JP2007041007A - Radiographic examination device having pet imaging apparatus and x-ray ct imaging apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiographic examination device whose energy resolution can be improved, and which can perform highly precise diagnosis, and produces a reduced coercive feeling to a subject. <P>SOLUTION: The radiographic examination device comprises a bed 14 on which the subject H is mounted, and first and second imaging apparatuses 1 and 4 arranged in the longitudinal direction of the bed 14. The first imaging apparatus 1 has a plurality of semiconductor radiation detectors 21 for detecting γ-rays discharged from the subject H around the bed 14. The second imaging apparatus 4 comprises an X-ray source 45 for radiating X-rays to the subject H and a radiation detector 40 for detecting X-rays from X-ray source 45 having transmitted through the subject H. The bed 14 is shared by the first imaging apparatus 1 and second imaging apparatus 4. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

  The present invention relates to a radiation inspection apparatus using radiation, and is particularly suitable for performing a plurality of types of radiation inspection such as X-ray CT and positron emission computed tomography (hereinafter referred to as “PET”). The present invention relates to a radiation inspection apparatus.

The inspection technique using radiation can inspect the inside of a subject nondestructively. In particular, radiation inspection techniques for the human body include X-ray CT, PET, SPECT single photon emission computed tomography (hereinafter referred to as “SPECT”), and the like.
Each of these techniques is a technique for measuring a physical quantity to be inspected as an integral value in the radiation flight direction, and calculating and imaging the physical quantity of each voxel in the subject by back projecting the integral value. These technologies need to process enormous amounts of data, and with the rapid development of computer technology in recent years, it has become possible to provide high-speed, high-detail images.

  The X-ray CT technique is a technique for measuring the intensity of X-rays that have passed through a subject and imaging the morphological information of the subject from the passage rate of X-rays in the body. The subject is irradiated with X-rays from an X-ray source, the X-ray intensity that has passed through the body is measured by a detection element disposed on the opposite side of the subject, and the distribution of the integral absorption coefficient of the subject is measured. Based on this integral absorption coefficient, the filtered back projection method (Filtered Back Projection Method) described on page 21 of IEEE Transactions on Nuclear Science NS-21 volume is used. Thus, the absorption coefficient of each voxel is obtained and the value is converted into a CT value. A radiation source often used for X-ray CT is about 80 keV.

On the other hand, PET and SPECT are techniques capable of detecting functions and metabolism at a molecular biology level that cannot be detected by X-ray CT or the like, and can provide a functional image of the body. PET is a technique in which a radiopharmaceutical labeled with positron emitting nuclides such as 18F, 15O, and 11C is administered, and its distribution is measured and imaged. Drugs include fluorodeoxyglucose (2- [F-18] fluoro-2-deoxy-D-glucose, 18FDG), which uses the high accumulation in tumor tissues by sugar metabolism, Used to identify
The radionuclide taken into the body decays and releases positron (β +). The emitted positron emits a pair of annihilation γ-rays having energy of 511 keV when it annihilates by combining with electrons. Since this annihilation γ-ray pair is radiated in almost opposite directions (180 ° ± 0.6 °), the annihilation γ-ray pair is detected at the same time by a detection element arranged so as to surround the subject, and the radiation direction data thereof Can be stored to obtain projection data. By backprojecting the projection data (using the filtered back projection method or the like), it becomes possible to identify and image the radiation position (radiation nuclide accumulation position).

  SPECT is a technique in which a radiopharmaceutical labeled with a single photon emitting nuclide is administered and its distribution is measured and imaged. Single γ-rays having energy of about 100 keV are emitted from the drug, and the single γ-rays are measured by a detection element. In the measurement of a single gamma ray, the flight direction cannot be identified, so in SPECT, a collimator is inserted in front of the detection element, and projection data is obtained by detecting only gamma rays from a specific direction. Similar to PET, image data is obtained by back projecting projection data using a filtered back projection method or the like. The difference from PET is that there is no need for simultaneous measurement due to the measurement of a single γ-ray, the number of detection elements is small, etc., and the apparatus configuration is simple and relatively inexpensive. On the other hand, since SPECT uses a collimator, the detection rate of γ rays is low and the image quality is generally poor.

As described above, PET and SPECT can extract a region where medicines are accumulated with good contrast in order to obtain a functional image using metabolism in the body, but there is a problem that the positional relationship with surrounding organs cannot be grasped. Thus, in recent years, attention has been focused on a technique for performing more advanced diagnosis by merging morphological images such as X-ray CT and functional images such as PET and SPECT (see Patent Document 1).
Japanese Patent Laid-Open No. 7-20245 (paragraph number 0010, FIG. 1)

By the way, in a radiation inspection apparatus that fuses conventional morphological images such as X-ray CT and functional images such as PET and SPECT, a scintillator is used as a γ-ray detector in order to obtain functional images such as PET and SPECT. Yes. The scintillator performs a process of once converting incident γ-rays into visible light and then converting it back to an electric signal by a photomultiplier tube (photomal). The scintillator has a problem that the number of photons generated at the time of visible light conversion is small and the two-step conversion process is required as described above, so that the energy resolution is low and a high-precision diagnosis cannot always be performed. Was. The decrease in energy resolution is a cause that quantitative evaluation cannot be performed particularly during 3D imaging of PET. This is because, since the energy resolution is low, the energy threshold of γ-rays must be lowered, and a lot of internal scattering, which is noise that increases during 3D imaging, is detected. Therefore, a scintillator type PET apparatus generally has a 2D imaging function in order to perform a highly quantitative test. In this 2D imaging, a scepter is inserted inside the scintillator to prevent the incidence of γ rays from other than the 2D region.
Further, in order to obtain a more accurate image in the PET inspection, it is necessary to perform absorption correction. Compared to the body surface, γ rays from the deep part of the body are more likely to be absorbed in the body, and by correcting the amount of γ rays absorbed in the body (absorption correction), high image quality and high quantification can be achieved. In a 1 gantry-type PET apparatus, it is necessary to hold a radiation source in order to perform absorption correction, and perform measurement while rotating the radiation source inside the scintillator.
Therefore, in a scintillator type PET apparatus, the size of the gantry is increased by a photomultiplier, a scepter, and an absorption correction radiation source. Furthermore, in a conventional radiological inspection apparatus in which the arrangement of X-ray CT and PET, SPECT, etc. are arranged in series, Furthermore, it was easy to enlarge, and there was a difficulty of giving a sense of intimidation to the subject. Therefore, it is necessary to solve such a problem.
It is an object of the present invention to provide a radiation inspection apparatus that can improve energy resolution and perform highly accurate diagnosis.

In order to solve the above problems, in the first invention, the detection accuracy is increased by using a semiconductor radiation detector for the first imaging device (PET imaging device). In this configuration, since the radiation is directly detected using the semiconductor radiation detector, the position resolution and the energy resolution can be improved. Moreover, since energy resolution can be improved, internal scattering is removed (noise is reduced). Accordingly, the quantitativeness at the time of 3D imaging can be improved, and a septum for 2D imaging is not required, and the apparatus can be miniaturized. Further, the detection element can be downsized, and a radiation inspection apparatus that contributes to downsizing of the entire apparatus can be obtained.
In addition, since the absorption correction can be performed using the second imaging device (X-ray CT imaging device), it is not necessary to provide the first imaging device with an absorption correction radiation source (gamma rays or the like), so that it can be further improved. The size of the apparatus can be reduced.

  In the second invention, a semiconductor radiation detector is used as the radiation detector of the second imaging device (X-ray CT imaging device). In this configuration, it is possible to reduce the size of the detection element, thereby obtaining a radiation inspection apparatus that contributes to downsizing of the entire apparatus.

