JP2007078671A - Radiation ray detection module, printed circuit board and diagnostic device of nuclear medicine - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation ray detection module, printed circuit board and diagnostic device of nuclear medicine capable of improving spatial resolution without increasing the number of channels and capable of diagnosing highly precisely. <P>SOLUTION: The module includes radiation ray detectors 21 and at least a plurality of wiring boards 24 arranged in the progressing direction of the radiation ray. The wiring board 24 keeps the adjacent radiation ray detectors 21 connected to the adjacent detectors in the progressing direction of the radiation ray mutually electrically into one detector body (detector channel) 21A. The electrodes to be connected mutually of the radiation ray detectors 21 are fixed in the state mutually opposite to each other on the wiring board. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

The present invention relates to a radiation detection module, a printed board, and a nuclear medicine diagnostic apparatus, and more particularly to a radiation detection module, a printed board, and a nuclear medicine diagnostic apparatus using a semiconductor detection element.

Conventionally, as radiation detectors for detecting radiation such as γ rays, CdTe (cadmium telluride), CdZnTe (cadmium, zinc, tellurium), TlBr (thallium bromide), GaAs
There is a semiconductor radiation detector provided with a semiconductor detection element made of a semiconductor material such as (gallium arsenide). Semiconductor detectors convert electric charges generated by the interaction between radiation and semiconductor materials into electrical signals, so semiconductor radiation detectors that use semiconductor detectors are more efficient in converting electrical signals than those that use scintillators. It is good and can be miniaturized.

A semiconductor radiation detector includes the semiconductor detection element and electrodes formed on both sides of the semiconductor detection element. By applying a DC high voltage between these electrodes, X-rays,
Electric charges generated when radiation such as gamma rays enters the semiconductor detection element are taken out as signals from the electrodes.
When a semiconductor radiation detector is used in a medical radiation imaging apparatus or the like (nuclear medicine diagnostic apparatus), a radiation detector is formed by connecting the semiconductor radiation detector on a wiring board (for example, see Patent Document 1). In addition, a technique in which a plurality of semiconductor radiation detectors are arranged on a radiation detector support plate has been proposed (see, for example, Patent Document 2).

JP 2003-84068 A (paragraph 0024, FIG. 3) Japanese Patent Laying-Open No. 2003-167058 (paragraphs 0020 to 0021, FIG. 3)

By the way, in order to obtain a highly accurate image in a PET (Positron Emission Tomography) apparatus which is a kind of nuclear medicine diagnostic apparatus, there is a demand to increase the spatial resolution. In addition, there is a demand for improving the γ-ray detection sensitivity, for example, increasing the arrangement density of radiation detectors, in order to shorten the inspection time in the PET apparatus. These requests are SPECT
(Single Photon Emission Computer Tomography) devices and other nuclear medicine diagnostic devices of γ cameras.

In order to increase the spatial resolution, it is conceivable to increase the number of signals output from the radiation detector (increase the number of channels). However, increasing the number of channels in this way increases the number of processing circuits, which increases the size of the substrate for mounting the processing circuit, increases the power consumption, and further increases the amount of heat generated by the entire substrate. Since it increases, there exists a possibility that the detection sensitivity of a radiation detector may fall.

An object of the present invention is to provide a radiation detection module, a printed circuit board, and a nuclear medicine diagnostic apparatus that can improve spatial resolution and increase diagnostic accuracy without increasing the number of channels.

In order to achieve the above-described object, in the present invention, a radiation detection module includes a radiation detector, and a wiring board to which a plurality of the radiation detectors are arranged and attached in a radiation traveling direction. Since the radiation detectors adjacent to each other in the radiation traveling direction are electrically connected to each other to form one detector body (detection channel), a pair of adjacent radiation detectors are elongated in the radiation traveling direction. It can be used as a single detector body (detection channel) formed, and the radiation is easily captured accordingly. In addition, since the detection channel configured by electrically connecting the radiation detectors adjacent to each other in the radiation traveling direction in this way constitutes one detection unit (channel), the radiation channel is substantially in the radiation traveling direction. The number of channels will be reduced. Since the number of detection channels increases in the X direction, the number of channels in the X direction can be increased. Thereby, in the nuclear medicine diagnosis apparatus, the number of channels in the X direction can be increased without increasing the total number of channels. Thereby, the spatial resolution of the obtained image is improved. In addition, since the total number of channels does not increase, an increase in heat generation of the signal processing circuit is also suppressed.

The radiation detector is arranged in the radiation traveling direction on the wiring board, and further, a semiconductor detection element having a cathode electrode on one surface and an anode electrode on the other surface is orthogonal to the radiation traveling direction. Since a plurality are arranged in parallel in the direction, it can be configured to detect radiation for each radiation detector in a direction orthogonal to the radiation traveling direction. Therefore, the detection area can be subdivided in the direction orthogonal to the radiation traveling direction,
Spatial resolution can be increased. In addition, the radiation detector is electrically connected to either one of the cathode electrode and the anode electrode between the radiation detector and another radiation detector adjacent to the radiation traveling direction on the wiring board. Since one detector body (detection channel) is paired with another radiation detector, each pair of adjacent radiation detectors is formed as one detector body (detection channel) formed long in the radiation traveling direction. Can be used and easy to capture radiation.
In addition, by using a detection module in which such a radiation detector is provided on a wiring board, a nuclear medicine diagnostic apparatus having excellent spatial resolution of an image obtained for a subject and excellent radiation capturing performance is provided. can get.

According to the present invention, it is possible to obtain a radiation detection module, a printed circuit board, and a nuclear medicine diagnostic apparatus that can improve spatial resolution and increase diagnosis without increasing the number of channels.

Next, a PET apparatus which is a preferred embodiment as a nuclear medicine diagnosis apparatus to which a radiation detection module having a detection channel is applied will be described in detail with reference to the drawings as appropriate.
As shown in FIG. 1A, the PET apparatus according to this embodiment includes a PET imaging apparatus 1, a bed 31 that supports a subject (examinee) H, a data processing apparatus (computer or the like) 2, and a display apparatus 3. I have. In the PET imaging apparatus 1, the subject H is placed on a bed 31 movable in the longitudinal direction and inserted into a measurement space 32 surrounded by the unit substrates U.

(PET device)
The PET imaging apparatus 1 has a unit substrate (printed substrate) U that surrounds a measurement space (measurement region) 32 in which a bed 31 is inserted and is arranged in the circumferential direction. As shown in FIG. 1B, the unit substrate U has an axial direction of the measurement space 32 (longitudinal direction of the bed 31 (FIG. 1A).
Also see))). As shown in FIGS. 2A and 2B, the unit substrate U includes a radiation detection module (hereinafter referred to as a detection module) 20A and an integrated circuit substrate (hereinafter referred to as an ASIC substrate) 20B. As shown in FIG. 2A, the detection module 20A includes a plurality of semiconductor radiation detectors (hereinafter simply referred to as detectors) 2 on a wiring board 24.
1 is provided. The detector 21 detects γ rays emitted from the body of the subject H. In the present embodiment, a direction orthogonal to the radiation traveling direction (FIG. 2) is formed on the wiring board 24 of the detection module 20A.
In (a), the direction of the arrow X (hereinafter simply referred to as the X direction) and the radiation traveling direction (FIG. 2 (a)).
In this case, a plurality of detectors 21 are arranged in an orderly manner in the arrow Y direction (hereinafter simply referred to as the Y direction). In this embodiment, the interval between the detectors 21 arranged in the X direction is set to be smaller than the interval between the detectors 21 arranged in the Y direction, and the detectors 21 are densely arranged in the X direction. ing. Here, the X direction corresponds to the circumferential direction in the PET imaging apparatus 1,
Further, the Y direction corresponds to the radial direction (radiation traveling direction) in the PET imaging apparatus 1 (FIG. 1).
(See (a)). On the wiring substrate 24, one detection channel 21A is configured by electrically connecting the anode electrodes A of a pair of detectors 21 adjacent in the Y direction to each other. The detection channel 21A is elongated in the radiation traveling direction. It can be said that the Y direction is a direction from the arrangement region of the detector 21 toward the arrangement region of the signal processing circuit (analog ASIC 28), and the X direction is a direction orthogonal to the direction.

