52. Each collimator 55 forms a large number of radiation paths through which radiation (for example, γ rays) passes. These radiation paths are provided for each detector 21 located at the tip of all detector substrates 20A of one radiation detection block 52.
Are provided in a one-to-one correspondence. All the coupling substrates 53 and the collimator 55 are provided on the rotary support 5
7 is disposed in the light shielding / electromagnetic shield 54 installed in the apparatus 7. The collimator 55 is attached to the light shielding / electromagnetic shield 54. The light shielding / electromagnetic shield 54 is a detector 21 for electromagnetic waves other than γ rays.
The influence on etc. is cut off.

The bed 14 on which the subject to whom the radiopharmaceutical has been administered is moved, and the subject is moved between the pair of radiation detection blocks 52. By rotating the rotation support base 56,
In each radiation detection block 52, the unit 11 turns around the subject. Gamma rays emitted from an accumulation part (for example, an affected part) C in the subject on which the radiopharmaceutical is accumulated enter the corresponding detector 21 through the radiation path of the collimator 55. The detector 21 outputs a γ-ray detection signal.
This γ-ray detection signal is processed by an analog ASIC 24A and a digital ASIC 26A which will be described later.

Since the configuration of the detector substrate 20A used in this embodiment (Embodiment 2) is the same as that in Embodiment 1, the description in this embodiment is omitted. A constituting the bonding substrate 53
The SIC substrate 53B will be described with reference to FIGS. The ASIC board 53B connected to the detector board 20A by the connector C1 similarly to the coupling board 20 includes a capacitor 22 and a resistor 23 provided for each detector 21, four analog ASICs 24A, and one digital ASIC 26A. Have.

The analog ASIC 24A includes 32 sets of analog signal processing circuits (analog signal processing devices) 33A each having a slow system and a fast system. The analog signal processing circuit 33A is provided for each detector 21. Here, the first system includes a trigger output circuit 24f that outputs a trigger signal that specifies detection of γ rays. The slow system is similar to the analog ASIC 24 in that the charge amplifier 24b, the polarity amplifier 24c,
A band pass filter 24d and a peak hold circuit 24e are provided connected in this order. One analog ASIC 24A includes 32 sets of analog signal processing circuits 33A.
It is converted to SI. The γ-ray detection signal output from the detector 21 and passed through the capacitor 22 and the resistor 23 is a charge amplifier 24b, a polarity amplifier 24c, and a bandpass filter 24.
The signal is input to the peak hold circuit 24e via d. The peak hold circuit 24e holds the peak value of the γ-ray detection signal. The γ-ray detection signal output from the bandpass filter 24d is input to the trigger output circuit 24f. The trigger output circuit 24f outputs a trigger signal when a γ-ray detection signal of a set level or higher is input in order to remove the influence of noise.

The digital ASIC 26A includes a packet data generation device 34A and a data transfer circuit 37.
These are LSIs. The above trigger signal is input to the ADC control circuit 36A of the packet data generation device 34A. All the digital ASICs 26A provided in the SPECT device 51 receive clock signals from a 64 MHz clock generator (crystal oscillator) (not shown) and operate in synchronization. Each digital ASIC26
The clock signal input to A is received by each ADC in all packet data generators 34A.
It is input to the control circuit 36A. The ADC control circuit 36A specifies the detector ID when the trigger signal is input. That is, the ADC control circuit 36A stores a detector ID for each trigger output circuit 24f connected to the ADC control circuit 36A, and a certain trigger output circuit 2
When the trigger signal is input from 4f, the detector ID corresponding to the trigger output circuit 24f can be specified. The ADC control circuit 36A outputs an ADC control signal including detector ID information to the ADC2
5 is output. The ADC 25 converts the peak value information output from the peak hold circuit 24e of the analog signal processing circuit 33A corresponding to the detector ID into a digital signal and outputs the digital signal. This peak value information is input to the ADC control circuit 36. The ADC control circuit 36A adds the peak value information to the detector ID and generates packet data. Packet data (including detector ID and peak value information) that is digital information output from the ADC control circuit 36A of each packet data generation device 34A is input to the data transfer circuit 37. The data transfer circuit 37 periodically transmits the packet data output from each ADC control circuit 36A to the unit integrated FPGA 31 of the detector unit 2A. The unit integrated FPGA 31 outputs the digital information to the information transmission wiring connected to the connector 38.

