JP2008089354A - Semiconductor radiation detector and positron emission tomography device - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To improve sensitivity of radiation detection by having a semiconductor element appropriately held, in a semiconductor radiation detector having a laminated structure of a semiconductor element and an electrode plate. <P>SOLUTION: The semiconductor radiation detector 1 has a laminated structure composed of semiconductor elements 11 converting received radiation into electric signals and metal electrode plates 12C and 12A which are alternately bonded with an electroconductive adhesive 14, wherein at least one of the electrode plates 12C and 12A composing the laminated structure is formed with a cutout part 13 in an area contacting with the semiconductor element 11. <P>COPYRIGHT: (C)2008,JPO&INPIT

Description

  The present invention relates to a semiconductor radiation detector and a positron emission tomographic imaging apparatus.

In recent years, nuclear medicine diagnostic apparatuses using radiation measurement technology have become widespread. Typical examples are a positron emission tomography apparatus (PET imaging apparatus), a single photon emission tomography apparatus (SPECT imaging apparatus), a gamma camera apparatus, and the like. The radiation detector mainly used in these apparatuses is a combination of a scintillator and a photomultiplier tube. The scintillator emits light when radiation is incident, and the weak light is amplified by a photomultiplier tube to detect the radiation. On the other hand, not only scintillators but also semiconductors can be used for radiation measurement. In semiconductors, when radiation is incident, charges such as holes and electrons are generated by the photoelectric effect, and these are moved by an electric field generated by an external voltage applied to the semiconductor. Since this amount of charge is proportional to the energy of the radiation, the energy of the radiation can be accurately known by accurately measuring the amount of charge. Therefore, the semiconductor has an advantage that the energy of radiation can be measured more accurately than the combination of a scintillator and a photomultiplier tube.
Here, there are several types of semiconductor elements used, but the effective atomic number is large in terms of sensitive measurement of gamma rays used in nuclear medicine diagnostic equipment, and room temperature operation is possible in terms of handling. A material using a semiconductor material such as cadmium telluride (CdTe) or zinc cadmium telluride (CdZnTe) is specifically mentioned.

  A semiconductor radiation detector formed using such a semiconductor element is superior to a scintillator in that it can accurately measure the energy of radiation (γ-rays) as described above. Is not necessarily excellent in terms of the accuracy of the time at which the light enters. The sensitivity tends to be higher as the atomic number of the substance is larger. The effective atomic number of CdTe or CdZnTe is about 50, whereas the scintillator is, for example, BGO (Bi4Ge3O12, bismuth germanate). And GSO (Gd2SiO5, gadolinium silicate) has an effective atomic number exceeding 60, and is more sensitive than a semiconductor radiation detector. Therefore, in the semiconductor radiation detector, in order to increase sensitivity, it is effective to increase the region through which γ rays pass as much as possible, that is, to increase the volume of the semiconductor element.

  However, CdTe and CdZnTe exhibit a characteristic that when the volume is increased, the ability to accurately measure the energy of γ rays decreases. This is presumably because the carrier lifetime and carrier moving speed of CdTe and CdZnTe are not sufficiently large, and when the volume of the semiconductor element is increased, the number of carriers that recombine and disappear in the middle increases. In other words, the variation in the amount of charge increases depending on the generation position of carriers inside the semiconductor element, and as a result, the accurate amount of charge cannot be measured. Furthermore, if the volume of the semiconductor element is increased, the carrier movement time becomes longer, so that the charge moves over a long period of time, leading to a decrease in time accuracy with respect to the time of incidence of γ rays. This is very disadvantageous particularly when simultaneous measurement of annihilation gamma rays is performed as in a PET imaging apparatus. This is because if the time accuracy decreases, it becomes difficult to identify the annihilation γ-ray, and the flight direction cannot be specified.

  Therefore, in a semiconductor radiation detector in which a semiconductor element is formed of, for example, CdTe, in order to avoid the above-described problem, it is conceivable that the semiconductor element is formed to be relatively thin and stacked. According to such a semiconductor radiation detector, since the semiconductor element is made relatively thin, the carrier movement time can be shortened, and the volume can be increased because of the laminated structure. Examples of the semiconductor radiation detector configured in this way include those disclosed in Patent Documents 1 and 2.

  Also, the electrode made of a conductive thin film directly bonded to the semiconductor element is divided into two regions, and the charge generated in the peripheral region of the semiconductor element is collected by one guard electrode and generated in the other region of the semiconductor element. In other words, there is disclosed a technique in which a noise signal other than a radiation detection signal is removed by collecting the charge to be collected by the other measurement electrode (see, for example, Patent Document 3).

JP-A-8-160147 JP-A-11-281747 JP-A-2005-77152

However, it has been found that in a semiconductor radiation detector in which semiconductor elements are formed thinner and stacked, the sensitivity per volume of the semiconductor elements is slightly reduced. For example, a semiconductor radiation detector in which four semiconductor elements having a thickness of 1.0 mm and an electrode plate are laminated, and a semiconductor radiation detector in which eight semiconductor elements having a thickness of 0.5 mm and an electrode plate are laminated are manufactured. When a sensitivity measurement test for 511 keV was performed, the sensitivity of a semiconductor radiation detector in which a 0.5 mm-thick semiconductor element was stacked was about 10% lower than that of a semiconductor radiation detector in which a 1.0 mm-thick semiconductor element was stacked. As a result.
On the other hand, a semiconductor radiation detector in which semiconductor elements having a thickness of 0.5 mm are stacked has a semiconductor radiation thickness of 1.0 mm with respect to characteristics such as energy resolution indicating accuracy of radiation energy and time resolution indicating accuracy of radiation incident time. It showed better characteristics than the detector. Therefore, it was confirmed that the characteristics were not deteriorated by forming the semiconductor element thin.