The ASIC substrate 20B has an application specific integrated circuit (ASIC 28, 29) for measuring the detected peak value and detection time of the γ-ray, and the detected peak value and detection time of the radiation (γ-ray). Is supposed to measure. The integrated circuit includes a plurality of signal processing devices that process radiation detection signals.

Next, details of the PET imaging apparatus 1 will be described.
(Detector assembly)
The detector assembly 41 applied to this embodiment will be described. The detector assembly 41 has two detectors 21. Each detector 21 includes two semiconductor detection elements (hereinafter referred to as detection elements, FIG. 3).
(See (a)) 211 and conductive members 22 and 23 (see FIG. 3B). Detection element 2
As shown in FIG. 3A, 11 includes a semiconductor element S made of a plate-like semiconductor material and thin film-like electrodes formed by vapor deposition or the like over the entire surfaces of both sides thereof.
An electrode formed on one surface of the detection element 211 is an anode electrode (hereinafter referred to as an anode) A
The electrode formed on the other surface is a cathode electrode (hereinafter referred to as cathode) C.
In each detector 21, detection elements 211 are arranged in parallel in the X direction so that the anodes A face each other, and a conductive member 23 is arranged between the anodes A and is attached to the anodes A with a conductive adhesive. Configured. Those anodes A are electrically connected by a conductive member 23. In addition, a conductive member 22 is attached to each cathode C of each detection element 211 of the detector 21 with a conductive adhesive. In the detector assembly 41,
The conductive member 22 positioned between the adjacent cathodes C of the two detectors 21 is shared by the detectors 21 and connected to the cathodes C. In the detector assembly 41,
The anodes A and cathodes C provided on the four detection elements 211 are alternately arranged, and the conductive members 22 and the conductive members 23 are also alternately arranged.

The semiconductor element S is a region that generates charges by interacting with radiation, and CdTe,
It is made of any single crystal such as CdZnTe or GaAs. The cathode C and the anode A are made of any material such as Pt, Au, and In. In the present embodiment, the detection element 211 uses, for example, CdTe for the semiconductor element S, Pt for the cathode C, and In for the anode A, and forms a pn junction diode.
Here, the relationship between the time at the thickness t of the semiconductor element S (see FIG. 3A) and the peak value curve will be described. When the reverse bias voltage (hereinafter referred to as the bias voltage) of the pn junction applied between the cathode C and the anode A is the same value, the semiconductor element S having a smaller thickness t has a faster increase (rise) of the crest value. The accuracy (energy resolution) of the peak value increases. When the crest value rises quickly, for example, the accuracy of simultaneous measurement in the PET imaging apparatus 1 (simultaneous counting resolution)
Will improve. The semiconductor element S having a small thickness t has a higher crest value rise rate and higher energy resolution (improves the charge collection efficiency) because the time for the electrons to reach the anode A and the holes for the cathode C This is because the time for arriving at is reduced, that is, the charge collection time is shortened. In addition, the holes that may disappear in the middle can reach the cathode C without disappearing because the thickness t is thin. Incidentally, the thickness t can also be expressed as an interelectrode distance between the cathode C and the anode A. The anode A is an electrode for extracting a radiation detection signal, and the cathode C is an electrode for applying a bias voltage. The thickness (interelectrode distance) t of the semiconductor element S is preferably 0.2 mm to 2 mm.

The conductive members 22 and 23 are, for example, copper (or a copper-based material such as phosphor bronze), and have the same size as each electrode of the detection element 211. Conductive members 22, 23
Is a non-sensitive region that does not detect radiation, and is preferably formed to be thin and have a thickness at which a detection signal is output satisfactorily, for example, about 10 μm to 100 μm, mainly 50 μm.
The degree is good. The conductive members 22 and 23 are connected portions 2 for constituting the detector channel 21A.
2A and 23A (see FIG. 3B). In this example, as will be described later, the connecting portions 23A of the detectors 21 adjacent in the Y direction are electrically connected to each other. Connection unit 22
A and 23A are projecting portions 22a and 23a projecting outward from the semiconductor element S (sideward, in the direction of arrow Y in FIG. 2A), and terminal portions 22b suspended from the projecting portions 22a and 23a. , 23b
And have. Specifically, the overhanging portion 22a of the conductive member 22 connected to the cathode C is located on one side surface (left side in FIG. 3C) of each detector 21 in the detector assembly 41.
The overhanging portion 23a of the conductive member 23 connected to the anode A is located on the other side surface (right side in FIG. 3C) of each detector 21, which is opposite to the side surface. That is, the detector 21 has three overhanging portions 22a (only one is shown in FIG. 3C) on one side surface and two overhanging portions 23a on the other side surface. The terminal portions 22b and 23b have bent portions 22c and 23c for electrically connecting to a connection member AP or a connection member CP on the wiring board 24 described later.

Thus, the conductive members 22 and 23 also serve as a fixing member for fixing the detector 21 to the wiring board 24. In particular, the overhang portions 22 a and 23 a serve as a fixing portion for attaching the detector 21 to the wiring board 24. Note that the material of the conductive members 22 and 23 is not limited to copper, and may be aluminum or an aluminum alloy, and the shape does not need to be a plate shape. Furthermore, the conductive members 22 and 23 are preferably the same size as the electrodes of the detection element 211, but need not be exactly the same size.
Further, as shown in FIG. 3B, the terminal portions 22b and 23b are formed to be narrow so as to have a protruding amount smaller than the protruding amount of the protruding portions 22a and 23a. Thereby, it is possible to minimize the scattering of γ rays at the portions of the terminal portions 22b and 23b. Further, since the terminal portions 22b and 23b are formed narrow, heat transmitted from the wiring board 24 to the detector 21 is reduced, and the stability of the detector 21 is improved.
In the present embodiment, the detector 21 is connected to the wiring board 24 by such terminal portions 22b and 23b.
It is attached to the state where it floats slightly. That is, in the detector 21, each bottom surface portion of the detection element 211 is attached to the wiring board 24 in a non-contact state. Thereby, it is possible to prevent the detector 21 from being rubbed and damaged on the wiring board 24 when the detector 21 is mounted. Also,
It is possible to positively prevent the detector 21 from being attached to the wiring board 24 in a state in which the insulating property is lowered due to a different member (dust or the like) being sandwiched between the detector 21 and the wiring board 24. . Note that an insulating material (not shown) may be coated on the bottom surface of the detector 21 to further prevent unexpected dielectric breakdown.

