Radiography device and radiography system

Abstract

This radiography device comprises a scintillator unit that converts incident radiation into visible light and a photoelectric conversion unit that converts the visible light into an electric signal, and is characterized in that the scintillator unit includes a first visible light absorbing member that separates a plurality of scintillators included in the scintillator unit in a direction perpendicular to the direction of incidence of the radiation and a second visible light absorbing member provided on a surface where the radiation is incident.

Description

Radiation imaging device and radiation imaging system

[0001] The present invention relates to a radiation imaging apparatus and a radiation imaging system.

[0002] As a radiographic imaging device that captures a radiographic image using radiation (such as X-rays) that has passed through a subject, an FPD (Flat Panel Detector) capable of displaying a radiographic image in real time has been proposed. The FPD has minute radiation detectors arranged in a matrix on a quartz glass substrate, each of which is a stack of a solid-state photodetector sandwiched between a transparent conductive film and a conductive film, and a scintillator that converts radiation into visible light. Known solid-state photodetectors include those that use photodetectors such as a CCD (Charge-Coupled Device) and a CMOS (Complementary Metal-Oxide Semiconductor). Other known solid-state photodetectors include those that use photodetectors such as an avalanche photodiode (APD) and a single-photon avalanche diode (SPAD).

[0003] Patent Document 1 discloses an X-ray CT device using an indirect photon-counting sensor. Patent Document 1 also discloses a configuration in which a scintillator is divided into pixels. This configuration can prevent visible light generated by the scintillator from entering adjacent pixels, thereby preventing degradation of image quality due to pile-up (when the intensity of visible light entering a pixel is too high and the number of times visible light is detected cannot be counted correctly).

[0004] Furthermore, the photodetector described in Patent Document 2 has a plurality of sub-pixels arranged in each pixel, and detects visible light photons emitted by a scintillator.

[0005] JP 2019-86443 A JP 2016-162772 A

[0006] However, in the photodetectors described in any of the patent documents, there is room for improvement in terms of measures to prevent degradation of image quality due to pile-up.

[0007] The present invention, which solves the above-mentioned problems, is a radiation imaging device having a scintillator section that converts incident radiation into visible light and a photoelectric conversion section that converts the visible light into an electrical signal, wherein the scintillator section has a first visible light absorbing member that separates multiple scintillators included in the scintillator section in a direction perpendicular to the incident direction of the radiation, and a second visible light absorbing member provided on the radiation incident surface.

[0008] Another invention is a radiation imaging device comprising a plurality of sub-pixels, each of which has a scintillator that generates visible light in response to irradiation with radiation, a photoelectric conversion element that generates electric charges in response to the visible light generated by the scintillator, and a sub-signal processing unit that generates electric signals in response to the charges generated by the photoelectric conversion element, and a signal processing unit that processes the electric signals generated by the sub-signal processing units included in the plurality of sub-pixels to generate pixel signals, wherein the visible light receiving area of ​​the photoelectric conversion element differs among the plurality of sub-pixels.

[0009] Another invention is a radiation imaging device comprising a plurality of sub-pixels, each of which has a scintillator that generates visible light in response to irradiation with radiation, a photoelectric conversion element that generates electric charges in response to the visible light generated by the scintillator, and a sub-signal processing unit that generates electric signals in response to the charges generated by the photoelectric conversion element, and a signal processing unit that processes the electric signals generated by the sub-signal processing units included in the plurality of sub-pixels to generate pixel signals, wherein the processing capability of the sub-signal processing units to process the electric charges differs among the plurality of sub-pixels.

[0010] Another invention is a radiation imaging device comprising: a scintillator that generates visible light in response to irradiation with radiation; a photoelectric conversion unit that receives the visible light generated by the scintillator and generates an electrical signal; and a signal processing unit that classifies and outputs the electrical signal generated by the photoelectric conversion unit, wherein the photoelectric conversion unit comprises a plurality of sub-pixels, each of which includes a photoelectric conversion element that receives the visible light and generates an electric charge, and a sub-signal processing unit that receives the electric charge generated by the photoelectric conversion element and generates the electric signal, and the plurality of sub-pixels are arranged so that the visible light detection sensitivity per unit area of ​​the photoelectric conversion unit is different from one another.

[0011] According to the present invention, pile-up in sub-pixels arranged within a pixel can be reduced, and visible light photons can be measured with high accuracy.

[0012] 1 is a schematic diagram of a radiation imaging device according to a first embodiment; sectional view of a radiation imaging device according to a first embodiment; sectional view of a radiation imaging device according to a second embodiment; sectional view of a radiation imaging device according to a second embodiment; sectional view of a radiation imaging device according to a third embodiment; sectional view of a radiation imaging device according to a third embodiment; sectional view of a radiation imaging device according to a fourth embodiment; sectional view of a radiation detection panel according to a fourth embodiment; sectional view of a photoelectric conversion substrate of a radiation detection panel according to a fourth embodiment; sectional view of a signal processing substrate of a radiation detection panel according to a fourth embodiment; sectional view of a subpixel circuit of a radiation detection panel according to a fourth embodiment; sectional view of a pixel circuit of a radiation detection panel according to a fourth embodiment; sectional view of a pixel circuit of a radiation detection panel according to a fourth embodiment; Schematic diagram showing driving; Configuration example of a photoelectric conversion substrate of a radiation detection panel according to a fourth embodiment; Plan view of a radiation detection panel according to a fourth embodiment; Diagonal cross-sectional view of a radiation detection panel according to a fourth embodiment; Cross-sectional view of a radiation detection panel according to a fourth embodiment; Cross-sectional view of a conventional radiation detection panel; Cross-sectional view of a radiation detection panel according to a first embodiment of the fourth embodiment; Cross-sectional view of a radiation detection panel according to a second embodiment of the fourth embodiment; Cross-sectional view of a radiation detection panel according to a third embodiment of the fourth embodiment; Cross-sectional view of a radiation detection panel according to a fourth embodiment; Configuration example of a radiation imaging system according to the present disclosure; Partial top view of a radiation detection panel according to a first embodiment of the fourth embodiment; Partial top view of a radiation detection panel according to a second embodiment of the fourth embodiment; Schematic diagram explaining an application example of a radiation imaging device according to the present disclosure to an X-ray diagnostic system.