  Therefore, in order to investigate the cause of the above-described decrease in sensitivity, a semiconductor radiation detector was fabricated using an extremely small electrode plate that only connects the peripheral part of the semiconductor element and an external circuit on a test basis, and its sensitivity. Was measured. As a result, the sensitivity of 511 keV γ rays was increased by about 10% in the case of stacking 1.0 mm-thick semiconductor elements, compared with the case of testing with a normal electrode plate. Further, in the case of stacking semiconductor elements having a thickness of 0.5 mm, the sensitivity increased by about 18%. Therefore, it was found that if the electrode plate is small, the sensitivity is improved accordingly.

  By the way, in the semiconductor radiation detector having the laminated structure as described above, the semiconductor element is held from both sides by the electrode plate, and the electrode plate has a holding function. Therefore, as described above, if the electrode plate is formed small for improving the sensitivity, the holding state of the semiconductor element becomes unstable, and the laminated structure cannot be suitably maintained. In particular, since CdTe is a soft and soft property, it is difficult to hold a semiconductor element in a structure in which an electrode plate is formed small.

  In addition, when an electrode plate for taking out a signal as described above is provided for the semiconductor radiation detector disclosed in Patent Document 3, the electrode plate is connected on the measurement electrode made of a thin film. For this reason, the semiconductor radiation detector disclosed in the cited document 3 has the same problem as that of the semiconductor radiation detector described in Patent Documents 1 and 2.

  Accordingly, an object of the present invention is to solve the above-described problems, and in a semiconductor radiation detector in which a semiconductor element and an electrode plate have a laminated structure, the semiconductor element can be suitably held and the sensitivity of radiation detection can be increased. An object of the present invention is to provide a semiconductor radiation detector and a positron emission tomography apparatus that can be improved.

  In order to achieve the above-described object, in the present invention, at least one electrode plate constituting the laminated structure of the semiconductor radiation detector has a configuration in which a notch is formed in a region in contact with the semiconductor element. According to this configuration, the notch formed in the electrode plate can be used as a passage space for primary electrons generated by the photoelectric effect due to the incidence of radiation. In other words, even if primary electrons generated by the photoelectric effect jump out toward the electrode plate, the primary electrons enter the adjacent semiconductor element through the notch, and the probability of being absorbed by the electrode plate is small. Become. Therefore, by making primary electrons annihilated in the conventional electrode plate incident on adjacent semiconductor elements, the primary electrons that generate carriers in the adjacent semiconductor elements increase, which is effective. Can be used as a signal. This improves the sensitivity of radiation detection. Moreover, since the notch formed in the electrode plate is formed in a region in contact with the semiconductor element, the semiconductor element can be suitably held by a portion other than the notch.

Further, by disposing the electrode plate in which the notch is formed only between the stacked semiconductor elements adjacent to each other, primary electrons are prevented from jumping out from the electrode plates stacked at both ends of the semiconductor radiation detector. Therefore, it is possible to obtain a semiconductor radiation detector having improved sensitivity of radiation detection while preventing the incidence of primary electrons to other adjacent semiconductor radiation detectors.
In addition, by giving magnetism to the electrode plate, handling using electromagnetic tweezers etc. can be performed, handling for stacking semiconductor elements and electrode plates, and mounting a semiconductor radiation detector on the substrate, etc. As a result, productivity can be improved and assembly cost can be reduced.

  Moreover, in the positron emission tomographic imaging apparatus using such a semiconductor radiation detector, the sensitivity of the semiconductor radiation detector is improved, so that the amount of radiopharmaceutical can be reduced and the measurement time can be shortened. Moreover, since the increase in the number of measurement persons can be expected by shortening the measurement time, the economic effect is high and the cost can be reduced.

  ADVANTAGE OF THE INVENTION According to this invention, in the semiconductor radiation detector by which the semiconductor element and the electrode plate were laminated structure, while being able to hold | maintain a semiconductor element suitably, the sensitivity of a radiation detection can be improved And a positron emission tomographic imaging apparatus.

Next, the semiconductor radiation detector of the present invention will be described in detail with reference to the drawings as appropriate.
(First embodiment)
A semiconductor radiation detector (hereinafter simply referred to as a detector) 1 of the present embodiment is arranged between four semiconductor elements 11 and between the semiconductor elements 11 and at both ends of the semiconductor element 11 as shown in FIG. The electrode plates 12C and 12A are provided in a laminated structure, and a notch 13 which is a characteristic configuration of the present embodiment is formed in the electrode plates 12C and 12A (see FIG. 2B).

  As shown in FIG. 2A, the semiconductor element 11 includes a semiconductor material (semiconductor crystal) 11a formed in a flat plate shape, and a thin film electrode is formed over the entire surface of both sides by vapor deposition or the like. ing. Here, an electrode formed on one surface of the semiconductor element 11 is a cathode electrode (hereinafter referred to as a cathode) C, and an electrode formed on the other surface is an anode electrode (hereinafter referred to as an anode) A.

  The semiconductor material 11a forms a region that generates electric charges by interacting with radiation (γ rays or the like), and is formed of any single crystal such as CdTe, CdZnTe, or GaAs. The cathode C and the anode A are made of any material such as Pt, Au, and In. In the present embodiment, for example, CdTe is used as the semiconductor material 11a, and a cathode C mainly composed of Pt and an anode A mainly composed of In are formed thereon to form a pn junction diode. Therefore, handling using electromagnetic tweezers or the like is possible.