The detector 21 having such conductive members 22 and 23 is a pair of detectors 2 adjacent in the Y direction.
One anode A (see FIG. 3 (c), the same applies hereinafter) are connected to each other at the connecting portion 23 of the conductive members 23, 23.
A detection channel 21A is configured by being electrically connected to each other via A and 23A.
In such a configuration, the detectors 21 adjacent to each other in the Y direction are mounted on the wiring board 24 so that the connection portions 23A of the conductive members 23 connected to the anode A are opposed to each other. Specifically, as shown in FIGS. 4 and 5, the detector 21 connects the terminal of the connection portion 23 </ b> A to the connection member AP for anode A (for example, a pad of the wiring substrate) provided on the surface of the wiring substrate 24. The portion 23b (not shown in FIG. 4A, the same applies hereinafter) is electrically connected to the connecting member CP for the cathode C and the terminal portion 22b of the connecting portion 22A (not shown in FIG. 4A, the same applies hereinafter). The wiring board 24 is installed in a state where it is electrically connected by the adhesive 25. That is, the pair of detectors 21 adjacent to each other in the Y direction are electrically connected to each other at the connection portions 23 </ b> A that face each other via the connection member AP on the wiring board 24.

As described above, the anodes A of the detectors 21 adjacent to each other in the Y direction are connected to each other via the connection member AP, so that two detection channels 21A are configured in the X direction. These detection channels 21A output detection signals from the respective connection members AP. On one side of the wiring board 24, 32 detection channels 21A are installed in the X direction. Here, when this embodiment is compared with a configuration example in which such a detector channel 21A is not configured and two detector aggregates 41 are simply arranged in the Y direction, the number of detection signals to be output is 2 However, the present embodiment is largely different in that γ rays can be detected in the two detection channels 21A formed by the two detector assemblies 41. As a result, since the number of channels in the X direction is doubled, the resolution in the X direction is doubled. On the other hand, the number of detection signals output as described above is the detection channel 2
Since 1A is not configured, it is not necessary to increase the number of signal processing circuits by configuring the detection channel 21A. In other words, there is no situation where the number of channels increases as the radiation detection area is subdivided, and the existing channel number (the number of channels when the detection channel 21A is not configured) is used as it is to improve the spatial resolution. To be able to.
In addition, since the detection channel 21A is formed elongated in the Y direction, the detection module 20A having excellent radiation capturing capability can be constructed. The detector assembly 41 is further divided in the X direction (for example, the detection elements 211 are in units of one), and the detectors 21 adjacent to each other in the Y direction are electrically connected to each other to form a detection channel. It is also possible to configure and output signals from these. However, such a configuration is preferable because a total of four signals are output from a pair of detector assemblies 41 adjacent in the Y direction, resulting in an increase in the number of channels of detection signals. Absent. In this regard, in the present embodiment, as described above, each detection channel 2 including the pair of detectors 21 adjacent in the Y direction.
Detection signals are output from 1A. For this reason, it is possible to improve spatial resolution and capture of radiation without substantially increasing the number of channels, and to perform highly accurate diagnosis that enables detection of small cancers. In addition, it is possible to avoid an increase in the size of the wiring board 24 due to an increase in the number of signal processing circuits (signal processing devices), and there is no increase in power consumption and high heat generation in the signal processing circuit.

As shown in FIG. 4, the plurality of connection members AP are installed between a pair of detectors 21 adjacent in the Y direction. A signal line (wiring) 24b embedded in the wiring board 24 is connected to each connection member AP. Thereby, the γ-ray detection signal output from each detection channel 21A is output through each signal line 24b. Each signal line 24 b is separately connected to a plurality of terminals 34 provided at the end of the wiring board 24.

Further, the four connecting members CP are in the Y direction of the wiring board 24 (the left-right direction in FIG. 4A).
In addition, they are arranged in parallel with the connection member AP sandwiched therebetween. All connection members CP are given the same potential in a state of being short-circuited by the wiring 24 a embedded in the wiring substrate 24. The wiring 24 a is connected to a terminal 33 provided at the end of the wiring substrate 24.

In the present embodiment, as described above, the six detectors 21 arranged in a row in the Y direction are connected to the conductive member 23 (connected to the anode A) in the two detectors 21 adjacent in the Y direction. 2
It arrange | positions so that 3A may face each other, and it is attached to applicable connection member AP with the adhesive agent 25 (refer FIG. 5). Further, the connection portions 22A of the respective detectors 21 are attached to the connection member CP by the adhesive 25. Here, of the four connection members CP, the two detectors 21 arranged on both sides sandwiching the two connection members CP1 located at the center,
The connection portions 22A of the respective conductive members 22 (connected to the cathode C) face each other. That is, these connection members CP1 are shared by the detection channels 21A arranged on both sides in the Y direction across these.

As the adhesive 25, a conductive paste, solder, or the like can be used. In consideration of the convenience of removing the detector 21 in an abnormal state such as a failure, in this embodiment, the detector assembly 41 from the wiring board 24, the adhesive 25 is preferably a thermoplastic adhesive.

In the detector, each semiconductor element S arranged in parallel has the thickness t (0.2 to 2 mm,
3), and the thickness of the cathode C and the anode A is at most several μm.
Degree.
The configuration of the detection channel 21A using the two detectors 21 as described above is based on the semiconductor element S.
The thickness t (see FIG. 3A) is reduced to increase the charge collection efficiency, increase the crest value rising speed to improve the energy resolution, and pass through the parallel arrangement of the semiconductor elements S. This is because the amount of γ rays is reduced to increase the interaction between the semiconductor element S and γ rays (to increase the count of γ rays). Increasing the count number of γ rays improves the sensitivity of the detection channel 21A.

Here, an outline of the principle of detection of γ rays by the detection channel 21A will be described. When γ rays enter the detection channel 21A from the Y direction and the γ rays interact with the semiconductor element S, holes and electrons are paired into the energy of the γ rays. Produced in proportion. Incidentally, the cathode C of the detection element 211 constituting the detection channel 21A.
And a bias voltage for collecting charges from a DC high-voltage power supply (not shown) (for example, cathode C is −500 V and anode A is close to the ground potential,
A reverse applied voltage is applied to the cathode C so that the anode A is increased by 500V. For this reason, holes corresponding to positive charges are attracted and moved to the cathode C, and electrons that are negative charges are attracted and moved to the anode A. When these holes and electrons are compared, the ease of movement (mobility) is relatively large, so that the electrons reach the anode A in a relatively short time. On the other hand, since the holes are relatively less likely to move, the holes reach the cathode C over time. Incidentally, holes may be trapped in the middle before reaching the electrodes.
In the detection channel 21A, the conductive member 23 disposed between the anodes A and the conductive member 22 attached to the cathode C are insensitive areas where γ rays are not detected. Note that the anode A and the cathode C are also insensitive areas.