  In the third invention, the first imaging device (PET imaging device) is formed to be smaller in shape than the second imaging device (X-ray CT imaging device), and is further on the near side than the second imaging device. Since it is arranged, the impression of the entire apparatus received by the subject can be made less intimidating.

  According to the present invention, energy resolution can be improved and highly accurate diagnosis can be performed.

  Next, a radiation inspection apparatus according to a preferred embodiment of the present invention will be described in detail with reference to the drawings as appropriate. In the following description, the PET imaging apparatus that is the first imaging apparatus constituting the radiation inspection apparatus of the present embodiment, the description of the X-ray CT imaging apparatus that is the second imaging apparatus, and the implementation of the semiconductor radiation detector and the like. The elements applied to the form will be described.

[Embodiment 1]
As shown in FIGS. 1 and 2, the radiation inspection apparatus according to the present embodiment includes a bed 14 that is a bed apparatus, a PET imaging apparatus 1 that is a first imaging apparatus, and an X-ray CT imaging that is a second imaging apparatus. As shown in FIG. 1, it includes a data processing device 12, a display device 13, and the like. A subject (subject) H is placed on a bed 14 configured to be movable back and forth in the body axis direction (X1, X2 direction) of the subject H, and the PET imaging apparatus 1 and the X-ray CT imaging apparatus 4. It is supposed to be taken with.

A. (PET imaging device)
The PET imaging apparatus 1 includes a large number of semiconductor radiation detectors 21 (see FIGS. 4, 8, and 11), and γ rays emitted from the body of the subject H are detected by a semiconductor radiation detector (hereinafter simply referred to as “semiconductor radiation detector”). It is detected by 21). The PET imaging apparatus 1 is provided with an integrated circuit (ASIC) for measuring the peak value and detection time of the γ-ray, and measures the peak value and detection time of the detected radiation (γ-ray). ing.

As shown in FIG. 3, in the PET imaging apparatus 1, in order to detect γ-rays emitted from the subject H, a combined substrate 20 having a large number of detectors 21 (details are shown in FIGS. 8A and 8B). A large number of detector units 2 (see FIG. 11 for details) are housed in a circumferential shape. A subject H who is a subject lies on the bed 14 and is positioned at the center of the PET imaging apparatus 1. At this time, each detector 21 surrounds the periphery of the bed 14. From the detector unit 2, the peak value information of the γ-ray, the time information of the γ-ray detection obtained based on the detection signal when the detector 21 interacts with the γ-ray, and the address information of the detector 21. (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.
Incidentally, the subject H is administered with a radiopharmaceutical, for example, fluorodeoxyglucose (FDG) containing 18 F with a half-life of 110 minutes. From the body of the subject H, gamma rays (annihilation gamma rays) are emitted when the positrons emitted from the FDG disappear.
As shown in FIG. 1, such a PET imaging device 1 is formed so that the shape of the housing is smaller than the shape of the housing of the X-ray CT imaging device 4 to be described later, and the body axis direction of the subject H In the direction of arrows X1 and X2 in FIG. 1, the X-ray CT imaging apparatus 4 is arranged on the near side.
Next, details of the PET imaging apparatus 1 will be described.

(Semiconductor radiation detector)
First, the detector 21 applied to this embodiment will be described. As shown in FIG. 4, the detector 21 has thin plate (film-like) electrodes (anode A, cathode C) on both sides of a semiconductor radiation detection element (hereinafter referred to as detection element) 211 made of a plate-like semiconductor material S. (Minimum configuration). Among these, the semiconductor material S is composed of any single crystal such as CdTe (cadmium telluride), TlBr (thallium bromide), GaAs (gallium arsenide), or the like. The electrodes (anode A, cathode C) are made of any material such as Pt (platinum), Au (gold), and In (indium). In the following description, it is assumed that the semiconductor material S is obtained by slicing a single crystal of CdTe. Further, it is assumed that the radiation to be detected is 511 KeV γ rays used in the PET imaging apparatus 1.

  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 the γ rays interact with the semiconductor material S constituting the detector 21, holes schematically indicated by “+” and “−” in the figure. (Holes) and electrons (electrons) are generated in pairs in an amount proportional to the energy of γ 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, the ease of movement (mobility) is relatively larger for electrons, so that electrons reach the anode in a relatively short 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.

  As shown in FIG. 5 which compares the “time-crest value curve” when the semiconductor material S (detection element 211) of the detector 21 is thick and thin, the semiconductor material S with the smaller thickness t is shown. However, the peak value rises (rises) quickly, and the peak value is high. 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). This is because the time to reach (C) (charge collection time) is shortened. In addition, holes that may have disappeared in the middle can reach the electrode (cathode C) without disappearing due to the short distance. Incidentally, the thickness t can also be expressed as a distance between the electrodes of the anode A and the cathode C facing each other.

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 is slowed and the maximum value of the crest value is also lowered. 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. Moreover, the measurement of the peak value can be performed more correctly.

  Incidentally, in the case of the PET imaging apparatus 1, since simultaneous measurement is performed, it is a problem to correctly measure the detection time of γ rays. For example, in FIG. 4, 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 large, 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 imaging apparatus 1 that performs simultaneous measurement in nsec (nanosecond) order. Therefore, also from this meaning, the detection time can be appropriately determined within the above-described thickness range. The detection time in the PET imaging apparatus 1 is determined by the LET method or the CFD method.

  As shown in FIG. 6 schematically showing the relationship between the thickness t of the semiconductor material S of the detector 21 and the peak value (maximum value), the maximum peak 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. 7, 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 having the laminated structure shown in FIG. 7 is not configured to detect the radiation independently of the other layers because each of the anodes A and the cathodes C is connected in common. . 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. In order to increase the interaction between the semiconductor material S and the γ-rays while increasing the charge collection efficiency and reducing the amount of the γ-rays passing through. Yes (to increase the number of counts).
By using the detector 21 having such a laminated structure, a better crest value rising speed (rise) and a more accurate crest value can be obtained, and the number of γ-rays that interact with the semiconductor material S (count) Number) can be increased (sensitivity can be increased).
Note that the detector 21 is not necessarily required to have such a laminated structure, and may be a single layer, or may have a two-layer to four-layer structure so as to have an appropriate layer structure.

The area s of the electrodes (anode A, cathode C) is preferably 4 to 120 square mm. The increase in the area s increases the capacitance (stray capacitance) of the detector 21, and noise increases easily due to the increase in the stray capacitance. Therefore, the electrode area s should be as small as possible. In addition, since the charge generated at the time of detecting γ-rays is partially accumulated in the stray capacitance, if the stray capacitance is increased, the amount of charge accumulated in the charge amplifier 24b of the analog ASIC 24 (see FIG. 9) and thus the output voltage (crest value). There is a problem that 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 imaging apparatus 1. The position resolution of the PET imaging apparatus 1 is determined by the range of positrons in addition to the size (arrangement pitch) of the detectors 21, but the range of the positrons of 18F is 2 mm. It is meaningless even if it is 2 mm or less. The mounting method in which the electrode area is the smallest is when the electrode surface is placed perpendicular to the radial direction of the PET imaging apparatus 1. From the above consideration, the lower limit value of one side of the electrode is 2 mm, and the lower limit value of the electrode area s is 4 square mm.