As shown in FIG. 2B, since the electrode plates 12C and 12A have the same shape, the electrode plate 12C will be described below, and the electrode plate 12A will be described as appropriate. The electrode plate 12C is a thin plate-like member in which a cutout portion 13 cut out in a quadrangular shape is formed in a region in contact with the semiconductor element 11 (see FIG. 2A, hereinafter the same), for example, iron-nickel. It is formed from at least one of an alloy, an iron-nickel-cobalt alloy, chromium, and tantalum. Here, 42 alloy (Fe 58%, Ni 42%) can be used as the iron-nickel alloy, and Kovar (Fe 54%, Ni 29%, Co 17%) can be used as the iron-nickel-cobalt alloy. . That is, the electrode plate 12C formed of such a material has magnetism.
The notch 13 is formed by being simultaneously punched when the electrode plate 12C is press-molded. In the present embodiment, a total of four notches 13 are formed at predetermined intervals on the top, bottom, left, and right of the electrode plate 12C. The electrode plate 12 </ b> C has an outer shape of a portion in contact with the semiconductor element 11 formed linearly along the outer shape of the semiconductor element 11, and has a rectangular frame shape as a whole with four notches 13 inside. It has become a thing.

Further, the outer shape of the electrode plate 12C, specifically, the outer shape of at least the portion in contact with the semiconductor element 11 is larger than the outer shape of the semiconductor element 11 as shown in FIG. Accordingly, the detector 1 can be configured by stacking the electrode plates 12C and 12A so that the upper end and the left and right side ends thereof protrude outward from the upper end and the left and right side ends of the semiconductor element 11 (see FIG. 1). In the detector 1 configured in this way, the upper end and the left and right side ends of the semiconductor element 11 are located inside the upper ends and the left and right side ends of the electrode plates 12C and 12A. Is prevented.
The outer dimensions of the electrode plate 12C may be the same size as the semiconductor element 11. The thickness of the electrode plate 12C is preferably about 10 μm to 100 μm, and preferably about 50 μm.

  Further, as shown in FIG. 2B, the electrode plate 12C is provided with a protruding portion 12a (12b) that hangs down from the semiconductor element 11 (on the side of the wiring board 24 shown in FIG. 1). Yes. The protruding portion 12a (12b) functions as a fixing portion for electrically attaching the detector 1 to the wiring board 24. The fixing is performed using solder (not shown) or the like.

  In the detector 1 including the semiconductor element 11 and the electrode plates 12C and 12A, the semiconductor elements 11 are arranged in parallel so that the cathodes C and the anodes A face each other as shown in FIG. The electrodes of the same type (the anodes C and the cathodes A) are electrically connected via the electrode plates 12C and 12A. That is, the electrode plate 12 </ b> C is disposed between the opposing cathodes C of the adjacent semiconductor elements 11, and is attached to each cathode C by the conductive adhesive 14. The electrode plate 12 </ b> A is disposed between the facing anodes A of the adjacent semiconductor elements 11 and is attached to each anode A by the conductive adhesive 14. Further, the electrode plate 12 </ b> C is bonded to the cathodes C located at both ends of the detector 1 by the conductive adhesive 14. As described above, the detector 1 is configured such that the cathodes C and the anodes A are alternately arranged, and the electrode plates 12C and 12A are also alternately arranged.

As the conductive adhesive 14, for example, a material in which conductive particles such as metal powder (silver) are dispersed in an insulating resin binder made of an organic polymer material is used. Usually, when the semiconductor element 11 and the electrode plates 12C and 12A are bonded by the conductive adhesive 14, a high-temperature heat treatment at about 120 to 150 ° C. is performed in order to cure the conductive adhesive 14.
In the present embodiment, as shown in FIGS. 1, 2A and 2C, the crossing portion 12d left in a cross shape at the center of the electrode plates 12C and 12A is bonded to the semiconductor element 11. As described above, the conductive adhesive 14 is applied to the central portion of the surface on which the cathode C and the anode A of the semiconductor element 11 are formed.
The application position of the conductive adhesive 14 is not limited to the intersecting portion 12d as long as the semiconductor element 11 can be held by the electrode plates 12C and 12A. For example, as shown in FIG. As such, the diagonal portions 11b and 11b of the electrode plate 12C may be used.

  As shown in FIG. 1, the detector 1 is connected to the connecting member CP for the cathode C provided on the wiring board 24 with the protruding portion 12a of the electrode plate 12C on the cathode C side. The protruding portion 12b of the electrode plate 12A on the anode A side is connected to the connecting member AP for the anode A provided on the substrate 24. In addition, the detector 1 is attached to the wiring board 24 in a non-contact state by these protrusions 12 a and 12 b, that is, in a state where the semiconductor element 11 has a predetermined gap S between the detector 1 and the wiring board 24. Thereby, the insulation fall resulting from dust etc. being pinched | interposed between the detector 1 and the wiring board 24 at the time of attachment can be prevented suitably. In addition, since the air permeability is increased through the gap S, the detector 1 can be cooled, and the characteristics of the detector 1 are stabilized. Note that an insulating material (not shown) may be coated on the bottom surface of the detector 1 to further prevent unexpected dielectric breakdown.

  Here, an outline of the principle of detection of γ rays by the detector 1 will be described. As shown in FIG. 3A, when γ-rays enter the detector 1 to generate primary electrons E, the γ-rays interact with the semiconductor material 11a of the semiconductor element 11 to generate holes. Electrons are generated in pairs by an amount proportional to the energy of the primary electrons E. By the way, between the cathode C and the anode A of the semiconductor element 11 constituting the detector 1, a reverse bias voltage for collecting charges from the DC high voltage power supply 20 (for example, the cathode C is −500 V, the anode A potential close to the ground potential, that is, a reverse applied voltage such that the anode A becomes 500 V higher than the cathode C) is applied. For this reason, holes corresponding to positive charges are attracted and moved to the cathode C, and electrons that are negative charges are attracted and moved to the anode A. When these holes and electrons are compared, the ease of movement (mobility) is relatively large, so that the electrons reach the anode A in a relatively short time. On the other hand, since the holes are relatively less likely to move, the holes reach the cathode C over time. Incidentally, electrons and holes may be trapped in the middle before reaching the anode A and the cathode C.