Due to the above-described arrangement of the detection channel 21A, the number of channels arranged on the wiring board 24 of the detection module 20A is as shown in FIGS. 2 (a) and 2 (b).
3ch in the Y direction (radial direction of the PET imaging device 1) heading from the first to the ASIC substrate 20B, 32ch in the X direction (circumferential direction of the PET imaging device 1) orthogonal to the Y direction, and Z which is the thickness direction of the wiring board 24 It becomes 2ch (both surfaces of the wiring board 24) in the direction (depth direction of the PET imaging device 1). A total of 96 detection channels 21A are installed on one side of the wiring board 24, and a total of 192 detection surfaces on both sides.
In the detection module 20A of the present embodiment, the arrangement density of the detectors 21 is increased by three configurations as described below, thereby achieving high accuracy of the detection channel 21A.
The first configuration is that the detectors 21 are arranged so that the connecting portions 22A and 23A face the Y direction. Thereby, since those connecting portions 22A and 23A are not directed in the X direction, the gap between the detectors 21 in the X direction can be narrowed. Thereby, coupled with the fact that the detector 21 adjacent in the Y direction forms the detection channel 21A, the subject H (FIG. 1)
The resolution in the direction around the body axis (see (a)) (see (a)) can be improved.
In the second configuration, each detector 21 has an electrode of the same polarity (for example, cathode C) in the X direction.
) The detectors 21 are configured by being arranged so as to face each other. With this arrangement, the insulation between the detectors 21 in the X direction can be reduced, so that the gap between the detectors 21 in the X direction can be reduced. In particular, by forming a detector assembly 41 in which two detectors 21 arranged in the X direction share one cathode C, it is possible to increase the density of the detectors 21 in the X direction.
Further, in the third configuration, the detectors 21 adjacent in the Y direction have the same polarity electrode (anode A).
Alternatively, the connection portions 22A and 23A (electrical connection portions with the outside of the detector 21) connected to the cathode C) are arranged to face each other. With this arrangement, insulation between the detectors 21 in the Y direction can be reduced, so that the gap between the detectors 21 in the Y direction can be narrowed, and the detectors 21 can be densified in the Y direction.

In the first and second configurations, the gap between the detectors 21 in the X direction is narrowed, so that γ-rays emitted from the body of the subject H on the bed 31 are converted into FIG. , When traveling in the Y direction (ie, in the radial direction of the PET imaging apparatus 1 (see FIGS. 1A and 1B)), the gamma rays that pass through without being detected by the detector 21 (adjacent detection) The ratio of γ rays passing through the gap formed between the vessels 21 can be reduced, so that the detection efficiency of γ rays is increased by each of the first and second configurations, and the spatial resolution of the obtained image There is also an advantage that the inspection time can be shortened.
In the third configuration, since the gap between the detectors 21 in the Y direction is narrowed, the γ rays traveling obliquely with respect to the wiring board 24 in the Y direction are mutually connected in the Y direction. The rate of passing through the gaps between them is reduced. Therefore, the third configuration can also increase the detection efficiency of γ rays and improve the spatial resolution of the obtained image, similarly to the first and second configurations. It also contributes to shortening the examination time.
Thus, in the detection module 20A having the first, second, and third configurations, the detection efficiency of γ rays can be increased, and the spatial resolution of the image can be improved.

In this embodiment, since the detectors 21 are installed on both sides of the wiring board 24, they are arranged in the depth direction (Z direction) of the PET imaging apparatus 1 as compared with the case where the detectors 21 are installed only on one side. The number of wiring boards 24 can be halved. For this reason, the detectors 21 can be arranged more densely in the Z direction. Also in this case, the detection efficiency of γ rays increases, which contributes to the improvement of the spatial resolution of the image. In addition, the examination time can be further shortened. In addition, as described above, since the number of wiring boards 24 (unit boards U) can be reduced by half, there is an advantage that the labor of mounting the unit boards U on the PET imaging apparatus 1 (see FIG. 1) can be saved. is there.

By the way, the detector 21 is preferably coated with an insulating material to avoid dielectric breakdown. The film made of an insulating material is immersed in an insulating material such as silicone rubber together with the detection module 20A.
Thereafter, the film can be formed to a thickness of several tens of microns by drying. in this case,
After the insulating film of the detector 21 is removed except for the connecting portions 22A and 23A, the connecting portions 22
The terminal portions 22b and 23b of A and 23A may be attached to the corresponding connecting members AP and CP of the wiring board 24. The conductive members 22, 23 are electrode connecting members AP,
In order to reduce the risk of dielectric breakdown with CP, it may be formed smaller than the detection element 211, and a part of the overhang portions 22 a and 23 a may be positioned between the detection elements 211.

Further, as shown in FIGS. 6A and 6B, the connecting portion 22A between the detectors 21 adjacent in the Y direction.
, 23A may be connected to the connection member AP on the wiring board 24 in a state where they are overlapped. With this configuration, the detectors 21 can be arranged closer to each other in the Y direction, and the detectors 21 can be arranged more densely in the Y direction. Moreover, the formation area of the connection member AP on the wiring board 24 can be reduced, and the insulation can be improved.

Furthermore, as shown in FIG. 7, each conductive member 23 may be formed in a size that spans adjacent detectors 21 and shared between the detectors 21. With this configuration, the distance in the Y direction between the detectors 21 can be further reduced, and dense mounting of the detectors 21 on the wiring board 24 can be realized. In addition, since the pair of detectors 21 are connected by the conductive member 23,
The rigidity of the detection channel 21A is increased. Further, the mounting on the wiring board 24 can be easily positioned, and the mounting accuracy is improved.

Although not shown in the figure, a plurality of detector assemblies 41 are arranged such that the detector assemblies 41 adjacent in the X direction share the cathodes C located at both ends with one conductive member 22. It is also possible to integrate and configure a larger detector assembly 41 in the X direction. With this configuration, the adjacent conductive members 22 (cathodes C) in the detector assemblies 41 adjacent to each other in the X direction in FIG. 2A can be reduced one by one, and the reduced space can be used in the X direction. The number of detectors 21 can be increased. Therefore, dense mounting of the detector 21 on the wiring board 24 can be realized. This can be realized by using an even number of detection elements 211 constituting the detector assembly 41 and using cathodes C as the electrodes disposed at opposite ends of the detector assemblies 41.
Even when such a large detector assembly 41 in the X direction is used, the anodes A of the detectors 21 adjacent in the Y direction are electrically connected to each other via the connecting portions 23A and 23A of the conductive members 23 and 23. A plurality of large detection channels 21A can be configured by connecting to the. By using such a detection channel 21A, dense arrangement of the detectors 21 in the X direction and the Y direction is realized.

(Unit board)
The detailed structure of the unit substrate U will be described with reference to FIGS. Unit board U
The detection module 20A in which the detection channel 21A is installed as described above, and the ASIC
And a substrate 20B. The ASIC board 20 </ b> B includes a capacitor 26, a resistor 27, an analog ASIC 28, and a digital ASIC 29.

(Detection module)
As shown in FIG. 4, the detection module 20 </ b> A is configured by a plurality of detectors 21 installed on a wiring board 24 to form a detection channel 21 </ b> A. As described above, for example, a voltage of 500 V is applied between the anode A and the cathode C of the detector 21 in order to collect charges. This voltage is supplied from the power supply wiring (not shown) installed on the ASIC board 20B via the connector C1 to the power supply wiring (installed on the wiring board 24 of the detection module 20A).
(Not shown) between the anode A and the cathode C of each detector 21.
The detection module 20A has a connector C1 (see FIGS. 2A and 2B) at the end of the wiring board 24.
The same shall apply hereinafter. The connector C1 has the terminal 33 and a plurality of terminals 34. The γ-ray detection signal output from each detection channel 21A is transferred to the AS via the connector C1.
Supplied to the IC substrate 20B side.

