JP4737201B2 - Radiation inspection equipment - Google Patents

Description

  The present invention relates to a radiation inspection apparatus, and more particularly to a radiation two-dimensional imaging apparatus, X-ray computed tomography (hereinafter referred to as X-ray CT), positron emission CT (positron emission CT). Computed Tomography (Positron Emission Computed Tomography) (hereinafter referred to as PET) and Single Photon Emission Computed Tomography (hereinafter referred to as SPET) The present invention relates to a radiological examination apparatus suitable for adapting to the above.

  Representative examples of a radiological examination apparatus that is a technique for imaging non-invasively the function and form of a body of a subject to be examined are a radiation two-dimensional imaging apparatus, X-ray CT, PET, and SPECT.

In PET examination, a radiopharmaceutical (referred to as a PET drug) containing a positron emitting nuclide ( ¹⁵ O, ¹³ N, ¹¹ C, ¹⁸ F, etc.), which is a radionuclide, is administered to the examinee. This is a test to check whether a lot is consumed at the site. The PET examination is an action of detecting γ-rays emitted from the body of the examinee due to the PET drug with a radiation detector. Specifically, a positron emitted from a radionuclide contained in a PET drug is combined with an electron in a nearby cell (cancer cell) and disappears. At that time, a pair of γ-rays having energy of 511 keV ( (Referred to as γ rays). Those γ rays are emitted in directions almost opposite to each other (180 ° ± 0.6 °). If this pair of γ-rays is detected by a radiation detector, it can be determined between which two radiation detectors the positron has been emitted. By detecting these many pairs of γ rays, it is possible to find a place where a lot of PET drug is consumed. For example, when a PET drug manufactured by combining a positron emitting nuclide and a sugar is used, it is possible to find a cancer lesion with intense sugar metabolism. An example of a radiation inspection apparatus used for PET is described in Patent Document 1. Note that the obtained data is converted into the data of each voxel by the filtered back projection method described in Non-Patent Document 1. The half-life of positron emitting nuclides ( ¹⁵ O, ¹³ N, ¹¹ C, ¹⁸ F, etc.) used for PET inspection is 2 to 110 minutes.

SPECT is a radiopharmaceutical (SPECT drug) containing a single photon emitting nuclide ( ⁹⁹ Tc, ⁶⁷ Ga, ²⁰¹ Tl, etc.) which is a radionuclide, and a substance (for example, sugar) having a property of accumulating in a specific tumor or a specific molecule And γ rays emitted from the radionuclide are detected with a radiation detector. The energy of γ rays emitted from a single photon emission nuclide often used at the time of inspection by SPECT is around several hundreds keV. In the case of SPECT, since a single gamma ray is emitted, the angle of the gamma ray incident on the radiation detector cannot be obtained. Therefore, angle information is obtained by detecting only γ-rays incident from a specific angle by a radiation detector using a collimator. SPECT is an inspection method for identifying a place where a large amount of SPECT drug is consumed by detecting γ rays generated in the body due to the SPECT drug. An example of a radiation inspection apparatus used for SPECT is described in Patent Document 2. Also in the case of SPECT, the obtained data is converted into data of each voxel by a method such as filtered back projection. Note that a transmission image may be taken even in SPECT. ⁹⁹ Tc, ⁶⁷ Ga, ²⁰¹ Tl used for SPECT is 6 hours to 3 days longer than the half-life of the radionuclide for PET.

Japanese Patent Laid-Open No. 7-20245 Japanese Patent Laid-Open No. 9-5441 IEEE Transaction on Nuclear Science NS-21, pp. 228-229

  For example, further improvement in diagnostic accuracy such as the position and size of an affected area such as a malignant tumor is desired, and improvement in the accuracy of an image including the affected area created by such a radiological examination apparatus is required. It is also an important issue to be able to replace a failed radiation detector in a short time.

An object of the present invention is to provide a radiological imaging apparatus that can be improve the accuracy of the image to be created.

A feature of the present invention that achieves the above-described object is that a detector support member disposed around a bed that supports a subject, and a plurality of detector units disposed in the longitudinal direction of the bed and around the bed. The detector unit includes a plurality of radiation detectors for detecting radiation, and the radiation detector includes a plurality of semiconductor members arranged in parallel and detection signals arranged alternately. An output electrode and a common potential electrode, the detection signal output electrode and the common potential electrode are disposed on the semiconductor member disposed between the electrodes, and the radiation detector is arranged in a radial direction of the detector support member. The radiation inspection apparatus includes a plurality of γ-ray detection signal processing devices that obtain positional information of the radiation detectors in the radial direction of the radiation detectors that are arranged in a plurality and that output the γ-ray detection signals.

  According to the present invention, the accuracy of an image to be created can be improved, and a failed radiation detector can be easily replaced.

Example 1
A radiation inspection apparatus according to a preferred embodiment of the present invention will be described below with reference to FIGS. The radiation inspection apparatus 1 of the present embodiment is used for PET inspection. The radiation examination apparatus 1 includes an imaging device 2, a signal processing device 40, a tomographic image creation device 35, an examinee holding device 18, a drive device control device 21, and an X-ray source control device 22. The examinee holding device 18 is installed at the upper end of the bed support 19 so that the bed 20 can move in the longitudinal direction of the bed 20.

  The imaging device 2 includes a casing 3, a large number of detector units 4, an annular detector support member 8, and an X-ray source circumferential direction moving device 13. As shown in FIGS. 3 and 4, the detector support member 8 includes an annular detector support portion 23 attached to the support member 39 and a cover member 24. The cover member 24 is attached to the detector support portion 23 so as to cover the signal discrimination unit storage space 44 formed in the detector support portion 23.

  The X-ray source circumferential direction moving device 13 includes a guide rail 12 and an X-ray source device 14. The annular guide rail 12 is attached to the side of the patient holding device 18 on the side of the detector support member 8, specifically on the side of the detector support 23 so as to surround the hole 41 into which the bed 20 is inserted. It is done. The X-ray source device 14 includes an X-ray source driving device 15, an extendable arm 16 and an X-ray source 17. The X-ray source driving device 15 is movably attached to the guide rail 12. Although not shown, the X-ray source driving device 15 includes a pinion that meshes with the rack of the guide rail 12 and includes a motor that rotates the pinion via a speed reduction mechanism. The telescopic arm 16 is attached to a casing (not shown) of the X-ray source driving device 15 and can extend and contract in the horizontal direction. The X-ray source 17 is attached to the distal end portion of the telescopic arm 16.

  Although not shown, the X-ray source 17 has a known X-ray tube. 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 filament by passing a current from the current source to the cathode. The electrons are accelerated by a voltage (several hundred kV) applied between the cathode and the anode from the voltage source, and collide with the target anode (W, Mo, etc.). X-rays of 80 keV are generated by the collision of the electrons with the anode. This X-ray is emitted from the X-ray source 17.

  The tomographic image creation device 35 includes a computer 36 and a storage device 37. The computer 36 is connected to the coincidence counting device 34, and the storage device 37 is connected to the computer 36. The computer 36 is a tomographic image creation unit. The display device 38 is connected to the computer 36.

  As shown in FIGS. 5 and 6, the detector unit 4 has a plurality of (for example, nine) radiation detectors 5 on one surface of the support substrate 6, and the connector portion 7 is provided on the support substrate 6. . Nine radiation detectors 5 are arranged on the support substrate 6 in three rows and three columns. The “circumferential direction” displayed in FIG. 4 is the circumferential direction of the detector support member 8, the “axial direction” is the axial direction of the detector support member 8, and the “radial direction” is the direction of the detector support member 8. Each means the radial direction (the same applies to FIGS. 10 and 12). The three radiation detectors 5 arranged in a row in the radial direction of the detector support member 8, that is, the cathode electrodes K 1, K 2, K 3 of the radiation detectors 5 A, 5 B, 5 C are connected to the ground wire 45. The ground wire 45 is connected to the connector terminal 7 </ b> D of the connector portion 7. The wiring 46 connected to the anode electrode A1 of the radiation detector 5A is connected to the connector terminal 7A of the connector portion 7. The wiring 47 connected to the anode electrode A2 of the radiation detector 5B is connected to the connector terminal 7B of the connector portion 7. The wiring 48 connected to the anode electrode A3 of the radiation detector 5C is connected to the connector terminal 7C of the connector portion 7. Similarly, the radiation detectors 5 included in the other two rows are also connected to other connector terminals provided in the connector portion 7. The ground wire 45 and the wirings 46, 47, 48 are all installed in the support substrate 6. The multiple detector units 4 are mounted and held on the detector support 23 by inserting connector terminals such as the connector terminals 7A provided on the detector units 4 into the connector 11 provided on the detector support 23. The A large number of detector units 4 surround the hole 41 and are arranged in the circumferential direction and the axial direction of the hole 41. These detector units 4 are detachably attached to the detector support portion 23.