Here, as shown in FIG. 3B, when the incident position of the γ-ray is in the vicinity of the cathode C and the anode A which are the end portions of the semiconductor material 11a (in the figure, in the vicinity of the anode A), γ The primary electrons E generated by the photoelectric effect due to the incidence of the line may jump out toward the adjacent semiconductor element 11 ′ before generating carriers in the semiconductor material 11a.
Here, the electrode plate 12A is disposed between the semiconductor element 11 and the adjacent semiconductor element 11 ′, and the notch 13 is formed in the electrode plate 12A, so that the primary electrons E are adjacent to each other. Even if it jumps out toward the semiconductor element 11 ′, the probability of passing through the notch 13 of the electrode plate 12A as it is and entering the semiconductor material 11a ′ of the adjacent semiconductor element 11 ′ increases. That is, the probability that the ejected primary electrons E are absorbed by the electrode plate 12A is reduced.

  As shown in FIG. 3C, if the electrode plate EB of the anode A is a normal electrode plate EB (electrode plate without the notch 13) having a plate-like surface as in the prior art. The ejected primary electrons E enter the electrode plate EB and are absorbed without generating carriers. That is, the primary electrons E are absorbed by the electrode plate EB without taking out an electrical signal from the primary electrons E.

  On the other hand, according to the detector 1 of the present embodiment, since the notches 13 are formed in the electrode plates 12C and 12A as described above, as shown in FIG. Even if the primary electrons E due to incidence jump out toward the adjacent semiconductor element 11 ′, the primary electrons E easily enter the adjacent semiconductor element 11 ′ through the notch 13 of the electrode plate 12 A. As a result, the adjacent semiconductor element 11 The primary electrons E that generate carriers in the “semiconductor material 11a” increase. That is, as shown in FIG. 3C, primary electrons E that have been annihilated in the past can be taken out as effective electrical signals by the adjacent semiconductor element 11 '. Therefore, the sensitivity of the detector 1 is improved.

As described above, the sensitivity of the detector 1 is determined by measuring the integral value of the current generated and generated when carriers (γ rays) are incident, that is, the amount of carrier charge proportional to the energy of the radiation. I can know.
Thus, the carrier charge amount was measured using the 137Cs 662 keV γ ray with a source intensity of 2500 kBq for the detector 1 of this embodiment, and the sensitivity was measured. The detector 1 was installed at a position 200 mm away from the radiation source. As a result, the count of the peak at 662 keV was 1,0056 times per hour.
As a comparative example, a detector using an electrode plate EB in which the notch 13 is not formed as shown in FIG. 3 (c) is prepared, and this is installed at a position 200 mm away from the radiation source, and the sensitivity is similarly obtained. Measurements were made. In addition, the detector of this comparative example was different only in the electrode plate EB, and used the semiconductor material 11a which has the same volume.

As a result, the detector of the comparative example had a peak count of 662 keV of 8995 times per hour.
Therefore, when the detector 1 of the present embodiment and the detector of the comparative example are compared and evaluated by counting, the detector 1 of the present embodiment is improved in sensitivity by 12% compared to the detector of the comparative example. The result was recognized.

Here, a modification of the detector 1 is shown in FIGS. The detector 1A shown in FIG. 4A is obtained by providing a plurality of circular cutouts 13 ′ on the electrode plates 12C ₁ and 12A ₁ , and other configurations are not changed. Since the electrode plates 12C ₁ and 12A ₁ have the same configuration, the electrode plate 12C ₁ will be described below. As shown in FIG. 4 (b), notch 13 ', the surface in contact with the semiconductor device 11 of the electrode plate 12C _1, 4 pieces in the longitudinal direction, are provided five laterally electrode plate 12C as a whole and 20 provided in the _1. In the detector 1A, while provided many notch 13 ', (the area of the portion in contact with the semiconductor device 11) area of cut remaining portion of the electrode plate 12C ₁ also said the detector 1 (see FIG. 1) since larger than the area of the cut remaining portion of the electrode plate 12C in the rigidity of the electrode plate 12C ₁ is high, and excellent retention of the semiconductor element 11. Also, not easily deformed or the like when punched electrode plates 12C ₁ to form the notch 13 ', it is excellent in shape retention of the electrode plate 12C _1. Therefore, by using such electrode plates 12C ₁ and 12A ₁ , it is possible to obtain a detector 1A having excellent sensitivity and high dimensional accuracy.

  In the detector 1A configured as described above, as described above, a 137Cs 662 keV γ ray with a source intensity of 2500 kBq is used, and the detector 1A is installed at a position 200 mm away from the source to measure sensitivity. It was. As a result, the peak count of 662 keV was 9511 times per hour, and an improvement in sensitivity was confirmed compared to the comparative example (see FIG. 3C).

5A, the electrode plates 12C ₂ and 12A ₂ are each formed with two large notches 13 ″ and 13 ″ on the top and bottom, and viewed from the front (viewed in the stacking direction). The electrode plates 12C ₂ and 12A ₂ have an H shape. Since the electrode plates 12C ₂ and 12A ₂ have the same configuration, the electrode plate 12C ₂ will be described below. Electrode plate 12C _2, as shown in FIG. 5 (b), the left and right vertical plate portion 12e, side plate portion 12f between 12e is one which is integrally connected, further than the detector 1 described above, The area of the portion in contact with the semiconductor element 11 is set small. The vertical plate portions 12 e and 12 e have a length that covers the left and right side edges of the semiconductor element 11. The left and right vertical plate portions 12e, 12e and the horizontal plate portion 12f are connected to each other by forming a wide triangular portion and reinforcing the widened portion and the vicinity thereof via the conductive adhesive 14. It is configured to be bonded to the semiconductor element 11.