[0013] Hereinafter, embodiments will be described in detail with reference to the accompanying drawings. Note that the following embodiments do not limit the scope of the invention claimed. Although multiple features are described in the embodiments, not all of these multiple features are necessarily essential to the invention, and multiple features may be combined arbitrarily. Furthermore, in the accompanying drawings, the same reference numerals are used to designate the same or similar components, and redundant explanations will be omitted.

[0014] In the following description, terms indicating specific directions or positions (for example, "upper," "lower," "right," "left," and other terms including these terms) are used as necessary. The use of these terms is for the purpose of facilitating understanding of the embodiments with reference to the drawings, and the meanings of these terms do not limit the technical scope of the present invention.

[0015] In this specification, a planar view refers to a view from a direction perpendicular to the light incident surface of the semiconductor layer. A cross-sectional view refers to a surface perpendicular to the light incident surface of the semiconductor layer. When the light incident surface of the semiconductor layer is a rough surface when viewed microscopically, the planar view is defined based on the light incident surface of the semiconductor layer when viewed macroscopically.

[0016] The semiconductor layer has a first surface and a second surface opposite to the first surface, through which light is incident. In this specification, the depth direction is the direction from the first surface toward the second surface of the semiconductor layer on which the APD is disposed. Hereinafter, the "first surface" may be referred to as the "front surface," and the "second surface" may be referred to as the "rear surface."

[0017] In the following description, the anode of the avalanche photodiode (APD) is set to a fixed potential, and a signal is extracted from the cathode side. Therefore, the first conductivity type semiconductor region having majority carriers of charges of the same polarity as the signal charge is an N-type semiconductor region, and the second conductivity type semiconductor region having majority carriers of charges of a polarity opposite to that of the signal charge is a P-type semiconductor region. Note that the present invention also applies when the cathode of the APD is set to a fixed potential, and a signal is extracted from the anode side. In this case, the first conductivity type semiconductor region having majority carriers of charges of the same polarity as the signal charge is a P-type semiconductor region, and the second conductivity type semiconductor region having majority carriers of charges of a polarity opposite to that of the signal charge is an N-type semiconductor region. The following description will be given of a case where one node of the APD is set to a fixed potential, but the potentials of both nodes may fluctuate.

[0018] In this specification, when the term "impurity concentration" is used simply, it means the net impurity concentration minus the amount compensated for by impurities of the opposite conductivity type. A region where the concentration of P-type doped impurities is higher than the concentration of N-type doped impurities is a P-type semiconductor region. Conversely, a region where the concentration of N-type doped impurities is higher than the concentration of P-type doped impurities is an N-type semiconductor region.

[0019] 22 shows an example of the basic configuration of a radiation imaging system 980 according to an embodiment of the present invention. The radiation imaging system 980 is configured to generate an electrical radiation image by electrically capturing an optical image formed by radiation. The radiation is typically X-rays, but may also be α-rays, β-rays, γ-rays, etc. The radiation imaging system 980 includes, for example, a radiation imaging device 990 and a control device 983 consisting of a computer that acquires image data from the radiation imaging device 990 and processes the acquired image data. In a preferred embodiment of the present invention, the radiation imaging system 980 further includes a display 982, an exposure control device 981, and a radiation generation device 984.

[0020] The radiation generating device 984 starts emitting radiation 985 in accordance with an exposure command (radiation command) from the exposure control device 981. The radiation 985 irradiated from the radiation generating device 984 passes through the subject 986 and enters the radiation imaging device 990. The radiation generating device 984 also stops emitting the radiation 985 in accordance with a stop command from the exposure control device 981.

[0021] The radiation imaging device 990 includes a radiation detection panel 991, a control circuit 993, and an image generation circuit 992. The radiation detection panel 991 generates a radiation image according to radiation 985 incident on the radiation imaging device 990 and transmits the generated radiation image to the computer 983. The control circuit 993 controls the operation of the radiation detection panel 991. For example, the control circuit 993 generates a stop signal for stopping the radiation generation device 984 from emitting radiation 985, based on an image signal obtained from the radiation detection panel 991. The stop signal is supplied to the exposure control device 981. In response to the stop signal, the exposure control device 981 sends a stop command to the radiation generation device 984.

[0022] The control circuit 993 may be configured with a dedicated circuit such as a PLD (Programmable Logic Device) such as an FPGA (Field Programmable Gate Array), or may be configured with a dedicated circuit such as an ASIC (Application Specific Integrated Circuit). Alternatively, the control circuit 993 may be configured with a combination of a general-purpose processing circuit such as a processor and a storage circuit such as a memory. In this case, the function of the control circuit 993 may be realized by the general-purpose processing circuit executing a program stored in the storage circuit.

[0023] The image generation circuit 992 stores the signal supplied from the radiation detection panel 991 in a memory, and generates a radiographic image based on this signal. The method of generating a radiographic image will be described in detail later. The image generation circuit 992 transmits the generated radiographic image to the computer 983.

[0024] The computer 983 has a control unit that controls the radiation imaging device 990 and the exposure control device 981, a receiving unit that receives radiation images from the radiation imaging device 990, and a signal processing unit that processes radiation images obtained by the radiation imaging device 990. Like the control circuit 993, the control unit, the receiving unit, and the signal processing unit may each be configured with a dedicated circuit or a combination of a general-purpose processing circuit and a memory circuit. In one example, the exposure control device 981 has an exposure switch, and when the exposure switch is turned on by a user, it sends an exposure command to the radiation generation device 984 and sends a start notification indicating the start of radiation irradiation to the computer 983. In response to the start notification, the computer 983 notifies the control circuit 993 of the radiation imaging device 990 of the start of radiation irradiation. When the exposure control device 981 and the computer 983 are not synchronously connected, the radiation detection panel 991 may detect the start of irradiation of radiation 985 based on pixel signals.