  The casing 3 is attached to the detector support portion 23 so as to cover these detector units 4 (FIG. 4). Further, the casing 3 is formed to extend in the horizontal direction and form a hole (through hole) 41 into which the bed 20 is inserted during inspection (FIG. 1).

  By installing these detector units 4, a large number of radiation detectors (for example, a total of 10,000) 5 are arranged in the casing 3 inside the annular detector support member 8. The radiation detectors 5 are arranged in multiple layers (for example, three layers) in the radial direction of the detector support member 8 and in a plurality of rows in the axial direction of the detector support member 8. Of the radiation detectors 5 arranged in each detector unit 4, the three radiation detectors 5 (radiation detectors 5 </ b> A) that are farthest from the connector portion 7 are located closest to the axis of the hole 41. This is called the first-layer radiation detector. The three radiation detectors 5 (radiation detectors 5C) located closest to the connector portion 7 are located farthest from the axis of the hole 41 and are referred to as third-layer radiation detectors. The three radiation detectors 5 (radiation detectors 5B) located between the first layer and the third layer in the detector unit 4 are referred to as second layer radiation detectors.

  The radiation detection device 43 includes the many radiation detection units 4 described above. The radiation detection device 43 includes the radiation detectors 5 arranged in one to three layers in the radial direction of the detector support member 8 and arranged in a large number in the axial direction of the detector support member 8.

  Typical radiation detectors include semiconductor radiation detectors and scintillators. The scintillator is not suitable for stacking (for example, the above-described three layers) because it is necessary to arrange a photomultiplier tube or the like behind the crystal (BGO, NaI, etc.) that is the radiation detection unit. The semiconductor radiation detector is suitable for a stacked arrangement because a photomultiplier tube or the like is unnecessary. In this embodiment, the radiation detector 5 uses a semiconductor radiation detector, and a 5 mm cube which is a detection unit is made of cadmium tellurium (CdTe). The detector may be composed of gallium arsenide (GaAs) or cadmium tellurium zinc (CZT).

  The signal processing device 40 includes a signal discriminating device 27, a γ-ray discriminating device 32, and a coincidence counting device 34. The signal discriminating device 27 is provided for each radiation detector 5 in the first layer. In addition, a γ-ray discriminating device 32 is provided for each of the second-layer and third-layer radiation detectors 5. These three signal discriminating devices 27 and six γ-ray discriminating devices 32 are installed on one substrate 26. A signal discriminating unit 25 is constituted by three signal discriminating devices 27 and six gamma ray discriminating devices 32 installed on the substrate 26. The substrate 26 is attached to the unit support member 66. As shown in FIG. 4, each signal discrimination unit 25 provided for each detector unit 4 is attached to a unit support member 66 disposed in the signal discrimination unit storage space 44. The unit support member 66 is attached to the detector support unit 23. The signal discriminating unit 25 is held on the detector support member 8 by being installed on the unit support member 66. It is also possible to directly attach the substrate 26 to the detector support portion 23 as a support substrate without using the unit support member 66.

  As shown in FIG. 7, the signal discriminating device 27 includes a changeover switch 28, a γ-ray discriminating device 32, and an X-ray signal processing device 33. The changeover switch 28 has a movable terminal 29 and fixed terminals 30 and 31. The γ-ray discriminating device 32 is connected to the fixed terminal 30, and the X-ray signal processing device 33 is connected to the fixed terminal 31. The connector terminal 7 </ b> A connected to the first-layer radiation detector 5 </ b> A comes into contact with the connector terminal 11 </ b> A provided in the connector unit 11 by the connection between the connector unit 7 and the connector unit 11. The movable terminal 29 is connected to the connector terminal 11A by a wiring 49. The wiring 49 is installed on the unit support member 66. The negative terminal of the power supply 50 is connected to the wiring 46 through the resistor 51, and the positive terminal of the power supply 50 is connected to the radiation detector 5A. The γ-ray discriminating devices 32 in all the signal discriminating devices 27 are connected to the coincidence counting device 34 by wiring 52. Further, the X-ray signal processing devices 33 in all the signal discriminating devices 27 are connected to the computer 36 by wiring 53.

  Of the six γ-ray discriminating devices 32 other than the signal discriminating device 27 provided in the signal discriminating unit 25, the three γ-ray discriminating devices 32 are connected to the connector terminal 11 </ b> B (not shown) of the connector section 11 by the wiring 54. Connected to. The connector terminal 11B is in contact with the connector terminal 7B to which the second-layer radiation detector 5B is connected. The remaining three γ-ray discriminating devices 32 are connected to a connector terminal 11C (not shown) of the connector portion 11 by another wiring 54. The connector terminal 11B is in contact with the connector terminal 7C to which the third layer radiation detector 5C is connected. The six γ-ray discriminating devices 32 other than the signal discriminating device 27 are connected to the coincidence counting device 34 by wires 55, respectively. In FIG. 1, the signal discriminating unit 25 and the wirings 54 and 56 are displayed outside the detector support member 8, but this is because of the signal discriminating device 27 and the γ-ray discriminating device 32 provided in the signal discriminating unit 25. This is to make the connection state by wiring easier to understand. The signal discriminating unit 25 is actually installed in the detector support member 8 as shown in FIGS. 3 and 4, and the wirings 52, 53, 55 are drawn from the detector support member 8.

Before specifically describing the radiation inspection in this embodiment, the principle of radiation detection in this embodiment will be described. X-ray CT images (tomographic images including visceral and bone images of the subject obtained by X-ray CT) are obtained by scanning X-rays emitted from an X-ray source in a specific direction for a predetermined time. The operation (scanning) of irradiating the subject and detecting the X-rays transmitted through the body by the radiation detector is repeated, and is created based on the intensities of the X-rays detected by the plurality of radiation detectors. In order to obtain accurate X-ray CT image data, in the X-ray CT examination, γ-rays emitted from the inside of the subject due to the PET drug to the radiation detector detecting the X-rays Is preferably not incident. In one radiation detector, if the X-ray irradiation time on the subject is shortened corresponding to the incidence rate of γ-rays, the influence of γ-rays can be ignored. Shortening was attempted. In order to determine the irradiation time T of the X-ray, first, the incidence rate of γ rays to one radiation detector is considered. In the PET examination, the radioactivity in the body based on the PET drug administered to the subject in the PET examination is N (Bq), the passing rate of the generated γ-ray in the body is A, and the incidence rate obtained from the solid angle of one radiation detector is B. , Where the sensitivity of the detection element is C, the rate α (number / sec) of γ rays detected by one radiation detector is given by equation (1). In the equation (1), the coefficient “2” is a pair (two) of γ at the time of annihilation of one positron.
α = 2NABC (1)
It means that a line is emitted. The probability W that γ rays are detected by one detection element within the irradiation time T is given by equation (2). (2) W = 1−exp (−Tα) (2) so as to reduce the value of W in the equation
By determining the irradiation time T, the influence of γ rays incident on one radiation detector is negligible during the X-ray CT examination.

An example of the X-ray irradiation time T will be described below. A specific X-ray irradiation time T was determined based on the equations (1) and (2). The intensity of radiation in the body due to the PET drug administered to the subject in the PET examination is about 370 MBq at the maximum (N = 370 MBq), and the γ-ray passage rate A in the body is a radius of 15 cm. Assuming water, it is about 0.6 (A = 0.6). For example, considering a case where a radiation detector having a side of 5 mm is arranged in a ring shape with a radius of 50 cm, the incidence rate B obtained from the solid angle of one radiation detector is 8 × 10 ⁻⁶ (B = 8 × 10 ⁻⁶ ). It is. The detection sensitivity C of the radiation detector is about 0.6 (C = 0.6) at the maximum when a semiconductor radiation detector is used. From these values, the detection rate α of γ rays of one radiation detector is about 2000 (pieces / sec). If the X-ray irradiation time T is 1.5 μsec, for example, the probability W that one radiation detector detects γ rays during X-ray detection is 0.003, and these γ rays can be almost ignored. When the radioactivity administered to the body is 360 MBq or less, if the X-ray irradiation time is 1.5 μsec or less, W <0.003, that is, the detection probability of γ rays is 0.3% or less and can be ignored.

  An X-ray CT inspection and a PET inspection in the present embodiment using the imaging device 2B to which the above principle is applied will be specifically described.