(ASIC board)
As shown in FIGS. 2A and 2B, the ASIC substrate 20B includes a wiring substrate (support substrate) 35.
Eight (four on one side) analog ASICs 28 are installed on both sides. In addition, the wiring board 3
One digital ASIC 29 is installed on one side of 5. Further, capacitors 26 and resistors 27 are installed on both sides of the wiring board 35 by the number of detection channels 21A. Also, these capacitor 26, resistor 27, analog ASIC 28 and digital ASI
A plurality of connection wirings (not shown) for electrically connecting C29 are provided in the wiring board 35. These connection wirings have a laminated structure in the wiring board 35. The arrangement of the capacitor 26, the analog ASIC 28, and the digital ASIC 29 on the wiring board 35 is matched to the order in which the signals supplied from the detector 21 of the detection module 20A are transmitted. The resistor 27 has one end connected to the input side of the capacitor 26 and the other end connected to a ground wiring (not shown) provided on the wiring board 35. The analog ASIC 28 is an application specific IC that processes an analog signal (γ-ray detection signal) output from the detection channel 21A.
ASIC (Application Specific Integrated Circuit), which is a kind of LSI. The analog ASIC 28 is provided with a signal processing circuit (signal processing device) for each detection channel 21A. These signal processing circuits have one corresponding detection channel 2
The γ-ray detection signal (radiation detection signal) output from 1A is input to obtain the crest value of the γ-ray.
The ASIC board 20 </ b> B has a connector C <b> 2 having a plurality of terminals connected to each capacitor 26 at the end of the wiring board 35.

The unit substrate U is provided in the PET imaging apparatus 1 so that the surface on which the detector 21 is installed faces the depth direction of the PET imaging apparatus 1 (the longitudinal direction of the bed 31 in the Z direction in FIG. 2B). It is installed on an annular support member 1b (partially shown in FIG. 1B). The annular support member is provided so as to surround the measurement space 32. The plurality of unit substrates U installed on the annular member are arranged in the circumferential direction and surround the measurement space 32. The detection module 20A is arranged on the inner side (measurement space 32 side) and the ASIC substrate 20B is located on the outer side. In the present embodiment, the plurality of unit substrates U are also arranged in the depth direction of the PET imaging apparatus 1. In the unit substrate U thus installed, the direction of the X direction shown in FIG. 2A, FIG. 3C and the like is the circumferential direction of the PET imaging apparatus 1 (the circumferential direction of the annular support member),
The direction in the Y direction shown in FIGS. 2A and 3C is the radial direction of the PET imaging apparatus 1 (the radial direction of the annular support member).

(Connection structure between detection module and ASIC board)
As shown in FIG. 2B, the detection module 20A and the ASIC board 20B are assembled by overlapping their end portions and connecting the connector C1 and the connector C2 existing in the overlap portion. The end portion of the detection module 20A and the end portion of the ASIC substrate 20B are detachably coupled (separately and freely connectable) with a fastening screw or the like at the overlap portion. The reason for coupling at the overlap portion in this way is as follows. That is, the unit substrate U in which the detection module 20A and the ASIC substrate 20B are combined is cantilevered in the PET imaging apparatus 1, so that the central portion (connection portion) of the unit substrate U depends on the installation position.
A force that bends or bends the unit substrate U acts on. The connecting portion is the wiring board 24.
If the end surfaces of the wiring board 35 are in contact with each other, it is not preferable because the connection portion is easily bent or bent.

Based on this point, in the present embodiment, the detection module 20A and the ASIC substrate 20B are:
As described above, coupling is performed at the overlap portion. For this reason, the unit substrate U used in the present embodiment has improved toughness against bending and bending compared to connecting the end surfaces of the wiring substrate 24 and the wiring substrate 35 with each other. When the toughness of the unit substrate U with respect to bending and bending is improved, for example, the positional deviation of the detector 21 is suppressed and the accuracy of specifying the generation position of γ rays is improved. Incidentally, as shown in FIG. 1A, in the PET imaging apparatus 1, a large number of unit substrates U are arranged in the circumferential direction and the depth direction, so that the left and right side portions (particularly right lateral portions) in FIG. The unit substrate U located is easily bent or bent. For this reason, the toughness with respect to bending and bending of the unit substrate U is important. In particular, the detection channel 21A includes detectors 21 and 21 adjacent to each other in the Y direction shown in FIGS. 2 (a), 3 (c), and the like.
Are electrically connected to each other, thus improving the toughness against bending and bending of the unit substrate U, the electrical connection of the detection channel 21A is effectively maintained, and Since the mounting accuracy such as the dimension between the detectors 21 is ensured, it is difficult to be affected by the secular change and the like, and a highly accurate PET image can be obtained over a long period.

Such a detection module 20A and ASIC using the connector C1 and the connector C2
By using an electrical connection structure with the substrate 20B, the γ-ray detection signal is detected by the detection module 20A.
To the ASIC board 20B can be transmitted with low loss. Incidentally, when the loss is reduced, for example, the energy resolution as the detector 21 is improved.
Since the detection module 20A is detachably attached to the ASIC board 20B with screws or the like, for example, when a defect such as a detection failure occurs in the detector 21 or the ASICs 28 and 29,
Only the defective part (detection module 20A or ASIC board 20B) needs to be replaced. Note that the electrical connection between the detection module 20A and the ASIC board 20B is made by the connector C1 such as the spring pin connector described above, so that the connection / connection release (bonding / bonding release) between the boards is easy. Yes, mounting accuracy can be easily secured. In the above configuration, one detection module 20A is connected to the ASIC board 20B, but the detection module 20A may be divided into a plurality of parts.

It is preferable that the length of the circuit and the length (distance) of the wiring for transmitting the γ-ray detection signal are short because the influence of noise on the way and the signal attenuation are small. Further, when performing simultaneous measurement processing with the PET imaging 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). Therefore, in this embodiment, the detector 21, the capacitor 26, the analog ASIC 28, and the digital ASIC 29 are arranged in this order on the unit substrate U from the central axis toward the outside in the radial direction of the PET imaging apparatus 1. This configuration can shorten the length (distance) of the wiring that transmits the weak γ-ray detection signal output from the detector 21 to the analog ASIC amplifier. For this reason, the influence of noise on the γ-ray detection signal is reduced, and attenuation of the γ-ray detection signal is also reduced.
In addition, the capacitor 26 and the resistor 27 can be provided inside the analog ASIC 28. In order to obtain an appropriate capacitor capacity and an appropriate resistance value, and the analog ASIC 2
In the present embodiment, for example, the capacitor 26 and the resistor 27 are reduced.
Are arranged outside the analog ASIC 28. The capacitor 26, the resistor 27, and the analog ASIC 28 provided on the ASIC board 20B may be provided not on the ASIC board 20B but on the detection module 20A. In this case, the capacitor 26, the resistor 27, and the analog ASIC 28 are located closer to the ASIC board 20B than the detector 21. Since the detection module 20A includes the detector 21 and the analog ASIC 28, the distance (the length of the wiring) between the detector 21 and the analog ASIC 28 can be further reduced. For this reason, the influence of noise is further reduced.