  In the detector 1B configured as described above, similarly to the above, using a 137Cs 662 keV γ ray with a source intensity of 2500 kBq, the detector 1B is installed at a position 200 mm away from the source, and the sensitivity is measured. It was. As a result, the peak count of 662 keV was 10122 times per hour, and an improvement in sensitivity was confirmed compared to the comparative example (see FIG. 3C) and the detector 1 described above.

Next, a PET imaging device 30 configured using the above-described detector 1 will be described.
As shown in FIG. 6, the PET imaging apparatus 30 of the present embodiment supports the imaging apparatus 31 having a columnar measurement space 31 a at the center and a subject (examinee) H and is movable in the longitudinal direction. A bed 32, a data processing device (such as a computer) 33, and a display device 34 are mainly provided.
The imaging device 31 has a large number of unit substrates U (printed circuit boards) arranged in the circumferential direction as shown in FIGS. 7 and 8A and 8B, and the subject H is as shown in FIG. It is placed on the bed 32 and inserted into the measurement space 31a surrounded by the unit substrate U.

  As shown in FIG. 7, a large number of unit substrates U arranged in the imaging device 31 are provided in the imaging device 31 so that the surface on which the detector 1 is installed faces in the depth direction (Z direction) of the imaging device 31. It is installed on an annular support member (not shown) and surrounds the measurement space 31a. The unit substrate U includes a detection module 20A on which a plurality of detectors 1 are installed and an ASIC (Application Specific Integrated Circuit) substrate 20B. The detection module 20A is on the inner side (measurement space 31a side), and the ASIC substrate 20B. Are arranged on the outer side (side away from the measurement space 31a). In the present embodiment, the plurality of unit substrates U are also arranged in the depth direction (Z direction) of the imaging device 31.

  As shown in FIGS. 8A and 8B, a plurality of detectors 1 are installed in the detection module 20A via the protruding portions 12a and 12b (see FIG. 1) of the electrode plates 12C and 12A. In the present embodiment, on the wiring board 24, 6ch in the Y direction (radial direction of the imaging device 31, see FIG. 7) from the detection module 20A toward the ASIC substrate 20B, the X direction (the circumference of the imaging device 31) perpendicular to the Y direction. 16 channels in the direction (see FIG. 7), and further 2 channels (arranged on both surfaces of the wiring board 24) in the Z direction (depth direction of the imaging device 31, see FIG. 7), which is the thickness direction of the wiring board 24. As a result, the detector 1 is installed with a total of 96 ch on one side of the wiring board 24 and a total of 192 ch on both sides.

  The ASIC board 20 </ b> B includes a capacitor 26, a resistor 27, an analog ASIC 28, and a digital ASIC 29. The γ-ray detection signal output from each detector 1 is supplied from the detection module 20A side to the ASIC board 20B side via connectors C1 and C2 (see FIG. 8B).

The ASIC board 20B has one digital ASIC 29 installed on one side and four analog ASICs 28 arranged on both sides. As many capacitors 26 and resistors 27 as the number of detectors 1 are provided on both surfaces of the ASIC substrate 20B.
A plurality of connection wirings (not shown) for electrically connecting the capacitor 26, the resistor 27, the analog ASIC 28, and the digital ASIC 29 are provided in the ASIC substrate 20B. The analog ASIC 28 is a kind of LSI for a specific application that processes an analog signal (γ-ray detection signal) output from the detector 1. The analog ASIC 28 is provided with a signal processing circuit for each detector 1. These signal processing circuits are adapted to obtain a peak value of γ-rays by inputting a γ-ray detection signal (radiation detection signal) output from one corresponding detector 1.

  In the present embodiment, the detector 1, the capacitor 26, the resistor 27, the analog ASIC 28, and the digital ASIC 29 are arranged close to each other in the radial direction of the imaging device 31 so that the accuracy of the detection time of the simultaneous measurement process is not impaired. The circuit length and the length (distance) of the wiring for transmitting the γ-ray detection signal are shortened to reduce the influence of noise on the γ-ray detection signal and to reduce the attenuation of the γ-ray detection signal.

The operation of the PET imaging apparatus 30 having the above configuration will be described. Before performing a radiological examination, first, a radiopharmaceutical for PET (including ¹⁸ F, for example) is administered to the subject H in advance so that its in-vivo radioactivity becomes, for example, about 370 MBq. The radiopharmaceutical is selected according to the examination purpose (for example, grasping the location of cancer or examining an aneurysm of the heart). The administered radiopharmaceutical gathers in the affected area of the subject H before long. In this state, the subject H is laid on the bed 32.
Examiners (medical radiographers and doctors) who perform PET examinations need information according to the purpose of the examination (areas (imaging areas or areas of interest) for which tomographic images are to be obtained, number of slices, slice interval, absorbed dose, etc.) Is input via the data processing device 33. In this case, an information input screen (not shown) is displayed on the display device 34, and necessary data is input using a keyboard, a mouse, or the like. Thereafter, the bed 32 is moved in the longitudinal direction, and the subject H is inserted into the measurement space 31a until the examination site of the subject H (for example, an affected part of cancer) comes to a predetermined position. Then, the PET imaging device 30 is operated.