[0025] First Embodiment An example of the schematic configuration of a radiation imaging apparatus according to a first embodiment of the present disclosure will be described with reference to Fig. 1 and Fig. 2. Fig. 1 is a schematic configuration diagram of the radiation imaging apparatus, and Fig. 2 is a cross-sectional view.

[0026] As shown in FIGS. 1 and 2 , the radiation detection panel 1100 (radiation detection panel 991 in FIG. 22 ) includes a scintillator section 1200 and a photoelectric conversion section 1300. The scintillator section 1200 is also referred to as a scintillator panel. X-ray photons 1400 are incident on the scintillator section 1200 and are converted and amplified into visible light photons 1500 by the scintillator section 1200. The surface on which the X-ray photons 1400 are incident on the scintillator section 1200 is also referred to as the radiation incident surface. The visible light photons 1500 are incident on the photoelectric conversion section 1300 and detected as an electrical signal. The photoelectric conversion section 1300 includes a photoelectric conversion substrate (hereinafter sometimes referred to as a photodiode substrate) 1310 and a signal processing substrate (hereinafter sometimes referred to as a counter circuit substrate) 1320.

[0027] The photodiode substrate 1310 and the counter circuit substrate 1320 are stacked and electrically connected. The photodiode substrate 1310 has a pixel region in which pixels 1301 having a sensor function are arranged in a matrix. Therefore, the photodiode substrate 1310 may be called a sensor member, a sensor substrate, a sensor chip, etc. The counter circuit substrate 1320 has a circuit region that processes signals detected in the pixel region. The counter circuit substrate 1320 may be called a circuit member, a circuit board, a circuit chip, etc. Furthermore, both substrates are semiconductor substrates made of silicon wafers.

[0028] The photodiode substrate 1310 has a first semiconductor layer 1311 and a first wiring structure 1312 that form a photodiode, which is a photoelectric conversion element, and these constitute the pixel 1301. As will be described later, since the scintillator section 1200 of this embodiment emits a small amount of light, an avalanche photodiode that has a function of amplifying a signal is suitable as the photoelectric conversion element. The counter circuit substrate 1320 has a second semiconductor layer 1321 and a second wiring structure 1322 that constitute circuits such as a signal processing section, corresponding to the pixel 1301.

[0029] The photoelectric conversion unit 1300 is a back-illuminated photoelectric conversion unit. That is, visible light photons 1500 are incident on a second surface of the photodiode substrate 1310, which is opposite to a first surface on which the first wiring structure 1312 is formed, and a counter circuit substrate 1320 is disposed on the first surface side of the photodiode substrate 1310. Note that the inventions according to each embodiment can also be applied to a photoelectric conversion unit having a front-illuminated structure.

[0030] The scintillator section 1200 and the photoelectric conversion section 1300 are stacked with an adhesive member 1201 interposed therebetween, as shown in FIG. 2 . The adhesive member 1201 can be an adhesive member that melts or softens when heated. The adhesive member 1201 can be, for example, a sheet-like or liquid adhesive material (also called a hot-melt resin) containing a thermoplastic elastomer such as a styrene-based, olefin-based, PVC-based, urethane-based, or amide-based material. The adhesive member 1201 can also be an acrylic-based, silicone-based, or other adhesive sheet that has adhesive properties at room temperature.

[0031] The scintillator section 1200 has a plurality of scintillators 1210 separated by partition walls (hereinafter sometimes referred to as separation regions) 1220. The plurality of scintillators 1210 are separated in a direction perpendicular to the incident direction of radiation (the incident direction of X-ray photons 1400). A support substrate 1230 is disposed on the X-ray incident side of the scintillator section 1200. Both the separation regions (partition walls) 1220 and the support substrate 1230 are visible light absorbing materials that have a light absorbing function. The partition walls (separation regions) 1220 and the support substrate 1230 can suppress the spread of visible light photons generated in the scintillators 1210. The partition walls (separation regions) 1220 and the support substrate 1230 can also reduce the amount of visible light photons generated in the scintillators 1210. Furthermore, by adjusting the size of the partition (separation region) 1220, it is possible to adjust the amount (volume) of the scintillator 1210, and therefore the amount of visible light photons generated in the scintillator 1210. The visible light absorbing material used as the partition (separation region) 1220 is an example of a first visible light absorbing material. The visible light absorbing material used as the support substrate 1230 is an example of a second visible light absorbing material.

[0032] An example of a manufacturing method for the scintillator section 1200 will be described. First, a support substrate 1230 made of, for example, glass, amorphous carbon, CFRP, a resin film, aluminum, or titanium is prepared. A photosensitive polyimide resin, for example, is applied thereon, and separation regions (partition walls) 1220 are formed by photolithography. Furthermore, a black resin or metal film is formed on the surface of the partition walls (separation regions) 1220. Then, particulate Gd2O2S (GOS), which is the scintillator 1210, is filled into the compartments formed by the partition walls (separation regions) 1220.

[0033] As described above, pileup may occur when the amount of visible light photons 1500 generated in the scintillator 1210 exceeds the processing capacity of the photoelectric conversion unit 1300, or when the visible light photons 1500 are diffused in the scintillator unit 1200. Improvement is particularly necessary in FPDs that require large areas and high image sharpness.

[0034] The scintillator 1210 of the radiation imaging device of this embodiment is separated by a partition (separation region) 1220 having a light-absorbing function and a support substrate 1230, thereby suppressing the diffusion of light to adjacent pixels and reducing the amount of generated light. Furthermore, increasing the size of the partition (separation region) 1220 reduces the amount (volume) of phosphor, thereby reducing the amount of visible light photons generated in the scintillator 1210. As a result, the amount of light incident on the photoelectric conversion unit 1300 can be reduced, thereby suppressing the occurrence of pile-up. Furthermore, since photolithography is used to form the separation region 1220, it is easy to increase the area and reduce manufacturing costs.