(Operation of PET imaging device)
The operation of the PET imaging apparatus 1 having the above configuration will be described. Before performing a radiological examination, first, a radiopharmaceutical for PET (including 18F, for example) is administered to the subject H in advance so that its in-vivo radioactivity becomes, for example, about 370 MBq. The radiopharmaceutical is selected according to the examination purpose (for example, grasping the location of cancer or examining the aneurysm of the heart). The administered radiopharmaceutical gathers in the affected area of the subject H before long. In this state, the subject H is laid on the bed 31.
Examiners (medical radiographers and doctors) who perform PET examinations need information according to the purpose of the examination (areas (imaging areas or areas of interest) for which tomographic images are to be obtained, number of slices, slice interval, absorbed dose, etc.) Is input via the data processing device 2 (see FIG. 1A). In this case, a method of displaying an information input screen (not shown) on the display device 3 and inputting necessary data using a keyboard, a mouse, or the like can be employed. After that, the bed 31 is moved in the longitudinal direction,
The subject H is measured in the measurement space 32 until the examination site of the subject H (for example, an affected part of cancer) comes to a predetermined position.
Insert inside. Then, the PET imaging apparatus 1 is operated.

In response to an instruction from the data processing device 2, a DC high voltage is applied between the anode A and the cathode C of each detector 21 of the detection channel 21A, and the PET imaging device 1 starts a PET examination. The γ-ray emitted from the body of the subject H due to the radiopharmaceutical is detected by the detection channel 21A.
More specifically, it is detected by the detector 21 in the detection channel 21A. That is, P
When the positrons emitted from the radiopharmaceutical for ET are extinguished, a pair of γ rays are emitted in opposite directions of about 180 ° and detected by separate detection channels 21A. The detection channel 21A outputs a γ-ray detection signal. This detection signal is input to a corresponding signal processing circuit (not shown) in the corresponding analog ASIC 28 via the signal line 24b, the connectors C1 and C2, and the capacitor 26. This signal processing circuit amplifies the γ-ray detection signal and obtains a peak value of the detected γ-ray. This crest value is converted from an analog / digital converter (not shown) on the ASIC board 20B.
ADC) is converted into digital peak value information. The digital ASIC 29 further outputs position information of the detection channel 21A that has detected γ rays and detection time information of γ rays. Digital peak value information, position information of the detection channel 21A, and detection time information of γ rays are input to the data processing device 2. The simultaneous measurement device (not shown) of the data processing device 2 uses the detection time information to count a pair of γ-rays generated by the disappearance of one positron as one, and to detect the pair of γ-rays. The position of the detection channel 21A is specified based on the position information. Further, a tomogram information creation device (not shown), which is an image information creation device of the data processing device 2, uses the count value obtained by simultaneous measurement and the position information of the detection channel 21A, that is, the radiopharmaceutical accumulation position, The tomographic image information (image information) of the subject at the malignant tumor position is created.
This tomographic image information is displayed on the display device 3.

Below, the effect acquired in this embodiment is demonstrated.
(1) In the present embodiment, the detectors 21 and 21 adjacent in the Y direction are electrically connected to each other on the wiring board 24 to form one detection channel 21A. Can be used for one detection channel 21A formed long in the Y direction, and radiation is easily captured accordingly. In this embodiment, the detection channel 21A configured by electrically connecting the detectors 21 adjacent to each other in the Y direction in this way constitutes one detection unit (channel). The number of directional channels will be reduced. Since the number of detection channels 21A increases in the X direction, the number of channels in the X direction can be increased. As a result, the PET imaging apparatus 1 can increase the number of channels in the X direction without increasing the total number of channels. Thereby, the spatial resolution of the obtained image is improved. Further, since the total number of channels does not increase, the number of signal processing circuits does not increase, and the increase in heat generation of the signal processing circuits is suppressed. Thereby, the inspection time is shortened. Further, the power consumption is not increased due to the improvement of the spatial resolution, and the mass-produced detector assembly 41 can be used as it is, so that the resolution can be increased while suppressing the cost.
(2) In the detector assembly 41, a plurality of detector elements 211 having the cathode C on one surface of the semiconductor element S and the anode A on the other surface are arranged in parallel with the cathode C and the anode A aligned in the X direction. Therefore, the detector aggregate 41 is arranged in the X direction as a unit of the detection element 211 (
For example, it can be divided by a detector 21) composed of two detection elements 211. For this reason, radiation can be detected for each of the divided parts (for example, the detector 21). Therefore, the detection region can be subdivided in the X direction, and thereby the spatial resolution can be increased. Moreover, the detector 21 is adjacent to the detector 21 on the wiring board 24 in the Y direction.
Since one of the electrodes of the cathode C and the anode A is electrically connected to each other to form one detection channel 21A, a pair of adjacent detectors 21 are formed long in the Y direction. Each detection channel 21A can be used, and radiation can be easily captured in each detection region.
(3) Since the detector 21 is arranged so that the cathodes C or the anodes A of the detection elements 211 face each other, the conductive members 22 and 23 can be shared. Therefore, it is not necessary to arrange an electrical insulating material between the detection elements 211, and the dense arrangement of the detection elements 211 in the X direction can be realized. Thereby, the sensitivity is improved and the inspection time can be shortened.
(4) The conductive members 22 and 23 protrude outward from the detection element 211 and have overhang portions 22a and 23a that are part of the conductive members 22 and 23. The overhang portions 22a and 23a are attached to the wiring board 24. Therefore, the conductive members 22 and 23 can be easily attached to the wiring board 24.
(5) Since the overhang portions 22a and 23a are attached to the connection members CP and AP of the wiring board 24 by the adhesive 25, the time required for connecting the overhang portions 22a and 23a to the connection members CP and AP can be shortened. . The adhesive 25 exhibits both functions of electrical connection and mechanical connection (holding) between the detector 21 and the wiring board 24. Further, the terminal portions 22b, 23b suspended from the overhang portions 22a, 23a are formed narrow so that the overhang amount is smaller than the overhang amount of the overhang portions 22a, 23a. As a result, at the portions of the terminal portions 22b and 23b, γ
It is possible to minimize the line confusion. Further, since the terminal portions 22b and 23b are formed narrow, heat transmitted from the wiring board 24 to the detector 21 is reduced, and the stability of the detector 21 is improved.
(6) Since the conductive members 22 and 23 are made of a conductive metal that is a rigid conductive member, the conductive members 22 and 23 also function as protective members for the semiconductor element S. In particular, the semiconductor element S is CdT.
In the case of being made of a mechanically brittle semiconductor material such as e, CZT, GaAs, etc., the conductive member 22,
23 can prevent the semiconductor element S from being damaged.
(7) Since the overhang portions 22a and 23a protrude on two different side surfaces of the detector 21, the electrical insulation of the detection module 20A can be enhanced. Further, the detection channel 21A can be easily configured by connecting the protruding portions 23a of the adjacent detectors 21 to the connection member AP, respectively, and the detector 2 can be used when such a detection channel 21A is not configured.
1 can be constructed on the wiring board 24 by work that is not different from the work for mounting 1 on the wiring board 24. Therefore, the detector channel 21A can be used without any special mounting work.
Can be constructed on the wiring board 24.
(8) Since each detector 21 is arranged so that the overhang portions 22a and 23a face the Y direction, X
The gap between the detectors 21 in the direction can be narrowed, and the proportion of γ rays passing through the gap can be reduced. Thereby, the detection efficiency of γ rays can be increased, and the spatial resolution of the obtained image can be improved.
(9) In the X direction, the conductive members 22 having the same polarity are arranged on both sides of each detector 21 so as to face each other, so that the insulation between the detectors 21 in the X direction can be reduced.
The gap between the detectors 21 in the X direction can be made narrower. As a result, (8
), The detection efficiency of γ rays can be increased, and the spatial resolution of the image can be improved.
(10) Since all the detectors 21 adjacent in the Y direction are arranged so that the overhanging portion 22a or the overhanging portion 23a connected to the electrode of the same polarity (anode A or cathode C) faces each other, Y The insulation between the detectors 21 in the direction can be reduced, and the detector 2 in the Y direction can be reduced.
The gap between the ones can be narrowed. Therefore, the detection efficiency of γ rays can be increased, and the spatial resolution of the image can be improved. Here, the arrangement in which the overhanging portions 22a and 23a face each other means that the overhanging portions 22a and 23a of the mutual detectors 21 facing each other are not only completely facing each other but also the mutual detectors 21 facing each other. Is displaced in the X direction orthogonal to the Y direction, and the positions of the two overhang portions 22a and 23a are also displaced in the X direction. In particular,
In the detectors 21 adjacent to each other in the Y direction, the side surfaces where the protruding portions 23a protrude are opposed to each other, which is a state where the protruding portions 22a and 23a are opposed to each other.
(11) Since the detectors 21 are installed on both surfaces of the wiring board 24, the number of wiring boards 24 in the PET imaging apparatus 1 can be halved, and the arrangement density of the detectors 21 in the PET imaging apparatus 1 can be improved. For this reason, the gamma ray detection efficiency and the spatial resolution of the image of the PET imaging apparatus 1 can be further improved.
(12) Since the detection module 20A and the ASIC board 20B are detachably attached to each other, if one of them fails, the failed detection module 20A or ASIC board 20B can be easily replaced.
(13) In the PET imaging apparatus 1 using the detection channel 21A, each detection channel 21A
ASI with many built-in amplifier circuits (one per detector channel 21A)
Since the signal processing circuit is formed using C or the like, it is possible to cope with downsizing of the detection channel 21A and further increase in the number of detection channels 21A. As a result, the spatial resolution can be further improved.
(14) Since a detection module 20A in which a large number of detection channels 21A composed of detectors 21 with high energy resolution can be arranged can be configured, inspection with high quantitativeness can be performed in 3D imaging.
(15) By covering the detector 21 installed on the wiring board 24 with an electrical insulator, the dielectric breakdown of the detection channel 21A can be prevented.
(16) Since the protruding portions 22a in the two detection channels 21A adjacent in the Y direction are connected to the connection member CP (for example, the connection member CP1), the detection channel 21A in the Y direction
The arrangement density of can be improved. This also improves the detection efficiency of γ rays and the spatial resolution of the image. In addition, this configuration can shorten the length of the wiring board 24 in the Y direction, and can shorten the length of the PET imaging apparatus 1 in the radial direction. This leads to downsizing of the PET imaging apparatus 1. Also by the arrangement of the detector 21 shown in (10) above, the length of the wiring board 24 in the Y direction can be shortened. For this reason, the PET imaging apparatus 1 can be further shortened in the radial direction, and is more compact. In addition, the overhanging portions 22a, 2 of the detector 21
3a is placed on the connecting members AP and CP so as to overlap each other.
, CP, the distance between the adjacent detectors 21 in the Y direction can be further reduced, and the arrangement density of the detectors 21 on the wiring board 24 in the Y direction can be further improved. Thereby, size reduction of the PET imaging device 1 can be achieved. Furthermore, this makes it possible to reduce the area of the connection members AP and CP on the surface of the wiring board 24, so that dielectric breakdown between them can be prevented.
(17) Since the wiring 24a connected to the connection member CP and shared by the plurality of connection members CP is provided on the wiring board 24, the wiring density of the wirings 24a and 24b provided on the wiring board 24 can be reduced. Therefore, wiring on the wiring board 24 can be easily performed.
(18) Since the surface of the wiring board 24 on which the detector 21 is attached is arranged facing the longitudinal direction of the bed 31, the detector 21 is densely arranged in the radial direction (X direction) of the PET imaging apparatus 1. Can be placed. For this reason, the detection efficiency of γ rays and the spatial resolution of the image can be improved.
(19) As described above, since the distance between the detectors 21 in the Y direction (radial direction of the PET imaging apparatus 1) can be shortened, the distance between the subject H and the detector 21 on the rear side in the Y direction can be reduced. Shrink. This brings about an effect of improving the γ-ray detection sensitivity in the PET imaging apparatus 1.
(20) In this embodiment, two detection channels 21A are configured using a pair of detector assemblies 41 arranged adjacent to each other in the Y direction, and the detector assemblies 41 are attached to the wiring board 24. The work time for attaching the detector 21 to the wiring board 24 can be shortened.