  In the above description, the semiconductor material S that interacts with γ rays is CdTe. However, it goes without saying that the semiconductor material S may be TlBr, GaAs, or the like. In addition, although the term “layered structure” or “upper layer / lower layer” is used, this term is based on FIG. 7, and when the viewing direction is 90 degrees sideways, the layered structure is, for example, a parallel structure, For example, it can be read as left and right. Further, the direction in which the γ-rays are incident may be upward / downward, left / right in FIG. In other words, the detector 21 has a structure in which a plurality of (for example, five) semiconductor materials S are arranged in parallel with the cathodes C and anodes A being alternately sandwiched therebetween.

(Bonded substrate)
The detailed structure of the combined substrate (unit substrate) 20 installed in the detector unit 2 (FIG. 11) will be described with reference to FIGS. The coupling substrate 20 includes a detector substrate (first substrate) 20A on which a plurality of detectors 21 are installed, a capacitor 22, a resistor 23, an analog ASIC 24, and an analog / digital converter (AD converter, hereinafter referred to as ADC). 25 and the ASIC board (second board) 20B on which the digital ASIC 26 is installed.

(Detector board)
8 A detector substrate 20A on which the detector 21 is installed will be described with reference to each drawing. As shown in FIG. 8A, 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 columns 21 = 16 horizontal x 4 vertical, for a total of 64). In the radial direction of the PET imaging apparatus 1, 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 PET imaging apparatus 1, that is, in the longitudinal direction of the bed 14 (advance / retreat direction). Further, as shown in FIG. 8B, since the detectors 21 are installed on both sides of the detector substrate 20A, a total of 128 detectors 21 are installed on one detector substrate 20A. Will be. Here, as the number of detectors 21 to be installed increases, it becomes easier to detect γ-rays, and the position accuracy at the time of γ-ray detection can be increased. For this reason, the detector 21 is installed on the detector substrate 20A as closely as possible. Incidentally, γ-rays emitted from the subject H (see FIG. 3) on the bed 14 are upward from below on the paper surface of FIG. 8A (in the direction of the arrow 32, that is, the radial direction of the PET imaging apparatus 1). If the arrangement of the detectors 21 in the left-right direction on the detector substrate 20A is made dense, the number of γ rays that pass through (the number of γ rays that pass through the gap between the detectors 21) can be reduced. It is preferable because it is possible. This increases the detection efficiency of γ rays, and can increase the spatial resolution of the obtained image.
As shown in FIG. 8B, the detector substrate 20A according to the present embodiment has the detector 21 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. 11) 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. 8C, each detector 21 has a stacked structure in which single crystals of the thin-plate semiconductor material S (detecting element 211) described above are stacked. The configuration and operation of this point are as already described with reference to FIG. 7, 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 supplied from the ASIC board 20B side to the detector board 20A side via the connector C1 (FIG. 8A). 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, the detector 21 may be provided so that each detection element 211 is perpendicular to the substrate body 20a.

(ASIC board)
Next, the ASIC substrate 20B on which the ASIC is mounted will be described. As shown in FIG. 8A, the ASIC substrate 20B is provided with two analog ASICs 24 and one digital ASIC 26 on one surface of the substrate body 20b. Further, as shown in FIG. 8B, since the analog ASICs 24 are installed on both surfaces of the board body 20b, one ASIC board 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. In addition, capacitors 22 and resistors 23 corresponding to the number of detectors 21 are provided on both surfaces of one substrate body 20b. Further, in order to electrically connect the capacitor 22, the resistor 23, the analog ASIC 24, the ADC 25, and the digital ASIC 26, the ASIC board 20B (board body 20b) has a board (not shown) similar to the detector board 20A described above. Internal wiring is provided. This intra-substrate wiring also has a laminated structure.
In the arrangement (in-board wiring) of these elements 22, 23, 24, 25, and 26, signals supplied from the detector board 20A are supplied in the order of the capacitor 22, the resistor 23, the analog ASIC 24, the ADC 25, and the digital ASIC 26. It has become so.
The ASIC board 20B includes a connector (spiral contact) C1 that is connected to an in-board wiring connected to each capacitor 22 and electrically connects to the detector board 20A, and a data processing apparatus side (unit integration described later). And a board connector C2 for electrical connection to the FPGA side). 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. The analog ASIC means an application specific integrated circuit (ASIC) that is an application specific IC that processes an analog signal, and is a kind of LSI.

(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) of each other, but as shown in FIG. The connectors C1 existing in the overlap portion are connected to each other. This connection is made detachable (separate and connectable) with a fastening screw or the like. Such a 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. 3, in the PET imaging apparatus 1, a large number of detector units 2 (see FIG. 11) including the coupling substrate 20 shown in FIG. The coupling substrate 20 positioned in the 3 o'clock direction or 9 o'clock direction corresponding to the reference horizontal direction 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, a connector (connector) C1 (FIG. 8) for electrically connecting the in-board wirings of both the boards 20A and 20B to the overlapping portions of the detector board 20A and the ASIC board 20B shown in FIG. 8 (a)) is provided. 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 as 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, even when a failure occurs in the detector 21 or the ASICs 24 and 26, only the defective portion 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 performed by the connector C1 such as the spiral contactor (registered trademark) described above. ) Is easy.

  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, the configuration may be such that eight detectors 21 in the horizontal direction and four detectors 21 in the vertical direction are mounted on one board, and 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 board on which the two 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. 9, the detector 21 is connected to the analog ASIC 24 via a connector C1, a capacitor 22 and a resistor 23 by electric 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. 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). For this reason, in this embodiment, as shown in FIG. 8A from the central axis toward the outside in the radial direction of the PET imaging apparatus 1, a detector 21, a capacitor 22, a resistor 23, an analog ASIC 24, an ADC 25, a digital The elements 21, 22, 23, 24, 25, and 26 are arranged (layout) in the order of the ASIC 26. This order is the same as the signal processing order by the elements 21, 22, 23, 24, 25, and 26 (see FIGS. 9 and 10). 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”. In other words, the arrangement order of the elements 21, 22, 23, 24 from the detector 21 to the digital ASIC 24 matches the processing order of signal processing by the elements 21, 22, 23, 24. 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. That is, in consideration of noise, there is no problem even if the wiring length after the analog ASIC 24 is increased. However, since the signal transmission is delayed when the wiring length is increased, the accuracy of the detection time may be impaired as described above.
In the present embodiment, since not only the detector 21 but also the analog ASIC 24 and the digital ASIC 26 are included in one combined substrate 20, the detector 21, the analog ASIC 24, and the digital ASIC 26 are moved in the longitudinal direction of the bed 14, that is, inspected. Since it can be arranged in a direction orthogonal to the body axis of the subject H to be received, the length in the longitudinal direction of the bed 14 of the PET imaging apparatus 1 does not have to be longer than necessary. Although it may be possible to arrange the analog ASIC 24 and the digital ASIC 26 in the longitudinal direction of the bed 14 outside the radial direction of the group of detectors arranged in a ring shape, the length of the PET imaging apparatus 1 in the longitudinal direction of the bed 14 is necessary. More than that. Further, since the semiconductor radiation detector is used as the detector 21 and the analog ASIC 24 and the digital ASIC 26 are used as the signal processing device, the length in the longitudinal direction of the coupling substrate 20 is shortened, and PET is compared with the case where a scintillator is used. The length of the imaging device 1 in the orthogonal direction can be significantly reduced. Furthermore, since the detector 21, analog ASIC 24, and digital ASIC 26 are sequentially arranged in the longitudinal direction of the coupling substrate 20, the length of the wiring connecting them can be shortened, and the wiring on the substrate is simplified. Therefore, the PET imaging apparatus 1 that contributes to downsizing is obtained.