[0035] As described above, the radiation imaging apparatus of the present disclosure can provide a high-image-quality radiation imaging apparatus that can reduce the occurrence of pile-up by arranging a visible light absorbing member on the radiation incident side of the photoelectric conversion unit 1300.

[0036] Second Embodiment An example of the schematic configuration of a radiation imaging apparatus according to a second embodiment of the present disclosure will be described with reference to Fig. 3 and Fig. 4. Fig. 3 is a schematic configuration diagram of the radiation imaging apparatus, and Fig. 4 is a cross-sectional view. The operating principle of the radiation imaging apparatus is the same as that of the first embodiment, and therefore a description thereof will be omitted. Differences from the first embodiment will be described in detail below.

[0037] 3 and 4 , the radiation detection panel 1100 includes a scintillator section 1200 and a photoelectric conversion section 1300. The photoelectric conversion section 1300 is composed of a single photoelectric conversion substrate 1330, which has semiconductor layers constituting photoelectric conversion elements and circuits such as a signal processing section on the same substrate. That is, a pixel 1302 has a first semiconductor layer 1311, a first wiring structure 1312, a second semiconductor layer 1321, and a second wiring structure 1322. Because one pixel is formed by the photoelectric conversion element and the signal processing circuit, the photoelectric conversion element of this embodiment has a smaller area than the photoelectric conversion element of the first embodiment. For this reason, an avalanche photodiode having a function of amplifying a signal is suitable for the photoelectric conversion element.

[0038] 4, each semiconductor layer is arranged to correspond to a separate scintillator 1210, but the second semiconductor layer 1321, which does not contribute to photoelectric conversion, does not necessarily need to be arranged within the compartment of the scintillator 1210, and may, for example, overlap with the separation region (partition) 1220. Furthermore, the scintillator section 1200 and the photoelectric conversion section 1300 may be overlapped with a shift so that the second semiconductor layer 1321 corresponds to an adjacent compartment.

[0039] An example of a manufacturing method for the scintillator section 1200 will be described. First, a material such as glass or silicon that will become the partition wall (separation region) 1220 is prepared. A photosensitive photoresist, for example, is applied thereon, and a mask pattern is formed using photolithography. Next, unnecessary portions are removed using dry etching, and the photoresist is then peeled off to form the partition wall (separation region) 1220. Powdered CsI:Tl raw material, which is the scintillator 1210, is then melted and filled into the compartments formed by the partition wall (separation region) 1220. From the standpoints of light absorption and processability, a silicon wafer is preferably used as the material for the partition wall (separation region) 1220, which is a visible light absorbing member. Furthermore, by using silicon, the same material as the semiconductor substrate that constitutes the photoelectric conversion section 1300, for the material of the partition wall (separation region) 1220, misalignment due to thermal expansion or the like can be suppressed when bonding the scintillator section 1200 and the photoelectric conversion section 1300 together.

[0040] As mentioned above, pile-up is a significant issue in photon-counting FPDs.

[0041] The scintillator 1210 of the radiation imaging device of this embodiment is separated by partitions (separation regions) 1220 having a light absorbing function, and therefore, like the scintillator 1210 of the first embodiment, it is possible to suppress the diffusion of light to adjacent pixels. It is also possible to reduce the amount of generated light. That is, by reducing the amount of light incident on the photoelectric conversion unit, it is possible to suppress the occurrence of pile-up. Furthermore, since photolithography technology is used to form the partitions (separation regions) 1220, it is easy to increase the area, and it is also possible to suppress manufacturing costs.

[0042] Furthermore, the support substrate 1230 and the partition (separation region) 1220 may be integrally formed of the same material (e.g., silicon, which is a visible light absorbing material). Forming the support substrate 1230 and the partition (separation region) 220 integrally from silicon has the effect of eliminating the need for the support substrate 1230, such as glass.

[0043] As described above, the radiation imaging apparatus of the present disclosure can provide a high-image-quality radiation imaging apparatus that can reduce the occurrence of pile-up by arranging a visible light absorbing member on the radiation incident side of the photoelectric conversion unit 1300.

[0044] [Third Embodiment] An example of the schematic configuration of a radiation imaging apparatus according to a third embodiment of the present disclosure will be described with reference to Fig. 5 and Fig. 6. Fig. 5 is a schematic configuration diagram of the radiation imaging apparatus, and Fig. 6 is a cross-sectional view. The operating principle of the radiation imaging apparatus is the same as that of the first embodiment, and therefore a description thereof will be omitted. Differences from the above embodiments will be described in detail below.

[0045] 5 and 6, the radiation detection panel 1100 includes a scintillator section 1200, a photoelectric conversion section 1300, and a visible light absorbing section 1600. The photoelectric conversion section 1300 and the visible light absorbing section 1600 are joined with an adhesive member 1601. Similar to the second embodiment, the photoelectric conversion section 1300 is configured from a photoelectric conversion substrate 1330 on which a semiconductor layer having photoelectric conversion elements and circuits such as a signal processing section is formed. The visible light absorbing member used in the visible light absorbing section 1600 is an example of a third visible light absorbing member.

[0046] As mentioned above, pile-up is a significant issue in photon-counting FPDs.

[0047] Since the scintillator 1210 has the same structure as that of the second embodiment, the radiation imaging apparatus of this embodiment can also reduce manufacturing costs and the occurrence of pile-up.

[0048] In this embodiment, a visible light absorbing section 1600 is further disposed between the scintillator section 1200 and the photoelectric conversion section 1300. The visible light absorbing section 1600 is preferably made of a material that emphasizes the linearity of visible light, such as a fiber optic plate (FOP). The FOP has the function of reflecting obliquely incident light by controlling the refractive index of the material and suppressing the incidence of visible light into the FOP. Furthermore, by disposing light absorbing members at predetermined intervals within the FOP, the FOP has the function of absorbing oblique light within the FOP. By using such a visible light absorbing member, the amount of visible light incident on the photoelectric conversion substrate can be reduced, thereby suppressing the occurrence of pileup.