In the embodiment, the overhanging portion 23a of the conductive member 23 connected to the anode A is connected to the connecting member A.
The protruding portion 22a of the conductive member 22 connected to the cathode C is connected to P (wiring) to the connecting member CP (wiring), but the protruding portion 23a is connected to the connecting member CP and the protruding portion 22a is connected to the connecting member CP. It is also possible to connect to the member AP. In this case, the cathode C serves as an electrode for outputting a γ-ray detection signal, and the anode A serves as an electrode for applying a bias voltage. If the voltage applied between the anode A and the cathode C is in the reverse direction, any pattern can be realized.
Further, the potential of the anode A is set to approximately ground, and the potential of the cathode C is set to −500V.
There is no restriction on the potential as long as it is in the reverse direction, and the voltage value may be set within a range that functions as the PET imaging apparatus 1. The cathode C can be used as a radiation detection signal extraction electrode, and the anode A can be used as a bias voltage application electrode.
Further, the cathodes C are arranged at both ends of the detector assembly 41, but the four detection elements 211 may be arranged so that the anode A is arranged at both ends of the detector assembly 41. Good.
Further, although the detector assembly 41 is configured by arranging four detection elements 211 in parallel, the number of parallel elements need not be limited to four. However, in order to improve the electrical insulation in the X direction, one detector assembly 41 may be configured by an even number of detection elements 211.
In the embodiment, as shown in FIGS. 4 to 7, the connection member AP and the connection member CP on the wiring board 24 are provided in addition to the portion to which the adhesive 25 is connected.
The connecting member CP may have a minimum necessary area to which the adhesive 25 is connected, and may be connected inside the wiring board 24 as necessary. Thereby, the electrical insulation between the connection member AP, the connection member CP, and the electrodes facing the connection member AP can be enhanced on the surface of the wiring board 24.
In the above-described embodiment, the detection channel 21A is configured by using a pair of detector assemblies 41 arranged adjacent to each other in the Y direction. However, the detector channels 41 are not adjacent to each other in the Y direction. The detection channel 21 </ b> A may be configured by using a pair of detectors 21 arranged in this manner. In this case, the detectors 21 of the detection channels 21A adjacent to each other in the X direction need to be arranged with a gap between them, so that the X direction can be compared with the above-described embodiment using the detector assembly 41. Although the density of the detector 21 is slightly reduced, the spatial resolution in the X direction is increased as compared with the conventional case.

In the above embodiment, the PET apparatus (see FIG. 1) has been described as an example of the nuclear medicine diagnosis apparatus. However, the present invention is not limited to the PET apparatus, but a single photon emission tomography apparatus (SPECT (Single Photon)
The detector and detection module of the present invention can also be applied to an Emission Computer Tomography apparatus) and a γ camera. Incidentally, the PET apparatus and the SPECT apparatus are common in taking a three-dimensional functional image of the subject, but the SPECT apparatus can perform simultaneous measurement because the measurement principle is to detect single photons. For this reason, a collimator that regulates the incident position (angle) of γ rays is provided. The γ camera includes a collimator that obtains 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 or SPECT apparatus, and X-ray CT.