Each packet data output from the unit integrated FPGA 31 is stored in the data processing device 12A.
Sent to. The rotation angle detected by an angle meter (not shown) connected to the rotation shaft of a motor (not shown) that rotates the rotation support base 57 is input to the data processing device 12A. This rotation angle indicates the rotation angle of each radiation detection block 52, specifically, the rotation angle of each detector 21. Based on this rotation angle, the data processing device 12A determines the position (positional coordinate) of each detector 21 that is turning on the turning track. For this reason, the position (positional coordinate) of the detector 21 at the time of detecting the γ-ray is obtained. Based on the calculated position of the detector 21, the data processing device 12 </ b> A counts γ rays whose peak value information is equal to or greater than a set value. This counting is performed for each region obtained by dividing the rotation support table 57 by 0.5 ° with respect to the rotation center. The peak value information is the peak value of each γ-ray detection signal of a plurality of detectors 21 (four detectors 21 arranged in a straight line in FIG. 7A) located on the extension line of the radiation path of the collimator 55. Is the cumulative value of The data processing device 12A uses the position information of the detector 21 and the count value (count information) of the detector 21 at the time when the γ-ray is detected, to collect the radiopharmaceuticals, that is, the position of the subject at the malignant tumor position. Create tomographic image information. This tomographic image information is displayed on the display device 13. Information such as the packet information, the count value obtained by simultaneous measurement, the position information of the detector 21, and the tomographic image information is stored in the storage device of the data processing device 12.

In the above embodiments, the PET apparatus 1 and the SPECT apparatus 51 have been described.
The present invention can also be applied to a γ camera. The γ camera includes a collimator that has a two-dimensional functional image and regulates the incident angle of γ rays. In addition, it is good also as a structure of the nuclear medicine diagnostic apparatus which combined PET apparatus 1 or SPECT apparatus 51, and X-ray CT.
Further, the mounting (storage) of the detector unit 2 to the camera 11 is not limited to the above-described unit support member 2A, and any mounting / storage means / method may be used.

1 is a perspective view showing a configuration of a PET apparatus as a nuclear medicine diagnostic apparatus according to an embodiment. It is the figure which showed typically the cross section of the circumferential direction of the camera of the PET apparatus of FIG. It is a figure which shows typically the structure of the minimum structure of a semiconductor radiation detector. It is the graph which compared the "time-crest value curve" when the thickness t of the semiconductor material of a semiconductor radiation detector is thick and thin. 6 is a graph schematically showing the relationship between the thickness t of semiconductor material of a semiconductor radiation detector and the peak value (maximum value). It is a figure which shows typically the structure of the semiconductor radiation detector which has a laminated structure on which the semiconductor material and the electrode (anode, cathode) were laminated | stacked. (A) is the front view which showed the combined substrate which integrated the detector board | substrate and the ASIC board | substrate of the semiconductor radiation detector concerning this embodiment. (B) is a side view of the same (a). (C) is the perspective view which showed typically the structure of the semiconductor radiation detector mounted on the detector board | substrate of (a). 3 is a block diagram schematically showing an analog detection circuit. FIG. 2 is a block diagram showing a schematic configuration of a digital ASIC and a connection relationship between the analog ASIC and the digital ASIC. FIG. 4 is a perspective view cited for explaining the structure of a detector unit containing a plurality of semiconductor radiation detectors. FIG. 11 is a side view of the detector unit of FIG. 10 with a side plate removed. (A) is a partially broken perspective view of the camera showing a state when the detector unit is mounted on the camera, and (b) is a cross-sectional view of the central part of the camera. It is a perspective view which shows the structure of the SPECT apparatus as a nuclear medicine diagnostic apparatus concerning other embodiment. It is the block diagram which showed typically the circuit structure of analog ASIC in the SPECT apparatus of FIG. FIG. 14 is a block diagram showing a schematic configuration of a digital ASIC in the SPECT apparatus of FIG. 13 and a connection relationship between the analog ASIC and the digital ASIC.