Here, in the present embodiment, each analog ASIC 24 is connected to 32 detectors 21 and processes signals obtained from the detectors 21. As shown in FIGS. 9 and 10, 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 charge amplifier (preamplifier) 24b, a polarity amplifier (linear amplifier) 24c, a bandpass filter (waveform shaping device) 24d, and a peak hold circuit for the purpose of obtaining the peak value of the detected γ-ray. (Peak value holding device) 24e is provided connected in this order. Incidentally, the slow system is named “slow” because it takes a certain amount of processing time to obtain the peak value. The γ-ray detection signal output from the detector 21 and passed through the capacitor 22 and the resistor 23 is amplified by the charge amplifier 24b and the polarity amplifier 24c. 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. One analog ASIC 24 is obtained by converting 32 sets of analog signal processing circuits 33 into an LSI.
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.

  With respect to the analog ASIC 24 shown in FIG. 9, in this embodiment, the slow output of the analog ASIC 24 is supplied to an ADC (analog-digital converter) 25. Further, the first output of the analog ASIC 24 is supplied to the digital ASIC 26.

  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 ASIC 26 are connected by 32 wires that transmit 32ch fast signals one by one. That is, one digital ASIC 26 is connected to four analog ASICs 24 by a total of 128 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. 5). The fast output signal output from the analog ASIC 24 to the digital ASIC 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). Note that the ADC 25 is connected to the digital ASIC 26 by a wiring for transmitting an ADC control signal and a wiring for transmitting peak value information.

As shown in FIG. 10, the digital ASIC 26 includes a plurality of packet data generation devices 34 including eight time determination circuits (time determination devices) 35 and one ADC control circuit (ADC control device) 36, and a data transfer circuit. (Data transmission device) 37 is provided, which is an LSI. All the digital ASICs 26 provided in the PET imaging 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 receives a timing signal for detecting γ rays from the time determination circuit 35 and specifies its detector ID. That is, the ADC 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 ADC control circuit 36 stores the detector ID in the time determination circuit 35. The corresponding 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. Packet data (including detector ID, time information, and peak value information) output from the ADC control circuit 36 of each packet data generation device 34 is input to the data transfer circuit 37.
The data transfer circuit 37 converts the packet data, which is digital information, output from the ADC control circuit 36 of each packet data generation device 34 into, for example, the detector unit 2 (see FIG. 11 and FIG. 12) is transmitted to a unit integration integrated circuit (unit integrated FPGA (Field Programmable Gate Array) 31) provided in one case 30 of FIG. The unit integrated FPGA (hereinafter referred to as “FPGA”) 31 outputs the digital information to the information transmission wiring connected to the connector 38.

  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 PET data processing unit 12a (see FIG. 16) of 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 PET data processing unit 12a performs simultaneous measurement processing (when two γ rays having a predetermined energy are detected in a time window of a set time, these γ Two detectors 21 that detect the pair of γ-rays by counting the pair of γ-rays measured at the same time as a pair of γ-rays generated by the annihilation of one positron. Are identified from their detector IDs. When there are three or more γ-ray detection signals detected within the above time window (there are three or more detectors 21 that have detected γ-rays), the PET data processing unit 12a Using the peak value information or the like, the two detectors 21 in which the γ-rays are incident first among the three or more detectors 21 are specified. The specified pair of detectors 21 are simultaneously measured to generate one count value (first information). Further, the tomographic image information creation device 12B of the PET data processing unit 12a uses the count value obtained by the simultaneous measurement and the position information (second information) of the detector 21 at the radiopharmaceutical accumulation position, that is, the malignant tumor position. Tomographic image information of the subject 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 26. 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. Since the detector substrate 20A includes the detector 21 and the analog ASIC 24, the distance (the length of the wiring) between the detector 21 and the analog ASIC 24 can be further shortened. For this reason, the influence of noise can be further reduced.
Furthermore, the coupling substrate 20 may be configured such that three substrates (detector substrate 20A, analog ASIC substrate, and digital ASIC substrate) are detachably connected via connectors. The detector board 20A includes a detector 21, the analog ASIC board includes a capacitor 22, a resistor 23, and an analog ASIC 24, and the digital ASIC board includes an ADC 25 and a digital ASIC 26. In this configuration, by separating the substrate on which the analog circuit is mounted from the substrate on which the digital circuit is mounted, noise on the digital circuit side is prevented from entering the analog circuit. In this configuration, since 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 20A) on which the detector 21 is installed 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 a substrate via a BGA (Ball Grid Array) by reflow, only the ASIC substrate can be soldered, which is preferable because the detector 21 does not need to be exposed to high temperatures. Of course, it is possible to avoid using the connector C1 by arranging all the elements 21 to 26 on the same substrate.

(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 combined substrates 20 are housed in a housing (frame body) 30 to constitute a detector unit (12 substrate units) 2. Incidentally, the PET imaging apparatus 1 has a configuration in which 60 to 70 detector units 2 are detachably arranged in the circumferential direction (see FIG. 13B), so that maintenance and inspection are easy. Yes.

(Storing in a housing)
As shown in FIG. 11, the detector unit 2 includes twelve coupling substrates 20, a high-voltage power supply device PS that supplies a voltage for collecting charges to the twelve coupling substrates 20, the above-described FPGA 31, A signal connector for transmitting and receiving signals, a power supply connector for receiving power supply from the outside, and the like are housed and held.

  11 and 12, the combined substrates 20 are arranged in three rows so as not to overlap in the depth direction (longitudinal direction of the bed 14), and in four rows in the frontage direction (circumferential direction of the PET imaging device 1). In the housing 30. That is, 12 coupling substrates 20 are accommodated in one housing 30. In order to store in this way, a guide member 39 having four rows of guide grooves (guide rails) G1 extending in the depth direction and spaced apart from each other in the circumferential direction is arranged in the housing 30 and is provided at the upper end of the housing 30. It is attached. 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 a single 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. 12). The guide grooves G <b> 1 and G <b> 2 have a depth enough to accommodate three (three) coupling substrates 20. An end portion of the coupling substrate 20 on the ASIC substrate 20B side is accommodated in the guide groove G1, and an end portion of the coupling substrate 20 on the detector substrate 20A side 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 and G2. It can be easily positioned at a predetermined location by sliding inside G2. 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. Note that the upper and lower portions of the casing 30 are when the casing 30 is taken out from the PET imaging apparatus 1, and when the casing 30 is provided in the PET imaging apparatus 1 as shown in FIG. In some cases, the top and bottom are reversed, the top and bottom are rotated 90 ° to the left or right, or are inclined.

  As shown in FIG. 12, the top plate 30a of the housing 30 is provided with an FPGA 31 and a connector 38 in addition to the four rows of guide grooves G1 described above. The connector 38 is connected to the FPGA 31. The FPGA 31 can create a program on site. This is different from the ASIC that cannot be programmed. Therefore, as in the present embodiment, in the FPGA 31, for example, even when the number and type of the combined substrates 20 to be accommodated are changed, it is possible to appropriately cope with the change in the number 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. 13A, the detector unit 2 is mounted via a unit support member 2A. Further, as shown in FIG. 13B, the detector unit 2 is mounted on the PET imaging apparatus 1 while being supported on one end 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 PET imaging apparatus 1 (as many as the number of detector units 2 to be mounted). Yes. 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. Alternatively, 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.

  When the detector unit 2 is attached to the PET imaging apparatus 1, the lid 11a is removed to expose the unit support member 2A, and the detector unit 2 is inserted from there until the flange portion comes into contact. ing. By inserting and attaching, the connector of the PET imaging device 1 and the detector unit 2 is connected, and the signal and power of the PET imaging device 1 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. 11, 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). Then, twelve bonded substrates 20 are accommodated in one housing 30 (that is, one detector unit 2). Therefore, a voltage is supplied to 128 × 12 = 1536 detectors 21 from the high-voltage power supply device PS.