  An X-ray CT inspection and a PET inspection in this embodiment will be described. By a method such as injection, a PET drug is previously administered to the examinee 42 as a subject so that the in-vivo radioactivity becomes 370 MBq. Thereafter, the examinee 42 waits for a predetermined time until the PET drug diffuses into the examinee's 42 and collects in the affected area (for example, an affected area of cancer) and becomes ready for imaging. The PET drug is selected according to the affected area to be examined. After the predetermined time has passed, the bed 20 on which the examinee 42 lies is inserted into the hole 41 of the imaging apparatus 2 together with the examinee 42. X-ray CT inspection and PET inspection are performed using the imaging apparatus 2. The examinee 42 to which the PET drug was administered was inserted into the hole 41 and a voltage was applied to each radiation detector 5 from the power supply 50, and then each radiation detector 5 was released from the examinee 42. Detect gamma rays. That is, the PET inspection is started. After the PET examination is started, the X-ray CT examination is started.

  An X-ray CT examination will be described. When starting the X-ray CT examination, the driving device control device 21 outputs a driving start signal and is connected to the motor of the X-ray source driving device 15 and connected to the power source (hereinafter referred to as a motor switch). Close. The rotational force of the motor is transmitted to the pinion via the speed reduction mechanism, and the X-ray source device 14, that is, the X-ray source 17 moves in the circumferential direction along the guide rail 12. The X-ray source 17 moves around the examinee 42 at a set speed while being inserted into the hole 41. At the end of the X-ray CT examination, the drive device controller 21 outputs a drive stop signal to open the motor switch. As a result, the movement of the X-ray source 17 in the circumferential direction is stopped. In this embodiment, all the radiation detectors 5 do not move in the circumferential direction and do not move in the axial direction of the hole 41. The drive device control device 21 and the X-ray source control device 22 are installed on the detector support member 8. A known technique that does not hinder the movement of the X-ray source device 14 is applied to the transmission of the control signal from the driving device control device 21 and the X-ray source control device 22 to the X-ray source device 14 that moves.

  The X-ray source control device 22 controls the emission time of X-rays from the X-ray source 17. That is, the radiation source control device 22 repeatedly outputs the X-ray generation signal and the X-ray stop signal. The first X-ray generation signal is output based on the drive start signal input to the X-ray source control device 22. The switch provided between the anode (or cathode) of the X-ray tube in the X-ray source 17 and the power source (hereinafter, not shown) is closed by the output of the X-ray generation signal, When the first set time elapses, the X-ray stop signal is output, the X-ray source switch is opened, and when the second set time elapses, the X-ray source switch is closed. 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. Under the control of the X-ray source control device 22, 80 keV X-rays are emitted from the X-ray tube in pulses. The irradiation time T, which is the first set time, is set to 1 μsec, for example, so that the detection probability of γ rays at the radiation detector 5 can be ignored. The second set time is a time T0 in which the X-ray source 17 moves between one radiation detector 5 and another radiation detector 5 adjacent to the radiation detector 5 in the circumferential direction. It is determined by the moving speed of the source 17. The first and second set times are stored in the X-ray source control device 22.

  By repeatedly outputting the X-ray stop signal and the X-ray generation signal, the X-ray source 17 emits X-rays during the first set time, that is, 1 μsec, and stops emitting X-rays during the second set time. . This emission and stop of X-rays are repeated during the movement period of the X-ray source 17 in the circumferential direction.

  The X-ray 57 emitted from the X-ray source 17 is irradiated to the examinee 42 in a fan beam shape. As the X-ray source 17 moves in the circumferential direction, the examinee 42 is irradiated with X-rays 57 from the surroundings. X-rays 57 that have passed through the examinee 42 (for example, X-rays that have passed through the affected area 56) 57 are centered around the radiation detector 5 located 180 degrees from the X-ray source 17 with the axial center of the hole 41 as the base point. It is detected by a plurality of radiation detectors 5 positioned in the direction. These radiation detectors 5 output detection signals of the X-rays 57. This X-ray detection signal is input to the corresponding signal discriminating device 27 via the corresponding wiring 49. Those radiation detectors 5 detecting the X-rays are referred to as first radiation detectors 4 for convenience.

  A 511 keV γ-ray 58 caused by the PET drug is emitted from the affected area (cancerous area) 56 of the examinee 42 on the bed 16. The radiation detectors 5 other than the first radiation detector 5 detect the γ-ray 58 and output a γ-ray detection signal. The radiation detector 5 that detects γ rays is referred to as a second radiation detector 5 for convenience. Among the second radiation detectors 5, the γ-ray detection signal output from the second radiation detector 5 positioned in the first layer is input to the corresponding signal discriminating device 27 via the corresponding wiring 49, and the second layer The γ-ray detection signal output from the second radiation detector 5 located in the third layer is input to the corresponding γ-ray discriminating device 32 via the wiring 54. Only the radiation detector 5 arranged in the first layer is connected to a signal discriminating device 61 having an X-ray signal processing device 33. This is because the X-ray energy is 80 keV, so that most (90% or more) X-rays transmitted through the examinee 42 are detected by the first radiation detector 5.

  In the signal discriminating device 27, the γ-ray detection signal output from the second radiation detector 5 of the first layer is transmitted to the γ-ray discriminating device 32, and the X-ray detection signal output from the first radiation detector 5 is X The signal is transmitted to the line signal processing device 33. Such transmission of each detection signal is performed by a switching operation of the changeover switch 28 of the signal discriminating device 27. The switching operation for connecting the movable terminal 29 of the switch 28 to the fixed terminal 30 or the fixed terminal 31 is performed based on a switching control signal that is an output of the drive device control device 21. At the time of X-ray CT examination, the drive device control device 22 selects the first radiation detector 5 among the first-layer radiation detection devices 5 and is movable in the signal discrimination device 27 connected to the first radiation detector 5. Terminal 29 is connected to fixed terminal 31.

  Selection of the first radiation detector 5 will be described. An encoder (not shown) is connected to the motor in the X-ray source driving device 15. The drive device control device 22 receives the detection signal of the encoder, obtains the position of the X-ray source drive device 15 in the circumferential direction of the detector support member 8 (hole 41), that is, the X-ray source 17, and this X-ray source. The radiation detector 5 positioned 180 ° opposite to the position 17 is selected using the stored position data of each radiation detector 5. Since the X-ray 57 emitted from the X-ray source 17 has a width that is the circumferential direction of the guide rail 12, the radiation detector 5 that detects the X-ray 57 that has passed through the examinee 42 is selected. In addition to the radiation detector 5, there exist a plurality in the circumferential direction. The driving device controller 22 also selects the plurality of radiation detectors 5. These radiation detectors 5 are the first radiation detectors 5. As the X-ray source 17 moves in the circumferential direction, the first radiation detector 5 also changes. As the X-ray source 17 moves in the circumferential direction, the first radiation detector 5 also appears to move in the pseudo circumferential direction. When the drive device controller 22 selects another radiation detector 5 as the X-ray source 17 moves in the circumferential direction, the movable device is newly connected to the radiation detector 5 that becomes the first radiation detector 5. The terminal 29 is connected to the fixed terminal 31. As the X-ray source 17 moves in the circumferential direction, the movable terminal 29 connected to the radiation detector 5 that is no longer the first radiation detector 5 is connected to the fixed terminal 30 by the drive device controller 22. Each radiation detector 5 in the first layer is a first radiation detector 5 when there is a relationship with the position of the X-ray source 17, and becomes a second radiation detector 5 when there is another. Therefore, one radiation detector 5 in the first layer outputs both an X-ray detection signal and a γ-ray detection signal with a time shift.

  The first radiation detector 5 detects X-rays that have been irradiated from the X-ray source 17 and transmitted through the examinee 42 during 1 μsec, which is the first set time. As described above, the probability that the first radiation detector 5 detects γ rays emitted from the examinee 42 during 1 μsec is negligibly small. A number of gamma rays 58 generated in the affected area 56 of the examinee 42 due to the PET drug are not emitted in a specific direction, but are emitted in all directions. As described above, these γ-rays 58 are emitted in pairs in substantially opposite directions (180 ° ± 0.6 °) and detected by any of the second radiation detectors 5.

  The signal processing of the signal discriminating device 27 when the X-ray detection signal and the γ-ray detection signal output from the first-layer radiation detector 5 are input will be described. The X-ray detection signal output from the first radiation detector 5 is input to the X-ray signal processing device 33 as described above. The X-ray signal processing device 33 integrates the input X-ray detection signal by the integrating device, and outputs an integrated value of the X-ray detection signal, that is, information on the measured X-ray intensity. The intensity information of the X-ray detection signal is transmitted to the computer 36 via the wiring 53 and stored in the storage device 37.