Explanation of symbols

1 PET equipment (nuclear medicine diagnostic equipment)
11 Camera 12 Data processing device 13 Monitor 14 Bed 2 Detector unit 20 Bonding substrate (semiconductor radiation detection device)
20A detector substrate (first substrate)
20a Substrate body 20B ASIC substrate (second substrate)
20b Substrate body 21 Detector (semiconductor radiation detector)
24 Analog ASIC (Analog Integrated Circuit)
26 Digital ASIC (Digital Integrated Circuit)
30 Housing 30a Top plate 30b Bottom plate PS High voltage power supply device 51 SPECT device

Claims (52)

A semiconductor radiation detector used in a nuclear medicine diagnostic apparatus,
The semiconductor radiation detector has a semiconductor region that generates charges by interacting with radiation, and an anode electrode and a cathode electrode are opposed to each other across the semiconductor region. The inter-electrode distance or the thickness of the semiconductor region sandwiched between the anode electrode and the cathode electrode is 0.2-2 mm,
A semiconductor radiation detector. The distance between the electrodes or the thickness of the semiconductor region is 0.5 to 1.5 mm;
The semiconductor radiation detector according to claim 1. The area of the electrode or the area of the semiconductor surface forming the electrode is 4 to 120 square mm;
The semiconductor radiation detector of Claim 1 or Claim 2 characterized by these. A plurality of semiconductor radiation detectors having a plurality of semiconductor regions that generate electric charges by interacting with radiation, the anode electrodes and the cathode electrodes facing each other across the semiconductor regions; and the plurality of semiconductor radiations A plurality of signal processing devices that are respectively provided for the detectors and that input radiation detection signals from the corresponding semiconductor radiation detectors, and outputs from the plurality of signal processing devices are input, and during a set time A counting device that counts the outputs corresponding to the detected pair of radiation detection signals as a pair, and a tomogram information creation device that creates tomogram information using the count information output from the counting device. ,
In the semiconductor radiation detector, a distance between the anode electrode and the cathode electrode, or a thickness of the semiconductor region sandwiched between the anode electrode and the cathode electrode is 0.2 to 2 mm.
Being
A positron emission tomography apparatus characterized by: The distance between the electrodes or the thickness of the semiconductor region is 0.5 to 1.5 mm;
The positron emission tomography apparatus according to claim 4. The area of the electrode or the area of the semiconductor surface forming the electrode is 4 to 120 square mm;
The positron emission tomography apparatus according to claim 4, wherein Obtained based on a plurality of semiconductor radiation detectors that receive radiation, an integrated circuit that processes radiation detection signals output from each of the plurality of semiconductor radiation detectors, and first information output from the integrated circuit A tomogram information creation device for creating tomogram information using the second information,
Having a plurality of unit substrates including the plurality of semiconductor radiation detectors and the integrated circuit;
A nuclear medicine diagnostic device. A plurality of semiconductor radiation detectors that receive radiation, an integrated circuit that processes a radiation detection signal output from each of the plurality of semiconductor radiation detectors, and an output from each of the integrated circuits are input during a set time. A counting device that counts the outputs corresponding to the detected pair of radiation detection signals as a pair, and a tomogram information creation device that creates tomogram information using the count information output from the counting device. ,
Having a plurality of unit substrates including the plurality of semiconductor radiation detectors and the integrated circuit;
A nuclear medicine diagnostic device. The integrated circuit processes an analog integrated circuit that processes a signal output from the semiconductor radiation detector, an AD converter that converts an analog signal output from the analog integrated circuit into a digital signal, and a digital that processes an AD-converted signal Including integrated circuits,
The nuclear medicine diagnosis apparatus according to claim 7 or 8, characterized in that. The semiconductor radiation detector, the analog integrated circuit, the AD converter, and the digital integrated circuit are arranged in that order from one end of the unit substrate to the other end in the longitudinal direction of the unit substrate. about,