  In response to an instruction from the data processing device 33, a DC high voltage is applied between the cathode C and the anode A (see FIG. 1) of each detector 1, and the imaging device 31 starts a PET inspection. The detector 1 detects γ rays emitted from the body of the subject H due to the radiopharmaceutical. That is, when the positrons emitted from the radiopharmaceutical for PET are extinguished, a pair of γ rays are emitted in opposite directions of about 180 ° and detected by separate detectors 1. The detector 1 outputs a γ-ray detection signal. The γ-ray detection signal is input from the wiring board 24 side to the corresponding signal processing circuit (not shown) in the corresponding analog ASIC 28 via the connectors C1 and C2 and the capacitor 26. This signal processing circuit amplifies the γ-ray detection signal and obtains a peak value of the detected γ-ray. This peak value is converted into digital peak value information by an analog / digital converter (ADC) (not shown) in the digital ASIC 29. The digital ASIC 29 further outputs position information of the detector 1 that has detected γ rays and detection time information of γ rays. The digital peak value information, the position information of the detector 1 and the γ-ray detection time information are input to the data processing device 33. The simultaneous measurement device (not shown) of the data processing device 33 uses the detection time information to count a pair of γ-rays generated by the disappearance of one positron as one and detect the two γ-rays detected by the pair of γ-rays. The positions of the detectors 1 and 1 are specified based on the position information. Further, a tomogram information creation device (not shown), which is an image information creation device of the data processing device 33, uses the count value obtained by the simultaneous measurement and the position information of the detector 1, that is, the radiopharmaceutical accumulation position, The tomographic image information (image information) of the subject at the malignant tumor position is created. This tomographic image information is displayed on the display device 34.

In the PET imaging apparatus 30 operating in this way, the carrier charge amount is measured using the 137Cs 662 keVγ ray with the source intensity of 2500 kBq, as in the sensitivity measurement of the detector 1 described above, and the sensitivity is measured. went. The radiation source was arranged at the center of the measurement space 31a.
As a comparative example, many detectors (not shown) using an electrode plate EB in which the notch 13 is not formed as shown in FIG. 3C are prepared, and these are provided on the wiring substrate 24 of the PET imaging device 30. Attach and measure the sensitivity as well. In addition, the detector of this comparative example was different only in the electrode plate EB, and used the semiconductor material 11a which has the same volume.

As a result, the PET imaging device 30 using the detector 1 of the present embodiment was 16% more sensitive than the PET imaging device using the detector of the comparative example. That is, an improvement in sensitivity higher than the improvement in sensitivity (12%) when measured with the single detector 1 described above was obtained. This is because the PET imaging apparatus 30 measures the pair annihilation γ-rays simultaneously at two locations. The sensitivity for detecting a pair of γ-rays generated by the annihilation of one positron within a set time window for simultaneous measurement. This is because the square is improved on both sides.
Since the sensitivity improvement of 16% is obtained in this way, in the PET imaging apparatus 30, the inspection time (measurement time) of the subject H can be shortened by the sensitivity improvement, that is, 16%.

Below, the effect acquired in this embodiment is demonstrated.
(1) Since the notch 13 is formed in the region in contact with the semiconductor element 11 in the electrode plate 12C (12A) constituting the laminated structure of the detector 1, the notch formed in the electrode plate 12C (12A) 13 can be used as a passage space for primary electrons E generated by the photoelectric effect caused by the incidence of γ rays. That is, even if the primary electrons E generated by the photoelectric effect jump out toward the electrode plate 12C (12A), the primary electrons E pass through the notch 13 and are adjacent to the semiconductor element 11 (in FIG. 3B, the semiconductor element). 11 ′, the same applies hereinafter), and the probability of being absorbed by the electrode plate 12C (12A) is reduced. Therefore, the primary electrons E that would have been annihilated by the electrode plate 12C (12A) in the prior art can be made incident on the adjacent semiconductor element 11. As a result, the primary electrons E that generate carriers in the adjacent semiconductor element 11 can be increased, and this can be used as an effective signal. Therefore, the sensitivity of γ-ray (radiation) detection is improved.
(2) Since the notch 13 formed in the electrode plate 12C (12A) is formed in a region in contact with the semiconductor element 11, the semiconductor element 11 is preferably used by a portion other than the notch 13 of the electrode plate 12C (12A). Can be held in.
(3) Since the electrode plate 12C (12A) is magnetized, it can be handled using electromagnetic tweezers, etc., and a process or wiring board for stacking the semiconductor element 11 and the electrode plate 12C (12A) Therefore, the detector 1 can be easily handled at the time of mounting, and productivity can be improved and assembly cost can be reduced.
In particular, in the small stack type detector 1 used in the PET imaging apparatus 30 as in the present embodiment, when the area of the electrode plate 12C (12A) becomes small, the electrode plate 12C is used by using vacuum tweezers or the like by suction. Since it becomes difficult to handle (12A), the electrode plate 12C (12A) can be effectively controlled by providing magnetism and the like by giving magnetism to the electrode plate 12C (12A).
(4) In the detector 1A shown in FIG. 4 (a) ~ (c), while provided many notch 13 ', the area of the cut remaining portion of the electrode plate 12C _1, the detector 1 (Fig. since larger than the area of the cut remaining portion of the electrode plate 12C in first reference), the stiffness of the electrode plate 12C ₁ is high, and excellent retention of the semiconductor element 11. Further, cutout portions 13 'and hardly be deformed or the like when punching the electrode plate 12C _1, is excellent in shape retention of the electrode plate 12C _1. Therefore, the detector 1A having excellent sensitivity and high dimensional accuracy can be obtained.
(5) In the detector 1B shown in FIGS. 5A to 5C, the area of the portion in contact with the semiconductor element 11 can be minimized, and the sensitivity can be further improved. The vertical plate portion 12e of the electrode plate 12C _2, 12e, so has a length to cover the right and left side edge portions of the semiconductor element 11, preventing damage or the like due to contact or the like of the left and right side edge portions of the semiconductor element 11 can do.
(6) In the PET imaging device 30 using the detector 1, the sensitivity of the detector 1 is improved, so that the amount of radiopharmaceutical can be reduced and the measurement time can be shortened. Moreover, since the increase in the number of persons to be inspected (number of persons to be measured) can be expected by shortening the inspection time (measurement time), the economic effect is high and the cost can be reduced.
In addition, since the detector 1 configured by laminating a plurality of thin semiconductor elements 11 is used, the moving distance of electrons and holes can be shortened, energy resolution and recognition accuracy of incident time can be improved, and The advantage is that the volume occupancy can be increased and the volume of the detector 1 can be increased, and the performance of the PET imaging apparatus 30 can be improved.