  The γ-ray detection signal output from the second radiation detector 5 of the first layer is input to the γ-ray discriminating device 32 by the action of the changeover switch 28. The energy of γ rays emitted from the affected part 56 due to the disappearance of the positrons emitted from the PET drug is 511 keV. However, when γ rays are scattered in the body of the examinee 42, the energy is lower than 511 keV. In order to remove scattered γ-rays, the γ-ray discriminating device 32 uses, for example, a filter (not shown) that passes a γ-ray detection signal having energy equal to or higher than the energy set value, with 400 keV being lower than 511 keV as an energy set value. ). This filter receives the γ-ray detection signal output from the fixed terminal 30. Here, as an example, the energy setting value is set to 400 keV because the variation of γ-ray detection signals generated when 511 keV γ-rays enter the radiation detector 5 is taken into consideration. The γ-ray discriminating device 32 generates a pulse signal having a predetermined energy when a γ-ray detection signal having an energy equal to or higher than the energy set value (400 keV) is input. The γ-ray discriminating device 32 is a γ-ray detection signal processing device, and adds time information and position information indicating the position of the radiation detector 5 connected to the γ-ray discriminating device 32 to the output pulse signal. The time information is any information of the time when the γ-ray detection signal is input to the γ-ray discriminator 32 and the time when the pulse signal is output from the γ-ray discriminator 32.

  The radiation detectors 5 in the second layer and the third layer are all second radiation detectors. The γ-ray discriminating device 32 connected to the radiation detectors 5 of the second layer and the third layer by the wiring 54 also exhibits the same function as the γ-ray discriminating device 32 in the signal discriminating device 27 described above.

  The coincidence counting device 34 receives the pulse signals output from all the γ-ray discriminating devices 32. The coincidence counting device 34 includes two second radiation detectors (approximately 180 ° ± 0.6 ° to be exact, centered on the axis of the hole 30) that detect the respective γ rays 58 of the γ ray pair. ) Simultaneous counting is performed using each pulse signal for each γ-ray detection signal output from a pair of second radiation detectors 5 present at different positions), and the count value for these γ-ray detection signals ( gamma ray counting information). The coincidence counting device 34 determines whether each pulse signal corresponds to a detection signal of each γ ray of the γ ray pair based on each time information given to those pulse signals. That is, if the difference between the two pieces of time information is within a set time (for example, 10 nsec), it is determined that the pulse signals are for a pair of γ rays 58 generated by the disappearance of one proton. Further, the coincidence counting device 34 converts each position information given to these pulse signals into data as each position of the corresponding pair of second radiation detectors 5, that is, position information of each γ-ray detection point. The coincidence counting device 34 outputs the count value information for each γ-ray detection signal and the position information of the two detection points where the γ-ray is detected. The count value and the position information are transmitted to the computer 36 and stored in the storage device 37.

  The computer 36 executes processing based on the processing procedure of steps 60 to 65 shown in FIG. The computer 36 that executes such processing creates first tomogram information using the first information (specifically, the γ-ray counting information and the position information of the γ-ray detection point), and the second information (specifically, Specifically, second tomogram information (specifically, X-ray CT image data) is created using X-ray intensity information and X-ray detection position information, and the first tomogram information and second tomogram information are used. And a tomographic image creation unit for creating third tomographic image information (specifically, synthetic tomographic image data) including such tomographic image information. The count value information of the γ-ray detection signal counted by the coincidence device 34, the position information of the γ-ray detection point output from the coincidence device 34, the X-ray intensity information output from the X-ray signal processing device 33, and the X X-ray detection position information given to the line intensity is input (step 60). The input count value information of the γ-ray detection signal, position information of the γ-ray detection point, X-ray intensity information, and X-ray detection position information are stored in the storage device 37 (step 61).

  Using the X-ray intensity information and the X-ray detection position information, a tomographic image of the cross section of the examinee 42 (hereinafter, the cross section refers to a cross section in a state where the examinee is standing) is reconstructed ( Step 62). The reconstructed tomographic image is referred to as an X-ray CT image. A specific process for the reconstruction of the tomographic image will be described. First, using the X-ray intensity information, the attenuation rate of X-rays in each voxel in the body of the examinee 42 is calculated. This attenuation factor is stored in the storage device 37. In order to reconstruct the X-ray CT image, using the attenuation rate of the X-ray detection signal read from the storage device 37, the position of the X-ray source 17 and the position of the radiation detector 5 (X The line attenuation coefficient in the body of the person to be examined 42 is obtained from (obtained from the line detection position information). The position of the X-ray source 17 at the time of movement detected by the encoder is given to the X-ray intensity information by each X-ray signal processing device 33 and transmitted to the computer 36. The CT value in each voxel is calculated based on the value of the linear attenuation coefficient in each voxel obtained by the filtered back projection method using the linear attenuation coefficient. X-ray CT image data is obtained using these CT values and stored in the storage device 37. In step 62, an X-ray CT image in a cross section passing through the affected area where the PET drug is accumulated is also reconstructed.

  A cross-sectional tomographic image of the examinee 42 including the affected part (for example, an affected part of cancer) is reconstructed using the count value of the γ-ray detection signal at the corresponding position (step 63). A tomographic image reconstructed using the count value of the γ-ray detection signal is referred to as a PET image. This process will be described in detail. Using the count value of the γ-ray detection signal read from the storage device 37, the pair of second radiation detectors 5 (identified from the position information of the γ-ray detection point) that detected the γ-rays generated by the annihilation of positrons. The number of γ ray pairs generated in the body between the semiconductor element portions (the number of γ ray pairs generated in response to the disappearance of a plurality of positrons) is obtained. Using this number of γ-ray pairs generated, the γ-ray pair generation density in each voxel is obtained by the filtered back projection method. Based on these γ-ray pair generation densities, PET image data can be obtained. The PET image data is stored in the storage device 37.

  The data of the PET image and the data of the X-ray CT image are synthesized, and the data of the synthesized tomographic image including both data is obtained and stored in the storage device 37 (step 64). By combining the PET image data at the position of the affected area and the X-ray CT image data at the position, the combined tomographic image data of the cross section of the examinee 42 at the position of the affected area is obtained. The synthesis of the PET image data and the X-ray CT image data can be easily and accurately performed by matching the position of the central axis of the hole 41 in both image data. That is, since the PET image data and the X-ray CT image data are created based on the detection signal output from the shared radiation detector 5, alignment can be performed with high accuracy as described above. The composite tomogram data is called from the storage device 37 and output to the display device 38 (step 65), and displayed on the display device 38. Since the synthetic tomographic image displayed on the display device 38 includes the X-ray CT image, the position of the affected part in the PET image in the body of the examinee 42 can be easily confirmed. That is, since the X-ray CT image includes images of the internal organs and bones, the doctor can specify the position where the affected part (for example, an affected part of cancer) exists in relation to the internal organs and bones.

  The radiation inspection apparatus 1 has a plurality of radiation detectors 5 stacked in the radial direction of the hole 41 (FIGS. 1 to 4), and this stacked layout can exhibit the following new functions. For example, as shown in FIG. 9A, two gamma rays 58a and 58b emitted from the gamma ray pair generation point 70 (in the affected area 56) in the body of the examinee 42 enter the radiation detectors 5f and 5g. Think about the case. Since it is not known at which position in the radiation detector the γ-ray is attenuated, in the conventional method, the line connecting the tip positions of the pair of radiation detectors 5f and 5h, that is, the line 71 shown in FIG. did. However, in the radiation inspection apparatus 1, since the radiation detectors 5 are stacked in the radial direction of the hole 41, a γ-ray detection signal of the radiation detector 5g located outside in the radial direction is obtained, and radiation detection is performed. A line 72 connecting the detector 5f and the radiation detector 5g can be a detection line. That is, it is possible to grasp the attenuation position in the depth direction of the radiation detector 5 that was not known in the conventional example. As a result, the detection line 72 accurately passes through the position where the γ ray pair is generated, so that the accuracy of the image is improved. As a result, the detection line becomes closer to the actual γ ray pair generation point, and the accuracy of the measurement data is improved.

  In this embodiment, since the radiation detection device 43 is composed of a plurality of radiation detectors 5 that output both X-ray detection signals and γ-ray detection signals, the radiation detection device 43 is a γ-ray detection unit and is an X-ray detector. It is also a detector. That is, the radiation detection device 43 has both functions of a γ-ray detection unit and an X-ray detection unit. In the present embodiment, the X-ray detection unit is located in a region formed between one end of the γ-ray detection unit and the other end of the γ-ray detection unit in the longitudinal direction of the bed 20. The radiation detection device 43 is an X-ray detection unit that detects X-rays 57 irradiated from the X-ray source 17 and transmitted through the examinee 42 and outputs a detection signal of the X-rays 57. Γ-ray 58 emitted from the PET drug is detected from a portion (affected part 56) through which X-rays 57 in the examinee 42 are transmitted at the position of the examinee 42 irradiating 57. This is a γ-ray detector that outputs 58 detection signals.