The nuclear medicine diagnosis apparatus according to claim 9. The analog integrated circuit has a function of performing signal amplification, and the digital integrated circuit has a function of determining a radiation detection time;
The nuclear medicine diagnosis apparatus according to claim 9. The unit substrate includes a first substrate and a second substrate,
The first substrate comprises the semiconductor radiation detector;
The second substrate comprises the integrated circuit;
The nuclear medicine diagnosis apparatus according to claim 7, wherein: The unit substrate includes a first substrate and a second substrate,
The first substrate comprises the semiconductor radiation detector and the analog integrated circuit;
The second substrate includes the digital integrated circuit;
The nuclear medicine diagnosis apparatus according to any one of claims 9 to 11, characterized in that: The unit substrate includes a first substrate, a second substrate, and a third substrate,
The first substrate comprises the semiconductor radiation detector;
The second substrate includes the analog integrated circuit;
The third substrate comprises the digital integrated circuit;
The nuclear medicine diagnosis apparatus according to any one of claims 9 to 11, characterized in that: The first substrate and the second substrate are detachably coupled to each other;
The nuclear medicine diagnosis apparatus according to claim 9, characterized in that: The second substrate and the third substrate are detachably coupled to each other;
The nuclear medicine diagnosis apparatus according to claim 14. A plurality of first substrates are detachably coupled to the second substrate;
The nuclear medicine diagnosis apparatus according to claim 12, wherein: The first substrate and the second substrate are joined with their respective ends overlapped;
The nuclear medicine diagnosis apparatus according to any one of claims 12 to 17, characterized by: The semiconductor radiation detector is disposed on both sides of the unit substrate;
The nuclear medicine diagnosis apparatus according to any one of claims 7 to 18, characterized by: A rotating body and a bed on which the subject is placed;
The plurality of unit substrates are attached to a support member provided on the rotating body and configured to turn around the bed,
The semiconductor radiation detector of the unit substrate is disposed on the bed side,
A collimator having a plurality of radiation paths facing the semiconductor radiation detector and disposed on the bed side of the semiconductor radiation detector is installed on the support member;
The nuclear medicine diagnosis apparatus according to claim 7. A bed for supporting a subject, wherein the plurality of unit substrates are arranged so as to surround the bed, and the semiconductor radiation detector of the unit substrate is arranged on the bed side,
The nuclear medicine diagnosis apparatus according to claim 7 or 8, characterized in that. A plurality of semiconductor radiation detectors that receive radiation; and an integrated circuit that processes a radiation detection signal output from each of the plurality of semiconductor radiation detectors,
The integrated circuit outputs radiation detection time information and identification information of the semiconductor radiation detector that detected the radiation,
Having a plurality of unit substrates including the plurality of semiconductor radiation detectors and the integrated circuit;
A positron emission tomography apparatus characterized by: An analog integrated circuit for processing the signal output from the semiconductor radiation detector by the integrated circuit, an AD converter for converting an analog signal output from the analog integrated circuit into a digital signal, and a digital integrated processing for an AD converted signal Including the circuit,
The positron emission tomography apparatus according to claim 22. The semiconductor radiation detector, the analog integrated circuit, the AD converter, and the digital integrated circuit are arranged in that order from one end of the unit substrate to the other end in the longitudinal direction of the unit substrate. about,
24. The positron emission tomography apparatus according to claim 23. The analog integrated circuit has a function of performing signal amplification, and the digital integrated circuit has a function of generating the time information;
25. The positron emission tomography apparatus according to claim 23 or 24. The unit substrate includes a first substrate and a second substrate,
The first substrate has the semiconductor radiation detector;
The second substrate comprises the integrated circuit;
26. The positron emission tomography apparatus according to any one of claims 22 to 25, wherein: The semiconductor radiation detector is disposed on both sides of the unit substrate;