[0049] The photoelectric conversion unit 1300 also includes a light intensity adjustment layer 1313 disposed on the X-ray incident side of the first semiconductor layer 1311, which forms the avalanche photodiode. The light intensity adjustment layer 1313 is positioned to reduce visible light incident on the avalanche photodiode. In other words, the light intensity adjustment layer 1313 can serve as a light-shielding pattern for reducing light incident on the avalanche photodiode, which is a photoelectric conversion element. The light intensity adjustment layer 1313, which is a visible light adjustment member, is preferably made of a material capable of absorbing visible light, such as a metal film or a black organic film. The light intensity adjustment layer 1313 further reduces the amount of light transmitted through the FOP, which is the visible light absorbing unit 1600, and the remaining light enters the first semiconductor layer 1311. This prevents pileup. The opening area of ​​the light intensity adjustment layer 1313 is preferably adjusted according to the processing capacity of the photoelectric conversion unit 1300. The visible light adjusting member used as the light amount adjusting layer 1313 is an example of a fourth visible light absorbing member.

[0050] As described above, the radiation imaging apparatus of the present disclosure can provide a high-image-quality radiation imaging apparatus that can reduce the occurrence of pile-up by arranging a visible light absorbing member on the radiation incident side of the photoelectric conversion unit 1300.

[0051] Next, the fourth embodiment of the present invention will be described with reference to several examples. First, the configuration common to each example of this embodiment will be described with reference to FIGS.

[0052] 7 is a diagram showing the configuration of a radiation detection panel 991 in which a scintillator section 500 is further laminated on a laminated photoelectric conversion section 100. Radiation photons 600 are incident on the scintillator section 500, and are converted into visible light photons 700 and multiplied by the scintillator section 500. Generally, the scintillator section 500 multiplies one radiation photon 600 by approximately 1,000 times, and emits it as several thousand visible light photons.

[0053] The multiplied visible light photons 700 are incident on the photoelectric conversion unit 100 and are detected as an electrical signal by the photoelectric conversion unit 100 .

[0054] The photoelectric conversion unit 100 is configured by stacking and electrically connecting two components: a photoelectric conversion substrate 11 (corresponding to the photodiode substrate in Embodiments 1 to 3) and a signal processing substrate 21 (corresponding to the counter circuit substrate in Embodiments 1 to 3). The photoelectric conversion substrate 11 has a first semiconductor layer having a photoelectric conversion element 102 (described later), and a first wiring structure. The signal processing substrate 21 has a second semiconductor layer having circuits such as a signal processing unit 103 (described later), and a second wiring structure. The photoelectric conversion unit 100 is configured by stacking the second semiconductor layer, the second wiring structure, the first wiring structure, and the first semiconductor layer in this order. The photoelectric conversion unit 100 is a back-illuminated photoelectric conversion unit in which visible light photons 700 are incident from the second surface and the signal processing substrate 21 is disposed on the first surface side. Note that the inventions according to each embodiment and example can also be applied to photoelectric conversion units having a front-illuminated structure, and can also be applied to photoelectric conversion units consisting of a single layer rather than a laminated type.

[0055] FIG. 8 is a diagram showing an example of the arrangement of the photoelectric conversion substrate 11. Subpixels 101 each having a photoelectric conversion element 102 including an avalanche photodiode (APD) are arranged in five rows and five columns to form a two-dimensional array of pixels 901 (corresponding to the pixel 1301 in the first to third embodiments) in a plan view. This two-dimensional array of pixels 901 forms a pixel region 12. The left-right direction in FIG. 8 is referred to as the "row direction," "horizontal direction," "x direction," etc., and the up-down direction in FIG. 8 is referred to as the "column direction," "vertical direction," "y direction," etc. Furthermore, directions perpendicular to the paper surface of FIG. 8 are referred to as the "depth direction," "z direction," etc. Furthermore, the pixel 901 constitutes one pixel, which is the smallest unit for generating a radiographic image, and may be simply referred to as the pixel 901 in the following description. Furthermore, a subpixel 101 is a divided region obtained by dividing the pixel 901 into multiple regions.

[0056] 9 is a configuration diagram of the signal processing board 21. It has a sub-signal processing unit 103 that processes the charges photoelectrically converted by the photoelectric conversion element 102 in FIG. 8, and a signal processing unit 902 that combines the signals from the sub-signal processing unit 103 into a signal for each pixel. It also has a readout circuit 112, a control pulse generation unit 115, a horizontal scanning circuit unit 111, signal lines 113, a vertical scanning circuit unit 110, and drive lines 116.

[0057] The photoelectric conversion element 102 in FIG. 8 and the sub-signal processing unit 103 in FIG. 9 are electrically connected via connection wiring provided for each sub-pixel.

[0058] The vertical scanning circuit unit 110 receives a control pulse supplied from a control pulse generating unit 115 and supplies the control pulse to each pixel via a driving line 116. The vertical scanning circuit unit 110 uses logic circuits such as a shift register and an address decoder.

[0059] The signal output from the photoelectric conversion element 102 of the sub-pixel is processed by the sub-signal processing unit 103 .

[0060] The sub-signal processing unit 103 is provided with a counter, a memory, etc., and a visible light photon count value is held in the memory. The visible light photon count value is reset once every time a radiation photon is detected.

[0061] The signal processing unit 902 reads out and integrates the visible light photon count values ​​of the 5 rows and 5 columns of sub-signal processing units 103 to estimate the energy of the radiation photons 600. For example, if the total of the visible light photon count values ​​of the 5 rows and 5 columns is 100 counts, it is recognized that 40 kV radiation photons 600 have been irradiated, and the 30 kV to 60 kV radiation photon count value is incremented by 1. If the total of the visible light photon count values ​​of the 5 rows and 5 columns is 200 counts, it is recognized that 80 kV radiation photons have been irradiated, and the 60 to 90 kV radiation photon count value is incremented by 1. The visible light photon count value varies depending on the configuration of the scintillator and reflective layer, and is therefore not limited to 100 counts or 200 counts.