  According to the present embodiment, the following effects can be obtained.

  (1) In this embodiment, since a plurality of detector units 4 are attached to the detector support member 8 via the connector portion, these detector units 4, specifically, a large number of radiation detectors 5 are attached. It can be done in a short time. For this reason, the manufacturing time of the imaging device 2, ie, the radiation inspection apparatus 1, can be shortened.

  (2) Since the detector unit 4 is detachably attached to the detector support member 8 via the connector portion, when the radiation detector 5 fails, the detector unit 4 including the failed radiation detector 5 is detected. The support member 8 can be easily removed. Further, it can be easily attached to the detector support member 8 at the position of the detector unit from which the new detector unit 4 is removed. As described above, in this embodiment, the failed radiation detector 5 can be easily replaced.

  (3) In this embodiment, the radiation detector 5 is used in a conventional PET inspection by arranging a plurality of radiation detectors not only in the axial direction and circumferential direction of the hole 41 (detector support member 8) but also in the radial direction. A γ-ray detection signal at a position subdivided in the radial direction of the hole 41 can be obtained without reducing the signal transmission substance as in the case of a radiation detector. For this reason, the present embodiment can obtain accurate position information (position information of the radiation detector 5 that outputs the γ-ray detection signal) where the γ-rays have reached in the radial direction of the hole 41. In the conventional PET inspection, one radiation detector is arranged in the radial direction of the hole 41, and a reflecting material is arranged inside the radiation detector so that a signal transmitting substance reaches the photomultiplier tube. The information of the position where the γ rays reached in the radial direction of the hole 41 was obtained. At this time, a part of the signal transmission material is attenuated in the radiation detector or reflected to the outside of the radiation detector by the reflecting material, so that the signal transmission material is reduced and the energy resolution is lowered.

  (4) In this embodiment, since a plurality of independent radiation detectors 5 are arranged in the radial direction of the hole 41, all of the signal transmitting substances of the respective radiation detectors 5 can be used for detecting γ rays. The energy resolution of the radiation detector 5 is improved. When the radiation detector 5 having a high energy resolution is used in the PET examination, it becomes possible to distinguish between γ-rays whose energy has been attenuated by scattering and γ-rays having a non-scattering energy of 511 keV. As a result, more scattered rays can be removed by the filter of the γ-ray discriminating device 32.

  (5) Since this embodiment can acquire information on the exact arrival position of γ rays in the radial direction of the hole 31 without reducing the number of signal transmission substances in the radiation detector, The use of information improves the accuracy of tomographic images, and the need for a reflector for the radiation detector prevents the reduction of signal transmission materials, improving the energy resolution and reconstructing tomographic images of scattered radiation. It became possible to suppress the influence. As a result, this embodiment can improve the accuracy of the tomographic image, that is, the diagnostic accuracy of the PET examination.

  (6) Since the present embodiment uses a semiconductor radiation detector as the radiation detector 5, a plurality of radiation detectors 5 can be arranged in the radial direction of the hole 41, and a plurality of such radiation detectors are detected. Even if the device 5 is arranged, the imaging device 2 does not become large.

  (7) In this embodiment, since the radiation detector 5 is a semiconductor radiation detector, a photomultiplier tube is not required compared to a radiation detector using a scintillator, and the imaging device 2 can be downsized. it can.

  (8) According to the present Example, the radiation detector 5 which is a semiconductor radiation detector was arrange | positioned on a support substrate, and it became possible to arrange the radiation detector 5 densely. In particular, since the radiation detectors 5 having a small detector width can be densely arranged in the circumferential direction of the hole 41, high resolution (small image voxel size) of the tomographic image can be achieved.

  (9) According to the present embodiment, the radiation detectors 5 can be arranged densely by the configuration in which the radiation detectors 5 are arranged on the support substrate 6. In particular, it is possible to arrange a plurality of radiation detectors 5 in the radial direction of the hole 41, so that high detection efficiency can be achieved. Furthermore, since each radiation detector 5 in the radial direction can independently detect γ-rays, the resolution in the radial direction is improved. In particular, in 3D (three-dimensional) -PET inspection, γ-rays may be incident on the radiation detector 5 obliquely. However, by improving the resolution in the radial direction, the incident direction of γ-rays can be captured more accurately. be able to. For this reason, it is possible to improve the image quality of the obtained PET image.

  (10) According to the present embodiment, in order to arrange the wiring connected to the radiation detector 5 in the support substrate 6, the distance between the radiation detectors 5 in the circumferential direction of the hole 41 and the axial direction thereof is shortened. it can. The shortening of the interval between the radiation detectors 5 reduces the detection leakage of γ rays between the radiation detectors 5 and increases the substantial detection efficiency of γ rays. PET inspection time can be shortened by substantially increasing the detection efficiency of γ rays.

  (11) Since the radiation detector 5 that detects γ-rays is used as the radiation detector 5 that detects X-rays, the radiation inspection apparatus 1 uses the radiation detector 5 that detects X-rays and the radiation detection that detects γ-rays. It is not necessary to provide the vessel 5 separately, and the configuration can be simplified and the size can be reduced. The radiation detector 5 outputs both an X-ray detection signal and a γ-ray detection signal.

  (12) In this embodiment, the X-ray detection unit is located in a region formed between one end of the γ-ray detection unit and the other end of the γ-ray detection unit in the longitudinal direction of the bed 20. The first tomogram created based on the first information obtained from the γ-ray detection signal output from the γ-ray detection unit even when the examinee 42 moves during the examination regardless of the movement of the bed 20. The tomogram of the examinee 42 created by combining the information of (PET image) and the information of the second tomographic image (X-ray CT image) obtained from the X-ray detection signal output from the X-ray detection unit. The accuracy of the image can be improved. This can improve diagnostic accuracy for the subject by using the tomographic image. Specifically, the position and size of the affected area of cancer can be recognized with high accuracy. In particular, cancer of lymph glands that are organelles can be diagnosed with high accuracy.

  (13) In the present embodiment, as described above, the radiation detection device 43 includes a plurality of radiation detectors 5 that output both X-ray detection signals and γ-ray detection signals (X-ray detection for obtaining X-ray detection signals). Is performed using the radiation detector 5 that detects γ-rays to obtain a γ-ray detection signal), and thus has both functions of a γ-ray detection unit and an X-ray detection unit. It can be said that the radiation detector 43 has the γ-ray detector and the X-ray detector arranged coaxially. For this reason, the present embodiment uses an X-ray detection signal that is one output signal of the radiation detector 5 arranged in the circumferential direction of the detector support member 8 to image the internal organs and bones of the examinee 42. The first tomographic image at the position of the affected part (PET drug accumulation) can be reconstructed, and the affected part of the subject 42 is examined using a γ-ray detection signal that is another output signal of the radiation detector 5 The second tomogram including the image can be reconstructed. Since the data of the first tomogram and the data of the second tomogram are reconstructed based on the output signal of the radiation detector 5 that detects both transmitted X-rays and γ-rays, the first tomogram at the position of the affected part And the data of the second tomographic image can be accurately aligned and synthesized. For this reason, it is possible to easily obtain an accurate tomographic image (synthetic tomographic image) including images of the affected area, internal organs, bones and the like. According to this synthetic tomographic image, the position of the affected part can be accurately known in relation to the internal organs and bones. For example, by combining the data of the first tomographic image and the data of the second tomographic image based on the axis of the detector support member 8 (or the hole 41) of the imaging device 2, the two tomographic images are simply synthesized. Image data can be obtained.

  (14) In this embodiment, the X-ray detection unit 42 detects X-rays 57 irradiated from the X-ray source 17 and transmitted through the affected part 56 of the examinee 42, and irradiates the examinee 42. In order to detect the γ-ray emitted from the site (affected part) through which X-rays pass through the body of the examinee 42 at the position of the examinee 42 by the γ-ray detection unit, the examinee 42 is moved by the bed 20. X-ray CT inspection and PET inspection can be performed at the same position without moving. During both examinations, the X-ray detection unit outputs an X-ray detection signal transmitted through the affected area 56 of the examinee 42, and the γ-ray detection unit outputs a detection signal of γ-ray emitted from the affected area 56. In order to synthesize the first tomographic image data at the position of the affected area 56 obtained based on the X-ray detection signal and the second tomographic image data at the position of the affected area obtained based on the γ-ray detection signal, Even when the examinee 42 moves on the bed 20 without being able to endure during the examination, the tomographic image data can be synthesized with high accuracy. That is, highly accurate synthetic tomographic image data can be obtained. For this reason, the diagnostic accuracy of the affected part 52 can be improved by using the synthetic tomographic image data (synthetic tomographic image image) at the position of the affected part 56 displayed on the display device 38. In particular, even when an affected part exists in a place where an organ is complicated, the position of the affected part can be appropriately grasped by the synthetic tomographic image obtained in the present embodiment, and the diagnosis accuracy of the affected part is improved.