27. The positron emission tomography apparatus according to any one of claims 22 to 26, wherein: The analog integrated circuit is provided for each of the semiconductor radiation detectors, and includes a plurality of signal processing devices that process the radiation detection signals including amplifiers that input the radiation detection signals output from the semiconductor radiation detectors. The digital integrated circuit outputs the time information and the identification information based on the output of the signal processing device,
25. The positron emission tomography apparatus according to claim 23 or 24. The analog integrated circuit inputs the radiation detection signal from the upstream side with respect to the slow system including the amplifier and a peak value output device that inputs the output of the amplifier and outputs the peak value of the radiation detection signal. A first system including a timing detection device that outputs a radiation detection timing signal, and the digital integrated circuit is provided for each semiconductor radiation detector, and generates time information based on the radiation detection signal Including equipment,
25. The positron emission tomography apparatus according to claim 23 or 24. When the digital integrated circuit further inputs time information from the time information generation device, one of the position information and the identification information of the semiconductor radiation detector connected to the time information generation device is obtained. An AD conversion control device to identify, an information integration device that integrates the one information, the time information and the peak value information,
The AD converter includes a wave from the peak value output device determined by the one information specified by the AD conversion control device among the peak value output devices of a plurality of signal processing devices included in the analog integrated circuit. Converting high value into peak value information, which is digital information, and outputting the information to the information integration device;
30. The positron emission tomography apparatus according to claim 29. A semiconductor radiation detection device used in a nuclear medicine diagnostic device,
A plurality of semiconductor radiation detectors and an integrated circuit for processing signals output from these semiconductor radiation detectors are installed on a unit substrate, and a plurality of signal lines for transmitting output signals of the respective semiconductor radiation detectors to the integrated circuit are provided. Provided on the unit substrate;
A semiconductor radiation detection apparatus. An analog integrated circuit for processing the signal output from the semiconductor radiation detector by the integrated circuit, an AD converter for converting an analog signal output from the analog integrated circuit into a digital signal, and a digital integrated processing for an AD converted signal Including the circuit,
32. The semiconductor radiation detection apparatus according to claim 31, wherein: The semiconductor radiation detector, the analog integrated circuit, the AD converter, and the digital integrated circuit are arranged in that order from one end of the unit substrate to the other end in the longitudinal direction of the unit substrate. about,
The semiconductor radiation detection apparatus according to claim 32, wherein: The unit substrate includes a first substrate and a second substrate,
The first substrate has the semiconductor radiation detector;
The second substrate comprises the integrated circuit;
The semiconductor radiation detection apparatus according to any one of claims 31 to 33, wherein: The first substrate and the second substrate are detachably coupled to each other with their ends overlapped;
The semiconductor radiation detection apparatus according to claim 34, wherein: The semiconductor radiation detector is disposed on both sides of the unit substrate;
36. The semiconductor radiation detection apparatus according to any one of claims 31 to 35, wherein: Obtained based on a plurality of semiconductor radiation detectors that receive radiation, an integrated circuit that processes radiation detection signals output from each of the plurality of semiconductor radiation detectors, and first information output from the integrated circuit A tomogram information creation device for creating tomogram information using the second information,
And a plurality of frames attached to the support member in a detachable manner;
A plurality of semiconductor radiation detectors, and a plurality of unit substrates including the integrated circuit,
A plurality of unit boards are detachably stored in the frame;
A nuclear medicine diagnostic device. A plurality of semiconductor radiation detectors that receive radiation, an integrated circuit that processes a radiation detection signal output from each of the plurality of semiconductor radiation detectors, and an output from each of the integrated circuits are input during a set time. A counting device that counts the outputs corresponding to the detected pair of radiation detection signals as a pair, and a tomogram information creation device that creates tomogram information using the count information output from the counting device. ,