Incidentally, in the past, a power supply with a voltage of 300 V with very little fluctuation was supplied from a remote precision power supply, but (1) as the distance from the precision power supply becomes longer, a wider range of high-voltage wiring is provided. (2) Since the voltage fluctuates due to temperature fluctuations of the detector 21, supply a precise voltage from a precision power supply device. However, the target detector 21 has a problem that the voltage is not as accurate as expected.
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), and 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 to 15 V) DC by the power supply wiring via the power supply connector 42 and the connector 38 provided on the top plate 30a. Connected to power. The high voltage side terminal of the high-voltage power supply device PS is connected to 12 connectors C3 provided on the top plate 30a via connectors 43 provided on the top plate 30a, respectively. The electrodes are connected to the electrodes C of the detectors 21 provided on the substrate body 20a via power supply wiring (not shown) in 20b, a connector C1 and power supply wiring (not shown) in the substrate body 20a. 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, a high-voltage power supply PS having an accuracy corresponding to the voltage fluctuation due to temperature is provided instead of the precision power supply. 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. As a result, it is possible to prevent noise from being placed on the signal wiring from the power supply system.

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

Incidentally, the voltage supplied to the high-voltage power supply device PS is boosted to 300 V by a DC-DC converter (not shown) in the high-voltage power supply device PS, and after being boosted, passes through the top plate 30a of the housing 30 and is combined with the combined substrate 20. , ASIC substrate 20B → detector substrate 20A → each detector 21 is supplied. 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 20 includes a voltage supply wiring that supplies the voltage supplied from the high-voltage power supply device PS via the substrate connector C <b> 2 to each detector 21.
B. (X-ray CT imaging device)
The X-ray CT imaging device 4 is arranged behind the PET imaging device 1 as shown in FIG. 1, and as shown in FIG. 14, the radiation detector 40, the X-ray source circumferential direction moving device 41, and the drive It has the apparatus control apparatus 42, the X-ray source control apparatus 43, and the casing 40a (refer FIG. 15). The X-ray source circumferential direction moving device 41 includes a disk-type holding unit 44, an X-ray source 45, and an X-ray source device holding unit 45a. The X-ray source device holding unit 45 a is attached to the outer surface of the disk type holding unit 44 at one end of the disk type holding unit 44. The X-ray source 45 is attached to the other end of the X-ray source device holding unit 45a.

  The radiation detector 40 is disposed at a position where X-rays transmitted through the subject H from the X-ray source 45 can be detected. A plurality of (about 100) radiation detectors 40 are arranged from the disk-shaped holding unit 44 via the detector holding unit 46, and rotate around the subject H in conjunction with the X-ray source circumferential direction moving device 41. It is configured as follows. In addition, a collimator 47 is attached to the radiation detector 40 so that only X-rays generated from the X-ray source 45 are incident on the radiation detector 40. In this embodiment, a scintillator detector is used as the radiation detector 40.

  The X-ray source 45 has a known X-ray tube (not shown). The X-ray tube includes an anode, a cathode, a cathode current source, and a voltage source for applying a voltage between the anode and the cathode in an outer cylinder. The cathode is a tungsten filament. Electrons are emitted from the cathode by passing a current from the current source to the cathode. The electrons are accelerated by a voltage (140 kV) applied between the cathode and the anode from the voltage source, and collide with the target anode (W, Mo, etc.). X-rays of 140 keV or less are generated by collision of electrons with the anode and are emitted from the X-ray source 45.

  The X-ray source control device 43 controls the X-ray emission time from the X-ray source 45. That is, the X-ray source control device 43 outputs an X-ray generation signal during the X-ray CT examination, and is a switch provided between the anode (or cathode) of the X-ray tube in the X-ray source 45 and the power source. (Hereinafter referred to as X-ray source switch not shown) is closed, when the first set time elapses, an X-ray stop signal is output to open the X-ray source switch, and when the second set time elapses The control of closing the X-ray source switch is repeated. A voltage is applied between the anode and the cathode during the first set time, and no voltage is applied during the second set time. By this control, X-rays are emitted from the X-ray tube in pulses.

  When starting the X-ray CT examination, the drive device control device 42 outputs a drive start signal to close a switch (not shown) connected to a power source (hereinafter referred to as a first motor switch). When the current is supplied, the first motor rotates, and the rotational force is transmitted to the pinion via the power transmission mechanism, so that the pinion rotates. Due to the rotation of the pinion, the disk-shaped holder 44, that is, the X-ray source 45 moves around the subject H at a preset speed. At the end of the X-ray CT examination, the drive device controller 42 outputs a drive stop signal to open the first motor switch. As a result, the movement of the X-ray source 45 in the circumferential direction is stopped. Since the radiation detector 40 is fixed to the disk-shaped holder 34 via the detector holder 46, it rotates together with the X-ray source 45. Therefore, when X-rays are irradiated from the X-ray source 45, X-rays transmitted through the subject H can be measured by the radiation detector 40.

The drive start signal output from the drive device controller 42 when starting the X-ray CT examination is input to the X-ray source controller 43. The X-ray source control device 43 outputs an X-ray generation signal based on the input of the drive start signal. Thereafter, the X-ray stop signal and the X-ray generation signal are repeatedly output. As described above, the X-ray stop signal and the X-ray generation signal are repeatedly output, so that the X-ray source 45 emits X-rays for a set time (for example, 1 μsec), and then stops emitting X-rays. . The emission and stop of the X-ray are repeated during the movement period of the X-ray source 45 in the circumferential direction. The X-rays emitted from the X-ray source 45 are irradiated to the subject H in a fan beam shape. This X-ray is detected by the radiation detector 40 that is transmitted through the subject H and is at a position 180 degrees opposite to the X-ray source 45 and rotating simultaneously with the X-ray source. These radiation detectors 40 output X-ray detection signals (hereinafter referred to as X-ray detection signals).
In the above description, the X-ray pulse irradiation method has been described, but the present invention is not limited to this. X-rays are continuously irradiated, and the charge accumulation time may be controlled and measured on the measurement side.

(Data processing device)
As shown in FIG. 16, the data processing device 12 has a PET data processing unit 12 a that processes data from the PET imaging device 1 and an X-ray CT data processing unit 12 b that processes data from the X-ray CT imaging device 4. The PET / CT control unit 12c is included.
The PET data processing unit 12a includes a storage device (not shown), a simultaneous measurement device 12A, and a tomographic image information creation device 12B. The PET data processing unit 12a captures packet data including the detected peak value of γ-rays, detection time data, and detector (channel) ID. The simultaneous measurement device 12A performs simultaneous measurement based on this packet data, in particular, the detection time data and the detector ID, specifies the detection position of the 511 KeV γ-ray, and stores it in the storage device. The tomographic image information creation device 12B creates a functional image based on the identified position and outputs it. Here, as the PET image reconstruction method, in the two-dimensional case, for example, the filtered back projection method may be used. When three-dimensional imaging is performed, for example, the Fourier rebinning method discussed in 1997, IEEE Transactions on Medical Imaging, Vol. 16, page 145 is used. Perform image reconstruction. Thereby, a PET image is obtained. For PET images, generation density information of gamma ray pairs is obtained.