  (15) In this embodiment, the X-ray source 17 can be moved in the axial direction of the radiation detection unit 65 during the radiation examination period using an X-ray source axial movement device (for example, the axial movement arm 16). Without moving the examiner 42 in the axial direction of the radiation detection apparatus 43, the X-ray CT inspection can be performed on the inspection target range while performing the PET inspection on the inspection target range. When the X-ray CT examination for the examination target range is executed by moving the bed 20 without moving the X-ray source 17 in the axial direction, the position of the site where the PET drug is accumulated is also the same. Move in the axial direction. This means that the position where the γ ray pair is generated is moved in the axial direction, noise for the creation of PET image data increases, and accurate PET image data cannot be obtained. In this embodiment, since the position where the γ ray pair is generated does not move in the axial direction, highly accurate PET image data can be obtained, and the accuracy of the combined tomographic image data can be improved.

  (16) In this embodiment, the radiation detector 5 included in the radiation detection device 43 can detect a plurality of pairs of γ rays emitted from the examinee 42 and move in the circumferential direction. X-rays emitted from 17 and transmitted through the examinee 42 can also be detected. For this reason, the prior art has required an imaging device that detects X-rays and another imaging device that detects γ-rays as an imaging device, but this embodiment is a single imaging device that detects X-rays and γ-rays. Therefore, the configuration of the radiation inspection apparatus capable of performing both the X-ray CT inspection and the PET inspection can be simplified.

  (17) In the present embodiment, the X-ray detection signal necessary for creating the first tomographic image and the radiation detector 5 sharing the γ-ray detection signal necessary for creating the second tomographic image are used. Therefore, the time required for the examination of the examinee 42 (examination time) can be remarkably shortened. In other words, the X-ray detection signal necessary for creating the first tomographic image and the γ-ray detection signal necessary for creating the second tomographic image can be obtained in a short examination time. In the present embodiment, unlike the prior art, it is not necessary to move the examinee 42 from an imaging device that detects transmitted X-rays to another imaging device that detects γ-rays. Contribute further to shortening

  (18) In this embodiment, since the X-ray source 17 is circulated and the radiation detection device 43 is not moved in the circumferential direction and the axial direction of the hole 41, the X-ray source 17 is not compared with a motor necessary for moving the radiation detection device 43. The capacity of the motor that circulates the radiation source 17 can be reduced. The power consumption required to drive the latter motor can also be less than that of the former motor.

  (19) Since the γ-ray detection signals input to the X-ray signal processing device 33, that is, the first signal processing device are significantly reduced, highly accurate first tomographic image data can be obtained. For this reason, the position of the affected part can be known more accurately by using image data obtained by combining the data of the first tomographic image and the data of the second tomographic image.

  (20) In this embodiment, since the X-ray source 17 circulates inside the radiation detection device 43, the inner diameter of the detector support member 8 becomes large, and the radiation detection can be installed in the circumferential direction inside the detector support member 8. The number of vessels 5 can be increased. An increase in the number of radiation detectors 5 in the circumferential direction brings about an improvement in sensitivity and resolution, and improves the resolution of tomographic images in the cross section of the examinee 42.

  (21) In this embodiment, since the axial movement arm 16 and the X-ray source 17 are located inside the radiation detection apparatus 434, they are emitted from the examinee 42 during the X-ray CT examination. , The radiation detector 5 located immediately behind them cannot detect the γ-rays, and there is a possibility that detection data necessary for creating a PET image may be lost. However, in this embodiment, since the X-ray source 17 and the axially moving arm 16 circulate in the circumferential direction by the X-ray source driving device 15 as described above, data loss is not a problem in practice. . In particular, the revolving speed of the X-ray source 17 and the axial movement arm 16 is about 1 second / 1 slice, which is sufficiently short compared with the time required for the PET inspection on the order of several minutes at the shortest. Even in this case, the loss of the data is not substantially a problem. When the X-ray CT inspection is not performed and the PET inspection is performed, the X-ray source 17 is accommodated in the X-ray source driving device 15, so that the X-ray source 17 and the axial movement arm 16 are used for γ-ray detection. It will not be an obstacle.

  Furthermore, the inspection time required to obtain an X-ray detection signal necessary for creating an X-ray CT image is shorter than the inspection time required to obtain a γ imaging signal necessary for creating a PET image. Therefore, during the examination time for obtaining the γ-ray detection signal, the examinee 42 is under examination by always irradiating the examinee 42 with X-rays from the X-ray source 17 to obtain the X-ray detection signal. Even in the case of movement, it is possible to correct the deviation of the PET image data associated with the swing of the examinee 42 from the continuous image of the X-ray CT image obtained based on the X-ray detection signal.

  Although the wiring connected to the radiation detector 5 is arranged in the support substrate 6, a through hole is formed in the support substrate 6, and the wiring is passed through the through hole from the surface of the support substrate 6 on the side where the radiation detector 5 is installed. The other side may be drawn out to the opposite side by a through hole, and the wiring may be placed on the surface of the support substrate 6 on the side where the radiation detector 5 is not installed. In that case, a groove may be formed in the surface of the support substrate 6 on the side where the radiation detector 5 is not installed, and wiring may be installed in the groove. Further, a multilayer wiring board may be used as the support substrate, and wiring may be installed in the multilayer wiring board. Further, the radiation detector 5 can be arranged on both surfaces of the multilayer wiring board by using the multilayer wiring board.

(Example 2)
A radiation inspection apparatus according to embodiment 2, which is another embodiment of the present invention, will be described below. The configuration of the detector unit of the radiation inspection apparatus according to the present embodiment is the same as that of the radiation inspection apparatus 1 according to the first embodiment. A detector unit 4A used in this embodiment, which is different in configuration from the detector unit 4 used in Embodiment 1, will be described with reference to FIGS.

In the detector unit 4A, a plurality of (for example, nine) radiation detectors 5D are installed on one surface of the support substrate 6 in three rows and three columns. Each radiation detector 5 </ b> D is the same semiconductor radiation detector as the radiation detector 5, and is arranged in three layers in the radial direction of the detector support member 8. In a row aligned in the radial direction of the detector support member 8, the radiation detector 5A ₁ to the first layer, the radiation detector 5B ₁ a two-layer, and a radiation detector 5C ₁ in a third layer are respectively disposed.

The radiation detector 5A ₁ has five detection elements, that is, detection elements 74A, 74B, 74C, 74D, and 74E. The detection elements 74A, 74B, 74C, 74D, and 74E are arranged in that order from the inside in the radial direction of the detector support member 8. The detection element 74A is disposed on the innermost side, and the detection element 74E is disposed on the outermost side. A cathode electrode 77A is provided on the inner surface of the detection element 74A. The detection element 74A and the detection element 74B are adjacent to each other with an anode electrode 78A provided on the outer surface of the detection element 74A and the inner surface of the detection element 74B in between. The detection element 74B and the detection element 74C are adjacent to each other with a cathode electrode 77B provided on the outer surface of the detection element 74B and the inner surface of the detection element 74C in between. The detection element 74C and the detection element 74D are adjacent to each other with an anode electrode 78B provided on the outer surface of the detection element 74C and the inner surface of the detection element 74D interposed therebetween. The detection element 74D and the detection element 74E are adjacent to each other with a cathode electrode 77C provided on the outer surface of the detection element 74D and the inner surface of the detection element 74E in between. An anode electrode 78C is provided on the outer surface of the detection element 74E.