And a plurality of frames attached to the support member in a detachable manner;
A plurality of semiconductor radiation detectors, and a plurality of unit substrates including the integrated circuit,
A plurality of unit boards are detachably stored in the frame;
A nuclear medicine diagnostic device. An analog integrated circuit for processing the signal output from the semiconductor radiation detector by the integrated circuit, an AD converter for converting an analog signal output from the analog integrated circuit into a digital signal, and a digital integrated processing for an AD converted signal Including the circuit,
A nuclear medicine diagnosis apparatus according to claim 37 or claim 38. The semiconductor radiation detector, the analog integrated circuit, the AD converter, and the digital integrated circuit are arranged in that order from one end of the unit substrate to the other end in the longitudinal direction of the unit substrate. about,
40. A nuclear medicine diagnostic apparatus according to claim 39. The unit substrate includes a first substrate and a second substrate,
The first substrate has the semiconductor radiation detector;
The second substrate comprises the integrated circuit;
40. The nuclear medicine diagnostic apparatus according to any one of claims 37 to 39, wherein: The frame body is provided with a plurality of guide members for guiding the unit substrate therein,
A nuclear medicine diagnosis apparatus according to claim 37 or claim 38. A power supply device including a voltage booster for boosting a voltage in the frame;
A voltage supply wiring for supplying a voltage to each of the semiconductor radiation detectors of the unit substrate from the power supply device;
40. The nuclear medicine diagnosis apparatus according to any one of claims 37 to 39. The frame has a light-shielding configuration;
45. The nuclear medicine diagnostic apparatus according to any one of claims 37 to 39, wherein: The frame is provided with a lid, and the lid includes an integrated circuit for integration that integrates and processes information output from the digital integrated circuits of the plurality of unit substrates.
The nuclear medicine diagnosis apparatus according to any one of claims 14 to 17, wherein: A plurality of semiconductor radiation detectors that receive radiation; and an integrated circuit that processes a radiation detection signal output from each of the plurality of semiconductor radiation detectors,
And a plurality of frames attached to the support member in a detachable manner;
A plurality of semiconductor radiation detectors, and a plurality of unit substrates including the integrated circuit,
A plurality of unit substrates are detachably stored in the frame body,
The integrated circuit outputs radiation detection time information and identification information of the semiconductor radiation detector that has detected the radiation;
A positron emission tomography apparatus characterized by: An analog integrated circuit for processing the signal output from the semiconductor radiation detector by the integrated circuit, an AD converter for converting an analog signal output from the analog integrated circuit into a digital signal, and a digital integrated processing for an AD converted signal Including the circuit,
47. The positron emission tomography apparatus according to claim 46, wherein: A semiconductor radiation detection apparatus according to any one of claims 31 to 36;
A frame having a guide member for detachably storing the plurality of semiconductor radiation detection devices,
A detector unit characterized by. The frame guide means stores a plurality of semiconductor radiation detection devices so as not to overlap in the depth direction, and stores a plurality of semiconductor detection devices in the frontage direction,
49. A detector unit according to claim 48. A power supply device including a voltage booster for boosting a voltage in the housing;
Voltage supply wiring for supplying a voltage from the power supply device to each of the semiconductor radiation detectors of the semiconductor radiation detection device;
50. A detector unit according to claim 48 or claim 49. The frame has a light-shielding configuration;
51. The detector unit according to any one of claims 48 to 50, wherein: The frame is provided with a lid, and the lid includes an integrated circuit for integration that integrates and processes information output from the digital integrated circuits of the plurality of unit substrates.
52. The detector unit according to any one of claims 48 to 51, wherein:

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