[0062] The horizontal scanning circuit section 111 inputs a control pulse to the signal processing section 902 for sequentially selecting each column in order to read out a signal from the memory of each pixel in which the radiation photon count value is held.

[0063] A signal is output to the signal line 113 from the signal processing unit 902 of the pixel selected by the vertical scanning circuit unit 110 for the selected column.

[0064] The signal output to the signal line 113 is output via the output circuit 114 to a recording unit or image generation circuit 992 external to the photoelectric conversion unit 100. Therefore, the generated image is a digital value obtained by estimating and counting the energy of radiation photons from the visible light photon count values ​​output from the multiple sub-pixels. For example, if the output value of one pixel is divided into three groups: 30 to 60 kV radiation photons, 60 to 90 kV radiation photons, and 90 to 120 kV radiation photons, and each group is represented by an 8-bit counter, then one pixel will be composed of a 24-bit digital value.

[0065] 8, the photoelectric conversion elements 102 may be arranged one-dimensionally in the pixel region. The function of the signal processing unit does not necessarily need to be provided for each photoelectric conversion element. For example, one signal processing unit may be shared by multiple photoelectric conversion elements 102, and signal processing may be performed sequentially.

[0066] 8 and 9 , a plurality of signal processing units 902 are arranged in a region overlapping the pixel region 12 in a planar view. Then, in a planar view, a vertical scanning circuit unit 110, a horizontal scanning circuit unit 111, a readout circuit 112, an output circuit 114, and a control pulse generation unit 115 are arranged so as to overlap between an end of the photoelectric conversion substrate 11 and an end of the pixel region 12. In other words, the photoelectric conversion substrate 11 has the pixel region 12 and a non-pixel region arranged around the pixel region 12. Then, in a region overlapping the non-pixel region in a planar view, the vertical scanning circuit unit 110, the horizontal scanning circuit unit 111, the readout circuit 112, the output circuit 114, and the control pulse generation unit 115 are arranged.

[0067] Furthermore, a vertical scanning circuit unit 110, a horizontal scanning circuit unit 111, a readout circuit 112, an output circuit 114, and a control pulse generation unit 115 may be arranged within the pixel region 12. In this case, the photoelectric conversion substrate 11 and the signal processing substrate 21 may be tiled in the x and y directions to increase the area.

[0068] Furthermore, a plurality of photoelectric conversion substrates 11 and signal processing substrates 21 can be stacked in the z direction on one wiring substrate to increase the area.

[0069] Fig. 10 is an example of a block diagram including the equivalent circuits of Fig. 8 and Fig. 9. In Fig. 10, the photoelectric conversion element 102 having the APD 201 is provided on a photoelectric conversion substrate 11, and the other members are provided on a signal processing substrate 21.

[0070] The APD 201 generates charge pairs in response to incident light through photoelectric conversion. A voltage VL (first voltage) is supplied to the anode of the APD 201. A voltage VH (second voltage) higher than the voltage VL supplied to the anode is supplied to the cathode of the APD 201. A reverse bias voltage is supplied to the anode and cathode so that the APD 201 performs avalanche multiplication. With such a voltage supplied, the charge generated by the incident light undergoes avalanche multiplication, generating an avalanche current.

[0071] When a reverse bias voltage is supplied, there are two modes: Geiger mode, in which the device operates with the potential difference between the anode and cathode greater than the breakdown voltage, and linear mode, in which the device operates with the potential difference between the anode and cathode close to or less than the breakdown voltage.

[0072] An APD operated in Geiger mode is called a single photon avalanche diode (SPAD). For example, the voltage VL (first voltage) is −30 V, and the voltage VH (second voltage) is 1 V. The APD 201 may be operated in either linear mode or Geiger mode. A SPAD is preferable because the potential difference is larger than that of a linear mode APD, resulting in a more pronounced effect of withstanding voltage.

[0073] The quench element 202 is connected to a power supply that supplies voltage VH and the APD 201. The quench element 202 functions as a load circuit (quench circuit) during signal multiplication by avalanche multiplication, suppressing the voltage supplied to the APD 201 and suppressing avalanche multiplication (quench operation). The quench element 202 also returns the voltage supplied to the APD 201 to voltage VH by passing a current equivalent to the voltage drop caused by the quench operation (recharge operation). Although a passive quench element has been described above, an active quench element (MOS transistor) may be used and the recharge operation may be performed by inputting a clock pulse.

[0074] The sub-signal processing unit 103 has a waveform shaping unit 210 and a counter circuit 211. In this specification, it is sufficient that the sub-signal processing unit 103 has either the waveform shaping unit 210 or the counter circuit 211.

[0075] The waveform shaping unit 210 shapes the potential change of the cathode of the APD 201 obtained when photons are detected, and outputs a pulse signal. For example, an inverter circuit is used as the waveform shaping unit 210. While Fig. 10 shows an example in which one inverter is used as the waveform shaping unit 210, a circuit in which multiple inverters are connected in series, or another circuit that has a waveform shaping effect, may also be used.

[0076] The counter circuit 211 counts the pulse signal output from the waveform shaping unit 210 and holds a visible light photon count value. When the control pulse pRES is supplied, the signal held in the counter circuit 211 is reset.

[0077] The electrical connection may be switched by disposing a switch such as a transistor between the quench element 202 and the APD 201 or between the photoelectric conversion element 102 and the sub-signal processing unit 103. Similarly, the supply of the voltage VH or the voltage VL to the photoelectric conversion element 102 may be electrically switched using a switch such as a transistor.

[0078] In this embodiment, a configuration using a counter circuit 211 has been described. However, instead of the counter circuit 211, the photoelectric conversion unit 100 that acquires the pulse detection timing may be configured using a time-to-digital converter (hereinafter referred to as TDC) and a memory. In this case, the generation timing of the pulse signal output from the waveform shaping unit 210 is converted into a digital signal by the TDC. To measure the timing of the pulse signal, a control pulse pREF (reference signal) is supplied to the TDC from the vertical scanning circuit unit 110 in FIG. 9 via a drive line 214. The TDC acquires, as a digital signal, a signal obtained when the input timing of the signal output from each pixel via the waveform shaping unit 210 is converted into a relative time based on the control pulse pREF.