The radiation detector 5B ₁ has five detection elements, that is, detection elements 75A, 75B, 75C, 75D, and 75E. The detection elements 75A, 75B, 75C, 75D, and 75E are arranged in that order from the inside in the radial direction of the detector support member 8. The detection element 75A is disposed on the innermost side, and the detection element 75E is disposed on the outermost side. A cathode electrode 79A is provided on the inner surface of the detection element 75A. The detection element 75A and the detection element 75B are adjacent to each other with an anode electrode 80A provided on the outer surface of the detection element 75A and the inner surface of the detection element 75B interposed therebetween. The detection element 75B and the detection element 75C are adjacent to each other with a cathode electrode 79B provided on the outer surface of the detection element 75B and the inner surface of the detection element 75C in between. The detection element 75C and the detection element 75D are adjacent to each other with an anode electrode 80B provided on the outer surface of the detection element 75C and the inner surface of the detection element 75D interposed therebetween. The detection element 75D and the detection element 75E are adjacent to each other with a cathode electrode 79C provided on the outer surface of the detection element 75D and the inner surface of the detection element 75E interposed therebetween. An anode electrode 80C is provided on the outer surface of the detection element 75E.

The radiation detector 5C ₁ has five detection elements, that is, detection elements 76A, 76B, 76C, 76D, and 76E. The detection elements 76A, 76B, 76C, 76D, and 76E are arranged in that order from the inside in the radial direction of the detector support member 8. The detection element 76A is disposed on the innermost side, and the detection element 76E is disposed on the outermost side. A cathode electrode 81A is provided on the inner surface of the detection element 76A. The detection element 76A and the detection element 76B are adjacent to each other with an anode electrode 82A provided on the outer surface of the detection element 76A and the inner surface of the detection element 76B in between. The detection element 76B and the detection element 76C are adjacent to each other with a cathode electrode 81B provided on the outer surface of the detection element 76B and the inner surface of the detection element 76C in between. The detection element 76C and the detection element 76D are adjacent to each other with an anode electrode 82B provided on the outer surface of the detection element 76C and the inner surface of the detection element 76D interposed therebetween. The detection element 76D and the detection element 76E are adjacent to each other with a cathode electrode 81C provided on the outer surface of the detection element 76D and the inner surface of the detection element 76E in between. An anode electrode 82C is provided on the outer surface of the detection element 76E.

  The ground wire 45 is connected to the cathode electrodes 77A, 77B, 77C, 79A, 79B, 79C, 81A, 81B, 81C. A wiring 74 is connected to the anode electrodes 78A, 78B, 78C. A wiring 75 is connected to the anode electrodes 80A, 80B, 80C. A wiring 76 is connected to the anode electrodes 82A, 82B, and 82C. The ground wire 45 is connected to the connector terminal 7 </ b> D of the connector portion 7. The wiring 74 is connected to the connector terminal 7 </ b> A of the connector unit 7. The wiring 75 is connected to the connector terminal 7 </ b> B of the connector unit 7. The wiring 76 is connected to the connector terminal 7 </ b> C of the connector portion 7. Similarly, the radiation detectors 5D included in the other two rows are also connected to other connector terminals provided in the connector portion 7. The ground wire 45 and the wirings 74, 75, 76 are all installed in the support substrate 6. The multiple detector units 4A are mounted and held on the detector support 23 by fitting connector terminals such as connector terminals 7A provided on the detector units 4A into the connector 11 provided on the detector support 23. The Similarly to the detector unit 4, the detector unit 4 </ b> A surrounds the hole 41 and is arranged in a large number in the circumferential direction and the axial direction of the hole 41.

The radiation detectors 5A ₁ , 5B ₁ , 5C ₁ have three or more detection elements having at least two surfaces, that is, semiconductor elements, and anode electrodes and cathode electrodes are alternately arranged between different semiconductor elements. A specific description will be given using the radiation detector 5A ₁ . The radiation detector 5A ₁ includes different detection elements, ie, between the detection elements 74A and 74B, between the detection elements 74B and 74C, between the detection elements 74C and 74D, and between the detection elements 74D and 74D. The anode electrode and the cathode electrode are alternately arranged between the detection element 74EC, for example, an anode electrode 78A between the detection element 74A and the detection element 74B, and a cathode electrode 77B between the detection element 74B and the detection element 74C. It is arranged.

By fitting the connector portion 7 into the connector portion 11, the first three radiation detectors 5 </ b> A ₁ are separately connected to the three signal discriminating devices 27 in the signal discriminating unit 25. Further, the three radiation detectors 5B ₁ in the second layer and the three radiation detectors 5C ₁ in the third layer are provided in the signal discrimination unit 25, and the six γ-ray discrimination devices 32 other than the signal discrimination device 27 are provided. Connected separately.

  The radiation inspection apparatus according to the present embodiment incorporating the detector unit 4A can obtain the effects (1) to (21) produced by the radiation inspection apparatus 1 according to the first embodiment. Furthermore, this embodiment can obtain the following effects (22) and (23).

  (22) According to this embodiment, since the radiation detector 5D has a laminated structure of a plurality of detection elements, the thickness of each detection element between the anode electrode and the cathode electrode is reduced, and recombination of electrons and holes is performed. The decrease in the detection signal due to the is suppressed. For this reason, energy resolution improves. Further, since the time until the detection signal is output is shortened, the time resolution is also improved. Since the energy threshold can be set higher by improving the energy resolution, it becomes possible to remove more γ-rays whose energy has been reduced by scattering. In addition, since the time window can be reduced by improving the time resolution, γ rays that are accidentally detected in the time window can be reduced. That is, detection of scattering events and incidental events that are noise components can be suppressed to a low level, so that the image quality of PET images can be improved.

  (23) According to the present embodiment, the radiation detector 5D has a laminated structure of a plurality of detection elements, so that the thickness of the detection element between the anode electrode and the cathode electrode is reduced, and the applied bias voltage is reduced. be able to. Due to the decrease in the bias voltage, the withstand voltage of the parts around the various wirings can be reduced. In addition, the power supply itself can be reduced in size.

(Example 3)
A radiation inspection apparatus according to embodiment 3, which is another embodiment of the present invention, will be described below. The configuration of the detector unit of the radiation inspection apparatus according to the present embodiment is the same as that of the radiation inspection apparatus 1 according to the first embodiment. A detector unit 4B used in this embodiment, which has a different configuration from the detector unit 4 used in Embodiment 1, will be described with reference to FIGS.

  In the detector unit 4A, a plurality of (for example, nine) radiation detectors 5E are installed on one surface of the support substrate 6 in three rows and three columns. Each radiation detector 5E is the same semiconductor radiation detector as the radiation detector 5, and is arranged in three layers in the radial direction of the detector support member 8. In the radial direction of the detector support member 8, the radiation detectors 5Aa, 5Ab, 5Ac on the first layer, the radiation detectors 5Ba, 5Bb, 5Bc on the second layer, and the radiation detectors 5Ca, 5Cb, 5Cc on the third layer, respectively. Be placed. The radiation detectors 5Aa, 5Ab, 5Ac arranged in the circumferential direction of the detector support member 8 are stacked in the circumferential direction. Similarly, the second-layer radiation detectors 5Ba, 5Bb, and 5Bc and the third-layer radiation detectors 5Ca, 5Cb, and 5Cc are stacked in the circumferential direction of the detector support member 8. This laminated structure will be described by taking the radiation detectors 5Aa, 5Ab, and 5Ac as an example.

  The radiation detector 5Aa has four detection elements, that is, detection elements 83A, 83B, 83C, and 83D. The detection elements 83A, 83B, 83C, 83D are arranged in that order in the circumferential direction of the detector support member 8. The radiation detector 5Ab includes detection elements 84A, 84B, 84C, and 84D. The detection elements 84A, 84B, 84C, 84D are arranged in that order in the circumferential direction of the detector support member 8. The radiation detector 5Ac includes detection elements 85A, 85B, 85C, and 85D. The detection elements 85A, 85B, 85C, and 85D are arranged in that order in the circumferential direction of the detector support member 8.

  A cathode electrode 86A is provided on one side surface of the detection element 83A. The detection element 83A and the detection element 83B are adjacent to each other with an anode electrode 87A provided on the other side surface of the detection element 83A and one side surface of the detection element 83B interposed therebetween. Here, one side surface refers to one side surface of the detection element in the circumferential direction of the detector support member 8, and the other side surface refers to the remaining side surface of the detection element in the circumferential direction of the detector support member 8. The detection element 83B and the detection element 83C are adjacent to each other with a cathode electrode 86B provided on the other side surface of the detection element 83B and one side surface of the detection element 83C interposed therebetween. The detection element 83C and the detection element 83D are adjacent to each other with an anode electrode 87B provided on the other side surface of the detection element 83C and one side surface of the detection element 83D interposed therebetween.

  The detection element 83D and the detection element 84A are adjacent to each other with a cathode electrode 88A provided on the other side surface of the detection element 83D and one side surface of the detection element 84A interposed therebetween. The detection element 84A and the detection element 84B are adjacent to each other with an anode electrode 89A provided on the other side surface of the detection element 84A and one side surface of the detection element 84B interposed therebetween. The detection element 84B and the detection element 84C are adjacent to each other with a cathode electrode 88B provided on the other side surface of the detection element 84B and one side surface of the detection element 84C interposed therebetween. The detection element 84C and the detection element 84D are adjacent to each other with the anode electrode 89B provided on the other side surface of the detection element 84C and one side surface of the detection element 84D interposed therebetween.