  The X-ray CT data processing unit 12b includes an amplifier circuit 12C and a sample hold circuit 12D, and receives an X-ray detection signal from the radiation detector 40 and converts the intensity of the X-ray detection signal into data. . Since the incident rate of X-rays emitted from the X-ray source 45 is overwhelmingly higher than that of the γ-rays, the X-ray CT data processing unit 12b generally has a so-called current mode (integration mode) measurement circuit. Consists of. The X-ray detection signal (current signal) from the radiation detector 40 is accumulated by the integrating amplifier circuit 12C, and the sample value of the signal is held by the sample hold circuit 12D. By repeating the above operation with a reset signal at a constant cycle (a maximum of several tens of milliseconds), the X-ray intensity for each fixed time is converted into data by the sample hold circuit 12D. As an image reconstruction method for X-ray CT data, for example, as described above, it is described on page 21 of NS-21 volume of IEEE Transactions on Nuclear Science. Performed using the Filtered Back Projection Method. The obtained image is a CT value in a rectangular parallelepiped or a cubic voxel obtained by dividing the body in the x, y, and z directions at equal intervals.

  The PET / CT control unit 12c is composed of a computer, a workstation, or the like, and creates a timing chart in the PET inspection and CT inspection therein, and based on the timing chart, the bed 14, the PET imaging apparatus 1, and the X-ray The CT imaging device 4, the PET data processing unit 12a, and the X-ray CT data processing unit 12b are instructed to perform a desired operation, and a tomographic image (PET image) is generated using the γ-ray imaging data from the PET data processing unit 12a. An X-ray CT image is reconstructed based on the X-ray imaging data from the line CT data processing unit 12b. Here, in obtaining the X-ray CT image, the attenuation coefficient of X-rays is used based on the X-ray imaging data, and the line attenuation coefficient in the subject H between the X-ray source 45 and the radiation detector 40 is calculated. Ask. Using this linear attenuation coefficient, the source weak coefficient of each voxel is obtained by a method such as a filtered back projection method. And the CT value in each voxel is obtained using the value of the linear attenuation coefficient of each voxel. X-ray CT image data is obtained using these CT values. Further, the amount of absorption with 511 keV γ rays is calculated from the CT value, whereby the in-vivo absorption in the subject H is corrected, and an accurate PET tomographic image is reconstructed. Both reconstructed tomographic images are displayed by the display device 13. Accordingly, it is not necessary to provide an absorption correction radiation source on the PET imaging apparatus 1 side.

  Here, the synthesis of the PET image and the X-ray CT image can be performed easily and accurately by matching the position of the central axis of the hole 50 (see FIG. 2) in both image data. An image may be synthesized with sinogram data and frequency space. Since the synthesized tomographic image displayed on the display device 13 includes the X-ray CT image, the position of the affected part in the PET image in the body of the subject H can be easily confirmed. That is, since the X-ray CT image includes images of the internal organs and bones, the position where the affected part (for example, an affected part of cancer) exists can be specified by the positional relationship with the internal organs and bones.

  In the present embodiment, the PET imaging apparatus 1 and the X-ray CT imaging apparatus 4 are arranged along the longitudinal direction of the bed 14 (back and forth in the body axis direction of the subject H). 1 and the X-ray CT image picked up by the X-ray CT image pickup device 4 can be taken independently, and there is no possibility that necessary data is lost due to interference or the like.

(Operation of radiation inspection equipment)
Next, the operation of the radiation inspection apparatus having the above configuration will be described.
Before performing a radiological examination, first, a radiopharmaceutical for PET is first administered to the subject H by a method such as injection so that the in-vivo radioactivity becomes 370 MBq. The radiopharmaceutical is selected according to the examination purpose (for example, grasping the location of cancer or examining an aneurysm of the heart). The subject H waits until the radiopharmaceutical is collected in a state where imaging is possible. The radiopharmaceutical collects in the affected area of the subject H as the predetermined time elapses. After the predetermined time has passed, the subject H is laid on the bed 14 (see FIG. 2). Depending on the type of examination, the radiopharmaceutical may be administered to the subject H laid on the bed 14.

  Examiners (medical radiographers and doctors) who carry out PET examinations and CT 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, the number of slices, slice intervals, CT Scan timing, absorbed dose, etc.) are input to the PET / CT controller 12c. This can be done by displaying an information input screen as shown in FIG. 17 on the display device 13 and inputting necessary data using a keyboard, mouse, or the like. As shown in FIG. 17, if a combo box, a radio button, or the like is arranged on the screen, input can be performed easily. In the PET / CT control, a PET inspection / CT inspection sequence (abbreviated as “inspection sequence” as appropriate) is created based on the input information. Incidentally, when the “display” button on the information input screen in FIG. 17 is clicked, an inspection sequence is created in the PET / CT control unit 12c and displayed on the display device 13, and when the “inspection start” button is clicked, the inspection is started. The In the PET / CT control unit 12c, the following parameters are all programmed in a series of inspection sequences, and are executed at a timing based on the number of clocks with respect to a reference time for starting an inspection.

(1) Operation and stop of PET imaging device 1 and X-ray CT imaging device 4 (2) Circulation of X-ray source 45, movement in body axis direction, irradiation amount (tube current, tube voltage)
(3) Operation and stop of PET data processing unit 12a and X-ray CT data processing unit 12b (4) Permit and prohibit transmission of γ-ray imaging data and X-ray imaging data (5) Movement control of bed 14

  Note that a sequence (including movement control of the bed 14) is performed so that imaging between the PET imaging apparatus 1 and the X-ray CT imaging apparatus 4 can be switched.

  Further, since the irradiation area of X-rays expands in the body axis direction, the radiation detector 40 is switched as an aggregate of up to about 10 detectors in the body axis direction. Supplementally, assuming that one radiation detector 40 has a size of 5 mm square, when 10 of these are arranged in series, it becomes 50 mm. This value of 50 mm matches the extent of X-rays (fan-shaped with a 5 ° spread in the body axis direction) in which the irradiation area is expanded in the body axis direction.

  First, the subject H laid on the bed 14 is moved to a predetermined position, and the PET imaging apparatus 1 is operated.

  In response to an instruction from the PET / CT control unit 12c, the PET data processing unit 12a operates to start the PET examination. The γ rays emitted from the body of the subject H are detected by the detector 21 and transmitted to the PET data processing unit 12c as γ ray imaging data. The PET data processing unit 12a generates γ-ray imaging data as described above and transmits it to the PET / CT control unit 12c. According to the sequence, PET inspection is performed for a while in this state, and when predetermined imaging is completed, the PET imaging apparatus 1 is stopped.

Prior to the CT examination, the X-ray source 45 is driven to generate X-rays. The X-ray generation intensity is stabilized at a predetermined value (tube current, tube voltage), and a standby state is set. And bed 14 is X-ray CT
It moves to the predetermined position of the imaging device 4 (refer FIG. 14).

  The radiation detector 40 included in the region irradiated with X-rays is connected to the X-ray CT data processing unit 12b, and the X-ray CT data processing unit 12b is operated to acquire X-ray imaging data. Here, since the X-ray source 45 irradiates the X-ray with a spread of about 5 ° in the body axis direction and about 60 ° in the circumferential direction, the X-ray source 45 is applied to the body axis direction irradiation region and the circumferential direction irradiation region (both not shown). A plurality of included radiation detectors 40 are connected to the X-ray CT data processing unit 12a (collecting and processing 5 ° spread in the body axis direction). The X-ray source 45 moves around the X-ray source 45 by rotating the X-ray source 45 in the circumferential direction by the X-ray source circumferential direction moving device 41 to perform CT examination.