[0079] FIG. 11 is an example of a block diagram including an equivalent circuit of the signal processing unit 902. The sub-signal processing units 103 shown in FIG. 10 are arranged in five rows and five columns, and the visible light photon count values ​​held in the counter circuits 211 of the sub-signal processing units 103 are added together in an accumulator 903. A comparator 904 compares the accumulated visible light photon count value with a threshold value. The count value is then classified into a 30 kV to 60 kV radiation photon counter 905, a 60 kV to 90 kV radiation photon counter 906, or a 90 kV to 120 kV radiation photon counter 907. The counter is then incremented according to the energy of the radiation photons. After one frame period of imaging has elapsed, a control pulse pSEL is supplied from the vertical scanning circuit unit 110 in FIG. 9 via the drive line 214 in FIG. 10. The selection circuit 212 then switches between electrical connection and disconnection between the 30 kV to 60 kV radiation photon counter 905, the 60 kV to 90 kV radiation photon counter 906, and the 90 kV to 120 kV radiation photon counter 907 and the signal line 113. The selection circuit 212 then reads out the count values. The selection circuit 212 includes, for example, a buffer circuit for outputting signals.

[0080] 12A and 12B schematically show the relationship between the operation of the APD and the output signal. Fig. 12A is a diagram excerpting the APD 201, quench element 202, and waveform shaping unit 210 from Fig. 10. Here, the input side of the waveform shaping unit 210 is referred to as node A, and the output side is referred to as node B. The upper part of Fig. 12B shows the waveform change at node A, and the lower part shows the waveform change at node B.

[0081] Between time t0 and time t1, a potential difference of VH-VL is applied to the APD 201. When a photon is incident on the APD 201 at time t1, avalanche multiplication occurs in the APD 201, an avalanche multiplication current flows through the quench element 202, and the voltage at node A drops. When the amount of voltage drop further increases and the potential difference applied to the APD 201 decreases, avalanche multiplication in the APD 201 stops, as at time t2, and the voltage level at node A no longer drops below a certain value. Thereafter, between time t2 and time t3, a current that compensates for the voltage drop from voltage VL flows through node A, and at time t3, node A settles to its original potential level. At this time, the portion of the output waveform at node A that exceeds a certain threshold is shaped by the waveform shaping unit 210 and output as a signal at node B.

[0082] FIG. 13 is an equivalent circuit diagram when the sub-signal processing unit 103 shown in FIG. 9 is not used.

[0083] In Fig. 9, a sub-signal processing unit 103 is provided for each sub-pixel 101. In contrast to this, in the equivalent circuit diagram shown in Fig. 13, the anodes of the APDs 201 of the sub-pixels arranged in 5 rows and 5 columns are all connected to the quench elements 202 connected to the cathodes, and are combined into one pixel 901. This makes it possible to reduce the number of connections between the photoelectric conversion board 11 and the signal processing board 21, and there is no need to provide a sub-signal processing unit 103. In the configuration of Fig. 13, the energy of radiation photons can be identified from the peak value that appears at node C.

[0084] 14 and 15 show the structure of the photoelectric conversion substrate 11 (sensor substrate) of the photoelectric conversion unit. Fig. 14 is a plan view showing the configuration of the pixel region, and schematically shows the semiconductor layer 300 of the photoelectric conversion substrate 11 as viewed from the first surface side. Fig. 15 is a cross-sectional view taken along line A-A (diagonal direction) in Fig. 14.

[0085] As shown in Fig. 14 , a plurality of APDs 201 are arranged in a two-dimensional array in the row direction (left-right direction in Fig. 14 ) and column direction (up-down direction in Fig. 14 ) on the semiconductor layer 300. For convenience of illustration, only nine APDs 201 arranged in three rows and three columns are shown in Fig. 14 , but in an actual product, for example, hundreds of thousands to millions of APDs 201 are formed. One APD 201 corresponds to one pixel.

[0086] The semiconductor layer 300 is provided with an isolation portion 330, which is an isolation structure for reducing crosstalk between adjacent APDs 201. The isolation portion 330 is formed in a lattice pattern by a plurality of row-direction isolation portions 330X extending in the row direction and a plurality of column-direction isolation portions 330Y extending in the column direction. An APD 201 is disposed in each of the sections divided by the lattice-shaped isolation portions 330. In this embodiment, the section corresponding to one pixel has a substantially square shape in a plan view. The boundaries of the sections are arranged to overlap the isolation portions 330, for example. The cathode wiring contact plug 326 formed in the wiring structure 320 is disposed approximately in the center of the pixel (section), and the anode wiring contact plugs 324 are disposed at the four corners of the pixel. That is, in this embodiment, four anode contact plugs 324 are provided for one pixel.

[0087] As shown in FIG. 15 , the photoelectric conversion substrate 11 (sensor substrate) has a structure in which a semiconductor layer 300 and a wiring structure 320 are stacked. The surface of the semiconductor layer 300 facing the wiring structure 320 is referred to as the first surface, and the surface opposite the first surface is referred to as the second surface. The semiconductor layer 300 is made of, for example, silicon. A fixed charge film 310, an insulating film 311, a planarization film 312, and the like are stacked in this order on the second surface of the semiconductor layer 300, and microlenses 313 corresponding to each pixel are further provided above them. That is, the photoelectric conversion unit of this embodiment has a so-called back-illuminated structure in which light enters the semiconductor layer 300 from the second surface side. The second surface is sometimes referred to as the light incident surface. While FIG. 15 shows the microlenses 313, their provision is not essential. Furthermore, the following structure can be applied to the light incident surface. For example, a concave-convex structure, such as at least one or more concave or convex portions, is arranged on the second surface, which is the light incident surface. The concave-convex structure is formed by the silicon constituting the semiconductor layer 300 and other materials. For example, an insulator such as a silicon oxide film, a silicon oxynitride film, or a silicon nitride film is disposed in a recess provided in the semiconductor layer 300. An interface having a refractive index difference that is not parallel to the second surface is formed. With this configuration, incident light is diffracted, thereby improving sensitivity to light in the infrared region.