  The detection element 84D and the detection element 85A are adjacent to each other with the cathode electrode 90A provided on the other side surface of the detection element 84D and one side surface of the detection element 85A interposed therebetween. The detection element 85A and the detection element 85B are adjacent to each other with an anode electrode 91A provided on the other side surface of the detection element 85A and one side surface of the detection element 85B interposed therebetween. The detection element 85B and the detection element 85C are adjacent to each other with the cathode electrode 90B provided on the other side surface of the detection element 85B and one side surface of the detection element 85C interposed therebetween. The detection element 85C and the detection element 85D are adjacent to each other with an anode electrode 91B provided on the other side surface of the detection element 85C and one side surface of the detection element 85D interposed therebetween. A cathode electrode 90C is provided on the other side surface of the detection element 85D.

  The ground wire 92 is connected to the cathode electrodes 86A, 86B, 88A, 88B, 90A, 90B, and 90C. A wiring 93 is connected to the anode electrodes 87A and 87B. A wiring 94 is connected to the anode electrodes 89A and 89B. A wiring 95 is connected to the anode electrodes 91A and 91B. The ground wire 92 is connected to the connector terminal 7 </ b> D of the connector portion 7. The wiring 93 is connected to the connector terminal 7 </ b> A of the connector portion 7. The wiring 94 is connected to the connector terminal 7B of the connector portion 7. The wiring 95 is connected to the connector terminal 7 </ b> C of the connector portion 7. Similarly, the radiation detectors 5E of the second layer and the third layer are also connected to other connector terminals provided in the connector portion 7. The ground wire 92 and the wirings 93, 93, 93 are all installed in the support substrate 6. The multiple detector units 4B are mounted and held on the detector support 23 by fitting connector terminals such as the connector terminals 7A provided on the connectors into the connector 11 provided on the detector support 23. The Similarly to the detector unit 4, the detector unit 4 </ b> B surrounds the hole 41 and is arranged in a large number in the circumferential direction and the axial direction of the hole 41.

  By fitting the connector portion 7 into the connector portion 11, the first-layer radiation detectors 5 Aa, 5 Ab, and 5 Ac are separately connected to the three signal discriminating devices 27 in the signal discriminating unit 25. The second-layer radiation detectors 5Ba, 5Bb, 5Bc and the third-layer radiation detectors 5Ca, 5Cb, 5Cc are provided in the signal discriminating unit 25, and the six γ-ray discriminators other than the signal discriminating device 27 are provided. Separately connected to device 32.

  The radiation inspection apparatus of the present embodiment incorporating the detector unit 4B has the effects (1) to (21) generated in the radiation inspection apparatus 1 of the first embodiment and the effects (22) generated in the radiation inspection apparatus of the second embodiment. (23) can be obtained. Furthermore, the present Example can obtain the following effect (24).

  (24) According to the present embodiment, since each radiation detector 5E has a laminated structure of an even number of detector elements, both side surfaces of the adjacent radiation detector 5E can be used as cathode electrodes. The cathode electrode can be shared by the radiation detector 5E. Therefore, the three radiation detectors 5E arranged in the circumferential direction of the detector support member 8 can be brought into close contact with each other. That is, it becomes possible to completely eliminate the interval between the radiation detectors 5E in the circumferential direction, and the detection leakage of γ rays between the radiation detectors 5E in the circumferential direction is remarkably reduced. This leads to a substantial increase in the detection efficiency of γ rays, leading to a reduction in inspection time.

  In the first to third embodiments, a detector unit in which a radiation detector is stacked in the radial direction of the detector support member 8 (the hole 41 into which the bed 20 is inserted) is used as an X-ray source from an X-ray source 17. It has the structure applied to the radiography apparatus which can detect the gamma ray and X-ray which can irradiate. However, the detector unit can also be applied to a PET radiological examination apparatus that detects only γ rays emitted from the subject due to the radiopharmaceutical that has passed through the subject without being irradiated with X-rays. . The detector unit can also be applied to a radiation inspection apparatus for SPECT.

It is a longitudinal cross-sectional view of the radiation inspection apparatus of Example 1 which is one suitable Example of this invention. It is II-II sectional drawing of FIG. It is an enlarged view of the III section of FIG. It is IV-IV sectional drawing of FIG. It is a perspective view of the detector unit of FIG. It is sectional drawing in the radial direction of the detector support member of the detector unit of FIG. It is a detailed structure figure of the signal discriminating device of FIG. It is explanatory drawing which shows the process sequence of the tomogram preparation performed with the computer of FIG. It is explanatory drawing which shows the state of the gamma ray detection in the Example of FIG. It is a perspective view of the detector unit applied to the radiation inspection apparatus of Example 2 which is another Example of this invention. It is sectional drawing in the radial direction of the detector support member of the detector unit of FIG. It is a perspective view of the detector unit applied to the radiation inspection apparatus of Example 3 which is another Example of this invention. It is sectional drawing in the radial direction of the detector support member of the detector unit of FIG.

Explanation of symbols

  DESCRIPTION OF SYMBOLS 1 ... Radiation inspection apparatus, 2 ... Imaging device, 4, 4A, 4B ... Detector unit, 5, 5D, 5E ... Radiation detector, 7, 11 ... Connector part, 8 ... Detector support member, 13 ... X-ray source Circumferential movement device, 14 ... X-ray source device, 15 ... X-ray source driving device, 17 ... X-ray source, 18 ... Examinee holding device, 20 ... Bed, 23 ... Detector support, 25 ... Signal discrimination unit , 27 ... signal discriminating device, 28 ... changeover switch, 32 ... γ-ray discriminating device, 33 ... X-ray signal processing device, 34 ... coincidence counting device, 35 ... tomographic image creating device, 36 ... computer, 40 ... signal processing device.

Claims (8)

A radiation detector including a detector support member disposed around a bed for supporting a subject, and a plurality of detector units disposed around the bed in the longitudinal direction of the bed;
The detector unit has a plurality of radiation detectors for detecting radiation,
The radiation detector has a plurality of semiconductor members arranged in parallel, and detection signal output electrodes and common potential electrodes arranged alternately, and the detection signal output electrode and the common potential electrode are between these electrodes. Arranged in the semiconductor member arranged,
A plurality of the radiation detectors are arranged in a radial direction of the detector support member,
A radiation inspection apparatus, comprising: a radiation detection signal processing device that obtains positional information of the radiation detector in the radial direction of the radiation detector that has output a γ-ray detection signal.   The detector unit includes a detector support substrate, a plurality of the radiation detectors installed on the detector support substrate, and is connected to each of the radiation detectors provided on the detector support substrate. The radiation inspection apparatus according to claim 1, further comprising: a plurality of wirings that transmit detection signals output from the detector.   The radiation inspection apparatus according to claim 2, further comprising an image creating apparatus that creates an image of the subject using an output signal of the radiation detector.   The radiation inspection apparatus according to claim 2, wherein the wiring is disposed on the detector support member. A detector support member that extends in the longitudinal direction of the bed that supports the subject and is disposed around the bed; and a plurality of detector units that are disposed in the longitudinal direction of the bed and the circumferential direction of the detector support member. A radiation detector including the plurality of detector units detachably attached to the detector support member,
The detector unit has a plurality of radiation detectors for detecting radiation, and a plurality of the radiation detectors are arranged at different positions in a radial direction of the detector support member,
The radiation detector has a plurality of semiconductor members arranged in parallel, and detection signal output electrodes and common potential electrodes arranged alternately, and the detection signal output electrode and the common potential electrode are between these electrodes. Installed in the semiconductor member arranged,
A radiation inspection apparatus, comprising: a radiation detection signal processing device that obtains positional information of the radiation detector in the radial direction of the radiation detector that has output a γ-ray detection signal. The detector unit includes a detector support substrate that is detachably attached to the detector support member, a plurality of radiation detectors installed at different positions in the radial direction of the detector support substrate, and the detector support substrate. The radiation inspection apparatus according to claim 5, further comprising: a plurality of wires that are connected to each of the radiation detectors and that transmit a γ-ray detection signal output from the radiation detector.   The radiation inspection apparatus according to claim 6, further comprising an image creating apparatus that creates an image of the subject using an output signal of the radiation detector.   The radiation inspection apparatus according to claim 6, wherein the wiring is disposed on the detector support member.

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