  Since the inspection sequence is set so that the rotation timing of the X-ray source 45 and the switching timing of each radiation detector 40 group are synchronized, the switching command is issued without detecting the position of the X-ray source 45. That is, the circuit switching timing of each radiation detector 40 can be set in the sequence program since the rotation start time and rotation speed (angular velocity) of the X-ray source 45 are known.

  Thereafter, the irradiation by the X-ray source 45 is stopped, the X-ray CT imaging apparatus 4 is stopped, the bed 14 is moved to the initial position, and the PET / CT examination is ended.

The PET / CT control unit 12c uses the γ-ray imaging data received from the PET data processing unit 12a to form a PET image, and uses the X-ray imaging data obtained from the X-ray CT data processing unit 12b to generate an X-ray CT image. Reconfigure. Then, the reconstructed X-ray CT image is displayed on the display device 13.
Below, the effect in this embodiment is demonstrated.
(1) In this embodiment, energy resolution is improved by using a semiconductor radiation detector, and scattered rays can be removed. In particular, at the time of 3D imaging, an increase in scattered radiation can be suppressed, the image quality of a PET image can be improved, and a quantitative test can be performed. FIG. 22 shows a simulation result simulating the time of 3D imaging. FIG. 22 is a plot of the relationship between the energy threshold value and the scattering rate. The energy resolution of the semiconductor radiation detector is several percent, and the energy threshold can be increased to about 475 KeV. From FIG. 22, when the energy threshold is 475 KeV, the scattering rate can be suppressed to 20% or less, which is the same as in 2D imaging, and high-quantitative inspection is possible.
(2) In this embodiment, the position resolution is improved by using a semiconductor radiation detector. In the scintillator, since the signals of several tens of scintillators are amplified by one photo and the detected scintillator position is calculated using the center of gravity calculation or the like, the position resolution is deteriorated. In addition, there is a limit to miniaturization of scintillator because of using photomaru. On the other hand, in the PET imaging apparatus using the semiconductor radiation detector of the present embodiment, an amplification circuit is formed for each semiconductor radiation detector, and the position resolution does not deteriorate. Furthermore, since the signal processing circuit is formed using ASIC or the like, the semiconductor radiation detector can be easily miniaturized and the position resolution can be further improved.
(3) In the present embodiment, the effect (1) enables inspection with high quantitativeness even in 3D imaging, so that 2D imaging is unnecessary and a 3D dedicated PET apparatus is possible. Therefore, it is not necessary to put a septa inside the detector, and the apparatus can be miniaturized.
(4) In this embodiment, since the semiconductor radiation detector is used and the ASIC is used for the signal processing, the size around the semiconductor radiation detector can be reduced as compared with the photomultiplier used in the scintillator. . Moreover, since the semiconductor radiation detector and the signal processing circuit are arranged in order on the integrated substrate, further miniaturization is realized.
(5) In this embodiment, the X-ray CT imaging device and the PET imaging device are arranged in series and the absorption correction is performed using the CT value. There is no need to rotate the detector inside, and the device can be miniaturized.
(6) In this embodiment, the X-ray CT imaging device and the PET imaging device are arranged in series according to the effects (1) to (5), and the entire device is miniaturized, so that the subject is not intimidated. In addition to being configured, it provides high-resolution images with high quantitativeness, enabling high-precision inspection.

[Embodiment 2]
The radiation inspection apparatus which is other embodiment is demonstrated. As shown in FIG. 18, the radiation inspection apparatus according to the present embodiment is different from the first embodiment in that a semiconductor radiation detector 21 is used as a radiation detector in the X-ray CT imaging apparatus 4. Specifically, a plurality of coupled substrates 60 as shown in FIG. 19 are provided in the slice direction via the detector holding unit 46 (four slices in this embodiment), and the X-ray source circumferential direction moving device 41 is provided. In conjunction with the rotation of the subject H. The configuration of the coupling substrate 60 includes the detector 21, the resistor 23, the analog ASIC 24A, the ADC 25, and the digital ASIC 26A as in the coupling substrate 20 in the first embodiment, and the number of the detector 21, the analog ASIC 24A, and the ADC 25 is small. It is the same except that it is configured. That is, 16 detectors 21 are provided in one row and 32 on both sides, and accordingly, one set of analog ASIC 24A and ADC 25 is provided.

  Imaging in the X-ray CT imaging apparatus 4 having such a detector 21 is performed by moving the bed 14 and moving the subject H between the X-ray source 45 and the detector 21. At that time, the X-ray source 45 and the detector 21 rotate around the subject H by rotating the disk-shaped holding unit 44. Then, the X-rays emitted from the X-ray source 45 are incident on the corresponding detector 21 with a desired spread, and the detector 21 outputs an X-ray detection signal. This X-ray detection signal is processed by an analog ASIC 24A and a digital ASIC 26A which will be described later.

  Here, the ASIC substrate 60B constituting the coupling substrate (unit substrate) 60 will be described with reference to FIGS. The ASIC board 60B connected to the detector board 60A by the connector C1 includes a resistor 23, one analog ASIC 24A, and one digital ASIC 26A provided for each detector 21. The analog ASIC 24A includes 32 sets of analog signal processing circuits (analog signal processing devices) 33A. The analog signal processing circuit 33A is provided for each detector 21. Here, a charge amplifier 24b, a polarity amplifier 24c, a band pass filter 24d, and a peak hold circuit 24e are connected in this order. One analog ASIC 24A is an LSI formed of 32 sets of analog signal processing circuits 33A. The X-ray detection signal output from the detector 21 and passed through the resistor 23 is input to the peak hold circuit 24e via the charge amplifier 24b, the polarity amplifier 24c, and the band pass filter 24d. The peak hold circuit 24e samples and holds the intensity of the X-ray detection signal.

  The digital ASIC 26A includes a packet data generation device 34 and a data transfer circuit 37, which are LSIs. The digital ASIC 26A receives a clock signal from a clock generator (crystal oscillator) of 64 MHz (not shown) and operates in synchronization. The ADC control circuit 36 receives the drive start signal output from the drive device control device 42 via the unit integrated LSI 31 and starts measurement. The ADC control circuit 36 has a counter inside and manages the measurement time. That is, the peak value information is read while operating the ADC 25 at predetermined time intervals and switching each channel. At the same time, the analog ASIC 24A is controlled to switch the peak value output channel, and the charge amplifier is reset. The ADC control circuit 36 generates packet data that is digital information including time information, detector ID, and peak value information. The packet data output from each ADC control circuit 36 is periodically transmitted from the data transfer circuit 37 to the unit integrated FPGA 31 of the detector unit 2A. The unit integrated FPGA 31 collectively outputs the packet data input from the data transfer circuit 37 of each coupling substrate 53 to the information transmission wiring connected to the connector 38.

  Each packet data output from the unit integrated FPGA 31 is transmitted to the data processing device 12A. The data processing device 12A uses the detector ID information and time information to determine the position coordinates of the detector 21 corresponding to the detector ID information at the time of X-ray detection. The position coordinates represent the position of the detector 21 that is turning along the turning track of the disk-type holding unit 44 when the X-ray is detected. Since the angular velocity at which the disc-shaped holding unit 44 turns is constant, the position (positional coordinates) of the detector 21 at the time when the X-ray is detected can be obtained using the time determined by the time determination circuit 35. The data processing apparatus 12A creates tomographic image information of the subject H based on the X-ray CT data detected at the position of each detector 21 obtained using the packet data. This tomographic image information is displayed on the display device 13. Information such as the packet information, the position information of the detector 21, and the tomographic image information is stored in the storage device of the data processing device 12A.