(A) is the perspective view which showed typically the structure of PET apparatus which is suitable embodiment of this invention, (b) is the figure along the longitudinal direction of the bed of the PET imaging device in (a). (A) is a front view of the unit board | substrate used for the PET imaging device shown in FIG. 1, (b) is a side view of a unit board | substrate. (A) is a schematic perspective view of a detection element, (b) is a schematic view of a conductive member, and (c) is a perspective view of a semiconductor radiation detector assembly using the detection element. (A) is the figure which showed typically the arrangement | positioning state of the some radiation detector in the detection module of Fig.2 (a), (b) is sectional drawing of a detection module. It is the perspective view which showed the arrangement | positioning state of a radiation detector. (A) is the figure which showed typically the arrangement | positioning state of the several radiation detector in another detection module, (b) is sectional drawing of a detection module. It is the perspective view which showed the arrangement | positioning state of the other radiation detector.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 PET imaging device 2 Data processing device 3 Display apparatus 20A Detection module 20B ASIC board | substrate 21 Detector (radiation detector)
21A Detection channel 22, 23 Conductive member 22A, 23A Connection portion 22a, 23a Overhang portion 22b, 23b Terminal portion 22c Bending portion 24 Wiring board 28 Analog ASIC
29 Digital ASIC
31 Bed 32 Measurement space 211 Detection element 41 Detector assembly A Anode AP, CP Connection member C Cathode H Subject S Detection element U Unit substrate

Claims (19)

A radiation detection module comprising: a radiation detector; and a wiring board on which the radiation detectors are arranged and attached at least in the radiation traveling direction,
A radiation detection module, wherein a plurality of radiation detectors adjacent to each other in the radiation traveling direction are electrically connected to each other on the wiring board to form one detection channel. The radiation detection module according to claim 1, wherein the semiconductor detection element, which is a constituent requirement of the radiation detector, is configured of any one of CdTe, CdZnTe, and GaAs. A radiation detection module comprising: a radiation detector; and a wiring board to which the radiation detector is mounted in a plurality of directions arranged in a direction perpendicular to the radiation traveling direction and the radiation traveling direction,
The radiation detector is
A plurality of the semiconductor detection elements having a cathode electrode on one side of the semiconductor detection element and an anode electrode on the other side are arranged in parallel in a direction perpendicular to the radiation traveling direction,
A pair of radiation detectors attached to the wiring board and adjacent to each other in a radiation traveling direction, and electrically connecting one of the cathode electrode and the anode electrode of the radiation detectors to each other; A radiation detection module comprising a plurality of detection channels. The cathode electrode and the anode electrode of the pair of radiation detectors are electrically connected to each other via a conductive material provided on the wiring board. The radiation detection module according to claim 1 or 3. The radiation detectors adjacent to each other in a radiation traveling direction are attached on the wiring board in such a manner that the connection portions of the electrodes that are electrically connected to each other face each other. The radiation detection module according to claim 4. 5. The radiation detection module according to claim 1, wherein the connection portion is provided to protrude on a facing surface side between the adjacent radiation detectors. 6. The radiation detection module according to claim 4, wherein the conductive material is a wiring or a conductive member. A radiation detection module comprising: a radiation detector; and a wiring board to which the radiation detector is mounted in a plurality of directions arranged in a direction perpendicular to the radiation traveling direction and the radiation traveling direction,
The radiation detector is
A plurality of the semiconductor detection elements having a cathode electrode on one side of the semiconductor detection element and an anode electrode on the other side are arranged in parallel in a direction perpendicular to the radiation traveling direction,
A pair of radiation detectors attached to the wiring board and adjacent to each other in a radiation traveling direction, and electrically connecting one of the cathode electrode and the anode electrode of the radiation detectors to each other; With multiple detection channels
The connection part between the electrodes includes a projecting part projecting in the facing direction from the lower part on the facing surface side of the radiation detectors, and a terminal part suspended from the projecting part,
The radiation detector module, wherein the radiation detector is attached in a non-contact state to the substrate surface of the wiring board by the terminal portion. 9. The radiation detector according to claim 1, wherein a plurality of the semiconductor detection elements are arranged in parallel so that the same poles are alternately opposed to each other in a direction orthogonal to a radiation traveling direction.
The radiation detection module according to any one of the above. The radiation detection module according to any one of claims 3 to 9, wherein the semiconductor detection element is made of any one of CdTe, CdZnTe, and GaAs. A wiring board, a plurality of detection channels provided in a first area of the wiring board and including a plurality of radiation detectors, and provided in a second area of the wiring board and separately connected to each of the detection channels. A signal processing device,
A plurality of the radiation detectors are arranged in a first direction from the first region to the second region, and in a second direction orthogonal to the first direction, respectively.
The printed circuit board, wherein the detection channel is configured by electrically connecting a pair of radiation detectors adjacent in the first direction. The printed circuit board according to claim 11, wherein the electrodes are connected by a conductive material provided on the wiring board. The printed circuit board according to claim 11, wherein the electrodes are electrically connected to each other at opposing surfaces of the semiconductor detection elements adjacent to each other in the radiation traveling direction. The printed circuit board according to claim 13, wherein the semiconductor detection element is made of any one of CdTe, CdZnTe, and GaAs. A plurality of detector assemblies including a plurality of radiation detectors arranged in the second direction are arranged in the first direction and the second direction, respectively.
The detector assembly is configured by connecting electrodes of the same polarity of a plurality of radiation detectors,
The detection channel is configured by electrically connecting the radiation detectors included in each of the pair of detector assemblies adjacent in the first direction, and is formed in at least two of the pair of detection assemblies. The printed circuit board according to claim 11, wherein: The wiring board has a first wiring board on which the plurality of detection channels are arranged, and a second wiring board on which the plurality of Shingo processing devices are arranged,
The printed circuit board according to any one of claims 12 to 15, wherein the first wiring board and the second wiring board are detachably coupled to each other. A plurality of radiation detectors, and a plurality of radiation detection modules having a wiring board to which the radiation detectors are attached, surrounding a measurement region in which a bed supporting a subject is inserted and arranged around the measurement region A nuclear medicine diagnostic apparatus comprising an image information creation device that generates an image using information obtained based on a radiation detection signal output from the radiation detector,
A nuclear medicine diagnosis apparatus comprising a plurality of detection channels configured such that a plurality of radiation detectors arranged adjacent to each other in the radial direction of the measurement region are electrically connected to each other. A plurality of radiation detectors, and a printed circuit board having a wiring board to which the radiation detectors are attached, surrounding a measurement region into which a bed supporting a subject is inserted, and arranged around the measurement region; A nuclear medicine diagnostic apparatus comprising an image information creation device that generates an image using information obtained based on a radiation detection signal output from the radiation detector,
The printed circuit board is
A wiring board, a plurality of detection channels provided in a first area of the wiring board and including a plurality of radiation detectors, and provided in a second area of the wiring board and separately connected to each of the detection channels. A signal processing device,
A plurality of the radiation detectors are arranged in a first direction from the first region to the second region, and in a second direction orthogonal to the first direction, respectively.
The nuclear medicine diagnostic apparatus, wherein the detection channel is configured by electrically connecting a pair of radiation detectors adjacent in the first direction. The nuclear medicine diagnosis apparatus according to claim 17 or 18, wherein the semiconductor detection element is made of any one of CdTe, CdZnTe, and GaAs.

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