(Second Embodiment)
The semiconductor radiation detector which is 2nd Embodiment of this invention is demonstrated. As shown in FIG. 9, the detector 1 </ b> D of the present embodiment has a configuration in which the electrode plates 12 </ b> C and 12 </ b> A in which the cutout portions 13 are formed are disposed only between the adjacent semiconductor elements 11 and 11. The electrode plate 12B disposed on both sides of the detector 1D is different from that of the face plate in which the notch 13 is not formed.

  According to such a detector 1D, since the electrode plates 12C and 12A in which the cutout portion 13 is formed only between the stacked semiconductor elements 11 and 11 are arranged, as described above, γ rays Is the end of the semiconductor material 11a, for example, in the vicinity of the anode A (see FIG. 3B), the primary electrons E are incident on the adjacent semiconductor element 11 ′ and adjacent to the adjacent semiconductor element 11 ′. Carriers are generated in the semiconductor element 11 ′. As a result, primary electrons E that have been annihilated in the electrode plate 12C (12A) in the past are incident on the adjacent semiconductor element 11 ′, thereby generating primary electrons E that generate carriers in the adjacent semiconductor element 11 ′. Increases and can be used as a valid signal. Therefore, the sensitivity of radiation detection is improved.

  On the other hand, when the incident position of γ rays is in the vicinity of the electrode plate 12B, the primary electrons E may jump out toward the electrode plate 12B as described above. In this case, since the notch 13 is not formed in the electrode plate 12B, the ejected primary electrons E are absorbed in the electrode plate 12B and disappear when entering the electrode plate 12B. Therefore, the primary electrons E are suppressed from jumping out from the electrode plates 12B stacked on both ends of the detector 1D.

  In the detector 1D configured as described above, as described above, a 137Cs 662 keV γ ray with a source intensity of 2500 kBq is used, and the detector 1A is installed at a position 200 mm away from the source to measure sensitivity. It was. As a result, the count of the 662 keV peak was 10089 times per hour, confirming an improvement in sensitivity compared to the comparative example (see FIG. 3C).

In the present embodiment, the effects (1) to (6) produced in the first embodiment can be obtained, and the effects described below are further produced.
(7) The detector 1D of the present embodiment is particularly effective when used in the PET imaging device 30 described above. That is, as described above, in the PET imaging apparatus 30, since it is necessary to provide a large number of detectors 1D on the wiring board 24, the incidence of primary electrons E from the detector 1D to other adjacent detectors 1D is suppressed. Thereby, the detection accuracy of each detector 1D is improved, and the accuracy of energy resolution and position resolution can be improved.

(Third embodiment)
The semiconductor radiation detector which is 3rd Embodiment of this invention is demonstrated. As shown in FIG. 10, the detector 1E of the present embodiment includes a total of two detector assemblies 11A and 11B, in which a plurality of semiconductor elements 11 are configured as one unit by electrical connection in the stacking direction. It has a different point.
The detector 1E includes a semiconductor element 11, electrode plates 12A ₃ and 12A ₄ having notches 13, and an electrode plate 12B having no notches 13, which are alternately stacked, and an electrode plate 12B disposed in the center. Two detector aggregates 11A and 11B are configured with reference to. In the present embodiment, the central electrode plate 12B serves as a common electrode for the two detector assemblies 11A and 11B.

As shown in FIG. 11, each detector assembly 11A, 11B includes two semiconductor elements 11, an electrode plate 12A ₃ (12A ₄ ) disposed between the two semiconductor elements 11, and two sheets It has electrode plates 12B (one electrode plate 12B is shared) disposed on both sides of the semiconductor element 11, and outputs independent γ-ray detection signals (radiation detection signals). That is, the detector assembly 11A side of the two γ-rays incident on the semiconductor element 11, as γ ray detection signal of the detector assembly 11A side, for the anode A from the electrode plates 12A ₃ through the protruding portion 12b ₁ is the output of the connecting member AP1, also, two γ-rays incident on the semiconductor device 11 of the detector assembly 11B side, as γ ray detection signal of the detector assembly 11B side, the protruding portion from the electrode plate 12A ₄ is output to the connecting member AP2 for the anode a via 12b _2.

Here, since the notches 13 are formed in the electrode plates 12A ₃ and 12A ₄ disposed between the two semiconductor elements 11 of the detector assemblies 11A and 11B, the following operation is performed. There is an effect.
That is, the detector assembly 11A will be described as an example. As shown in FIG. 3B, primary electrons E generated by the photoelectric effect due to the incidence of γ rays generate carriers in the semiconductor material 11a. Even if it jumps out toward the adjacent semiconductor element 11 ′ side (adjacent semiconductor element 11 ′ in the detector assembly 11A) before, it passes through the notch 13 as it is, and the semiconductor material of the adjacent semiconductor element 11 ′ It is incident on 11a 'and taken out as an effective electrical signal by the adjacent semiconductor element 11'. Accordingly, carriers due to primary electrons E are generated in the detector assembly 11A, thereby improving the sensitivity in the detector assembly 11A.