This embodiment has the effects described below in addition to the effects (1) to (6).
(7) In this embodiment, a semiconductor radiation detector is used as the detector of the X-ray CT imaging apparatus, an ASIC or the like is used for the signal processing circuit, and the X-ray CT is arranged in order on the integrated substrate. The imaging device is downsized. Therefore, the X-ray CT imaging apparatus and the PET imaging apparatus are arranged in series, and the entire apparatus is further miniaturized so as not to give a sense of intimidation to the subject.

  In each of the above embodiments, the mounting (housing) of the detector unit 2 to the PET imaging device 1 and the X-ray CT imaging device 4 is not limited to the unit support member 2A described above. Any mounting / storing means / method may be used.

It is a perspective view which shows the structure of the radiation inspection apparatus concerning this embodiment. It is the figure which showed typically the PET imaging device and X-ray CT imaging device of FIG. It is the figure which showed typically the cross section of the circumferential direction of the PET imaging device of FIG. It is the figure which showed 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. It is the graph which showed typically the relationship between the thickness t of the semiconductor material of a semiconductor radiation detector, and a crest value (maximum value). It is the figure which showed typically the structure of the semiconductor radiation detector which has the 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. It is the block diagram which showed schematic structure of digital ASIC, and the connection relation of analog ASIC and digital ASIC. It is the see-through | perspective perspective view quoted in order to demonstrate the structure of the detector unit which accommodated two or more semiconductor radiation detectors. It is the side view of the place which removed the side plate of the detector unit of FIG. (A) is the partially broken perspective view which showed the mode at the time of mounting | wearing a PET imaging device with a detector unit, (b) is sectional drawing of a center part. It is the figure which showed the X-ray CT imaging device typically. It is the figure which showed typically the cross section of the circumferential direction of the X-ray CT imaging device of FIG. It is the block diagram which showed the data processor typically. It is the figure which showed an example of the operation screen displayed on a display apparatus. It is the figure which showed typically the structure of the X-ray CT imaging device as a radiation inspection apparatus concerning other embodiment. (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). 3 is a block diagram schematically showing an analog detection circuit. FIG. It is the block diagram which showed schematic structure of digital ASIC, and the connection relation of analog ASIC and digital ASIC. It is a simulation result at the time of 3D imaging, and is a graph showing a relationship between an energy threshold value and a scattering rate.

Explanation of symbols

1 PET imaging device (first imaging device)
4 X-ray CT imaging device (second imaging device)
12 data processing device 13 display device 14 bed 2 detector unit 20 coupling substrate 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 PS High voltage power supply

Claims (24)

A bed on which a subject is placed; and first and second imaging devices arranged along the longitudinal direction of the bed,
The first imaging device includes a plurality of semiconductor radiation detectors that detect γ-rays emitted from the subject, surrounding the bed,
The second imaging device includes an X-ray source that irradiates the subject with X-rays, and a radiation detector that detects the X-rays transmitted through the subject and transmitted from the X-ray source,
The bed is shared by the first imaging device and the second imaging device.   The radiation inspection apparatus according to claim 1, wherein the radiation detector in the second imaging apparatus is a semiconductor radiation detector.   The first imaging device is formed to be smaller in shape than the second imaging device, and is disposed on the front side of the second imaging device in the longitudinal direction of the bed. The radiation inspection apparatus according to claim 1 or 2.   The radiation inspection apparatus according to claim 2 or 3, wherein the second imaging apparatus includes a plurality of semiconductor radiation detectors arranged in a longitudinal direction of the bed. Tomographic image information is created using an integrated circuit that processes a radiation detection signal output from each of the plurality of semiconductor radiation detectors and second information obtained based on the first information output from the integrated circuit. A tomogram information creation device,
5. The radiological examination according to claim 1, wherein the first imaging device includes a plurality of unit substrates including the plurality of semiconductor radiation detectors and the integrated circuit. 6. apparatus.   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 The radiation inspection apparatus according to claim 5, comprising an integrated circuit.   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. The radiation inspection apparatus according to claim 6. The unit substrate includes a first substrate and a second substrate,
The first substrate has at least the semiconductor radiation detector;
The radiation inspection apparatus according to claim 5, wherein the second substrate includes at least the integrated circuit.   The radiation inspection apparatus according to claim 1, wherein the first imaging apparatus is a positron emission CT apparatus.   The radiation inspection apparatus according to claim 9, wherein the first imaging apparatus is a 3D-dedicated positron emission CT apparatus that does not hold a scepter for 2D imaging and a radiation source for absorption correction.   The semiconductor radiation detector has a semiconductor region that generates charges by interacting with radiation, and an anode electrode and a cathode electrode face each other across the semiconductor region, and the anode electrode and the cathode electrode The radiation inspection apparatus according to claim 1, wherein a distance between the electrodes or a thickness of the semiconductor region sandwiched between the anode electrode and the cathode electrode is 0.2 to 2 mm.   The radiation inspection apparatus according to claim 1, wherein the distance between the electrodes or the thickness of the semiconductor region is 0.5 to 1.5 mm.   The radiation inspection apparatus according to claim 8, wherein the first substrate and the second substrate are detachably coupled to each other.   The radiation inspection apparatus according to claim 13, wherein the first substrate and the second substrate are joined with their end portions overlapped with each other.   The radiation inspection apparatus according to claim 5, wherein the semiconductor radiation detectors are disposed on both surfaces of the unit substrate. The first imaging device has a rotating body and a bed on which a 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,
The collimator which has a plurality of radiation passages which oppose the semiconductor radiation detector, and is arranged on the bed side rather than the semiconductor radiation detector is installed in the support member. 5. The radiological examination apparatus according to 5.   The radiological examination apparatus according to claim 6, wherein the analog integrated circuit has a function of performing signal amplification, and the digital integrated circuit has a function of generating the time information. 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. And the digital integrated circuit outputs the time information and the identification information based on the output of the signal processing device,
A simultaneous measurement device that performs simultaneous measurement based on the time information; and the tomogram information creation device that creates the tomogram information using the identification information and information obtained by the simultaneous measurement device. The radiation inspection apparatus according to claim 6, wherein the radiation inspection apparatus is characterized. The analog integrated circuit inputs the radiation detection signal from an upstream side of the amplifier, and a slow system including a peak value output device that inputs the amplifier and an output of the amplifier and outputs a 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 of the semiconductor radiation detectors and generates time information based on the radiation detection signal Including equipment and
A simultaneous measurement device that performs simultaneous measurement based on the time information; and the tomogram information creation device that creates the tomogram information using the identification information and information obtained by the simultaneous measurement device. The radiation inspection apparatus according to claim 6, wherein the radiation inspection apparatus is characterized. When the digital integrated circuit further inputs from the time information generation device, AD that specifies one of the position information and the identification information of the semiconductor radiation detector connected to the time information generation device A conversion control device, and 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 the plurality of signal processing devices included in the analog integrated circuit. The radiological examination apparatus according to claim 19, wherein a high value is converted into peak value information that is digital information and output to the information integration device. 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,
The radiation inspection apparatus according to claim 1, wherein a plurality of unit substrates are detachably accommodated in the frame.   The radiation inspection apparatus according to claim 21, wherein the frame body includes a plurality of guide members for guiding the unit substrate. A power supply device including a voltage booster for boosting a voltage in the frame;
23. The radiation inspection apparatus according to claim 21, further comprising a voltage supply wiring for supplying a voltage from the power supply device to each of the semiconductor radiation detectors of the unit substrate.   The radiation inspection apparatus according to any one of claims 21 to 23, wherein the frame body has a light-shielding configuration.

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