  On the other hand, the electrode plate 12B disposed between the detector assembly 11A and the detector assembly 11B is not formed with the notch 13, so that the electrode is similar to the operation in FIG. Even if the primary electrons E generated by the γ rays incident in the vicinity of the plate 12B jump out toward the electrode plate 12B, they are absorbed in the electrode plate 12B and disappear. Therefore, the primary electrons E are prevented from flying between the detector assemblies 11A and 11B. As a result, the γ rays incident on the detector aggregates 11A and 11B individually generate carriers due to the primary electrons E only in the detector aggregates 11A and 11B. Therefore, the detection accuracy in each of the detector aggregates 11A and 11B is substantially improved. This contributes to the improvement of energy decomposition and the accuracy of position resolution in the PET imaging apparatus 30.

In the present embodiment, the effects (1) to (6) produced in the first embodiment can be obtained, and the effects described below are further produced.
(8) In the present embodiment, the electrode plates 12A ₃ , 12A _{4 in} which the notch portions 13 are formed and the electrode plates 12B in which the notch portions 13 are not formed are alternately arranged, so that 2 One detector assembly 11A, 11B can be configured, and a detector 1E that can output a separate γ-ray detection signal from each detector assembly 11A, 11B can be configured.
According to the PET imaging apparatus 30 using such a detector 1E, the PET imaging apparatus 30 is excellent in energy resolution and position resolution with an inexpensive configuration in which the lamination of the electrode plates 12A and the electrode plates 12B is alternated. Can be obtained.
In each of the above embodiments, the semiconductor material 11a using CdTe has been described. However, the present invention is not limited to this, and various substances can be used. Further, the number of stacked semiconductor elements 11 can be set arbitrarily.
Further, a thin film portion having a thickness smaller than that of other portions is formed on the electrode plates 12C and 12A, and primary electrons E generated by the photoelectric effect are incident on the adjacent semiconductor elements 11 and 11 ′ through the thin film portion. You may comprise. The thin film portion may be formed together with the cutout portions 13 and 13 ′, or may be formed in place of the cutout portions 13 and 13 ′.

It is the perspective view which showed typically the semiconductor radiation detector of 1st Embodiment of this invention. (A) is a schematic perspective view of a semiconductor element, (b) is a perspective view schematically showing an electrode plate, and (c) is an exploded perspective view of a semiconductor radiation detector. (A)-(c) is a schematic diagram which shows the production | generation of the carrier by the incidence | injection of a gamma ray. (A) is the perspective view which showed typically the semiconductor radiation detector of the modification, (b) is the perspective view which showed the electrode plate typically, (c) is the disassembled perspective view of a semiconductor radiation detector. (A) is a perspective view schematically showing a semiconductor radiation detector according to another modification, (b) is a perspective view schematically showing an electrode plate, and (c) is an exploded perspective view of the semiconductor radiation detector. is there. 1 is a configuration diagram illustrating a positron emission imaging apparatus according to a first embodiment of the present invention. FIG. It is a schematic diagram which shows arrangement | positioning of the unit board | substrate in an imaging device. (A) is a front view of a unit board | substrate, (b) is a side view of a unit board | substrate. It is a perspective view which shows typically the semiconductor radiation detector of 2nd Embodiment of this invention. It is a perspective view which shows typically the semiconductor radiation detector of 3rd Embodiment of this invention. It is a disassembled perspective view of a semiconductor radiation detector similarly.

Explanation of symbols

1 Semiconductor radiation detector (detector)
1A, 1B, 1D, 1E Semiconductor radiation detector (detector)
11, 11 ′ Semiconductor element 11A, 11B Detector assembly 11a, 11a ′ Semiconductor material 12C, 12A Electrode plate 12B Electrode plate 12A ₁ electrode plate 12A ₂ electrode plate 12C ₁ electrode plate 12C ₂ electrode plate 12a Protruding portion 12b Protruding portion 13 Cutout portion 13 'Cutout portion 13''Cutout portion 14 Conductive adhesive 20A Detection module 20B ASIC substrate 24 Wiring substrate 30 PET imaging device A Anode electrode C Cathode electrode U Unit substrate

Claims (8)

A semiconductor radiation detector having a laminated structure in which a plurality of semiconductor elements that convert incident radiation into electrical signals and metal electrode plates are alternately bonded with a conductive adhesive,
A semiconductor radiation detector, wherein at least one of the electrode plates constituting the laminated structure has a notch formed in a region in contact with the semiconductor element.   The semiconductor radiation detector according to claim 1, wherein the electrode plate in which the notch is formed is disposed only between the adjacent semiconductor elements stacked. A semiconductor element that converts incident radiation into an electrical signal and a plurality of metal electrode plates are alternately bonded with a conductive adhesive to form a stacked structure, and the plurality of semiconductor elements are electrically connected in the stacking direction. Is a semiconductor radiation detector comprising a plurality of detector assemblies configured in a stacking direction,
A semiconductor having a notch formed in a region in contact with the semiconductor element in at least one of the other electrode plates excluding the electrode plate disposed between adjacent detector assemblies. Radiation detector.   The semiconductor radiation detector according to claim 3, wherein the electrode plate having the notch is disposed only between the adjacent semiconductor elements stacked in the detector assembly.   The semiconductor radiation detector according to claim 1, wherein the electrode plate has magnetism.   6. The semiconductor radiation detector according to claim 1, wherein a portion of the electrode plate in which the cutout portion is formed is in a frame shape in contact with the semiconductor element. .   The electrode plate in which the notch is formed has a portion in contact with the semiconductor element formed linearly along at least the outer shape of the semiconductor element. The semiconductor radiation detector according to item 1. A positron emission tomography apparatus using the semiconductor radiation detector according to any one of claims 1 to 7,
A wiring board having a plurality of the semiconductor radiation detectors attached thereto, surrounding a measurement region into which a bed supporting a subject is inserted, and a plurality of printed circuit boards arranged around the measurement region;
A positron emission tomography apparatus comprising: an image information creation device that generates an image using information obtained based on radiation detection signals output from the plurality of semiconductor radiation detectors.

Images