[0088] The fixed charge film 310 is made of a dielectric material having a negative fixed charge and is disposed over the entire second surface of the semiconductor layer 300. The material of the fixed charge film 310 is selected from, for example, hafnium oxide, aluminum oxide, zirconium oxide, titanium oxide, tantalum oxide, and ruthenium oxide, with aluminum oxide or hafnium oxide being preferred. The fixed charge film 310 may be configured with multiple layers. The insulating film 311 is disposed over the entire second surface on the fixed charge film 310. Suitable examples of the insulating film 311 include a silicon oxide film, a silicon oxynitride film, and a silicon nitride film. The insulating film 311 may be configured with multiple layers. Although not shown, in addition to the planarization film 312, a filter layer such as a color filter or an infrared light blocking filter may be provided on the second surface side of the semiconductor layer 300.

[0089] The wiring structure 320 is a structure in which multiple layers of wiring 321, 322, and 323, via plugs 325 and 327 that connect the wirings, a contact plug 324 for an anode wiring, a contact plug 326 for a cathode wiring, and the like are arranged in an insulating layer 329. The lower surface of the wiring structure 320 (the surface opposite to the semiconductor layer 300) is a bonding surface with the signal processing substrate 21, and a plurality of bonding portions 328 are provided on the bonding surface.

[0090] The semiconductor layer 300 includes a first semiconductor region 301, a second semiconductor region 302, a third semiconductor region 303, a fourth semiconductor region 304, a fifth semiconductor region 305, a sixth semiconductor region 306, a seventh semiconductor region 307, an eighth semiconductor region 308, and a ninth semiconductor region 309. Each semiconductor region is a region to which impurities are added by ion implantation or during the fabrication and epitaxial growth of the semiconductor substrate. Here, the first semiconductor region 301, the sixth semiconductor region 306, the seventh semiconductor region 307, and the eighth semiconductor region 308 are semiconductor regions of a first conductivity type (N-type in this embodiment). The second semiconductor region 302, the third semiconductor region 303, the fourth semiconductor region 304, the fifth semiconductor region 305, and the ninth semiconductor region 309 are semiconductor regions of a second conductivity type (P-type in this embodiment).

[0091] The first semiconductor region 301 is a semiconductor region of a first conductivity type (N-type in this embodiment) and is provided on a first surface of the semiconductor layer 300. The first semiconductor region 301 in this embodiment is formed in a circular shape in the center of the pixel (section) in plan view, as shown in Fig. 14. A cathode contact plug 326 is connected to the center position of the first semiconductor region 301.

[0092] The fifth semiconductor region 305 is a semiconductor region of the second conductivity type (P-type in this embodiment) and is arranged on the light incident side (closer to the second surface) than the first semiconductor region 301. The fifth semiconductor region 305 is formed in a layer shape at a predetermined depth so as to divide the first conductivity type epitaxial layer in one pixel (section) into upper and lower parts. The periphery of the fifth semiconductor region 305 is in contact with the isolation portion 330 that surrounds the pixel. The first conductivity type epitaxial layer closer to the first surface than the fifth semiconductor region 305 is the sixth semiconductor region 306, and the first conductivity type epitaxial layer closer to the second surface is the eighth semiconductor region 308.

[0093] The first semiconductor region 301 of the first conductivity type and the fifth semiconductor region 305 of the second conductivity type form an avalanche multiplication unit AM through a PN junction. Signal charges generated in the eighth semiconductor region 308 by photoelectric conversion are collected in the avalanche multiplication unit AM. In order to improve the sensitivity of the APD 201, it is preferable to increase the size of the eighth semiconductor region 308, which corresponds to the sensitive region.

[0094] The seventh semiconductor region 307 is a semiconductor region of the first conductivity type formed around the first semiconductor region 301. The seventh semiconductor region 307 is also formed in a circular shape in a plan view. The impurity concentrations of the first semiconductor region 301, the seventh semiconductor region 307, and the sixth semiconductor region 306 are set to satisfy the relationship: first semiconductor region 301 > seventh semiconductor region 307 > sixth semiconductor region 306. That is, the impurity concentration of the first semiconductor region 301 is the highest, and the impurity concentration of the seventh semiconductor region 307 is set to be between the first semiconductor region 307 and the sixth semiconductor region 306. This ensures electrical connection between the cathode and the first semiconductor region 301 (i.e., the APD 201). The seventh semiconductor region 307 also serves as a guard ring for electric field relaxation.

[0095] The ninth semiconductor region 309 is a buried layer of the second conductivity type provided over the entire second surface of the semiconductor layer 300. The ninth semiconductor region 309 serves to suppress noise from the second surface side. A voltage VL from the anode wiring can be supplied to the ninth semiconductor region 309 via the second semiconductor region 302. In this case, a potential gradient can be formed to collect charges.

[0096] FIG. 16 is a cross-sectional view showing an example of a configuration in which a scintillator section 500 is laminated on a laminated photoelectric conversion section 100.

[0097] The photoelectric conversion substrate 11 (sensor substrate) has a first semiconductor layer 300 and a first wiring structure 320. The above-described photoelectric conversion elements 102 and the like are formed on the first semiconductor layer 300. Pads 361 are exposed on the signal processing substrate 21. These pads 361 are aluminum electrodes used for connection to an external device. For example, they are connected to an external control unit 960 or an image processing unit (not shown) via bonding wires, an anisotropic conducting film (ACF), and flexible printed circuits (FPC) 962.

[0098] The signal processing board 21 (circuit board) has a second semiconductor layer 420 and a second wiring structure 410. A sub-signal processing section 103 having a plurality of semiconductor elements is formed in the second semiconductor layer 420. Furthermore, wiring 401 related to the sub-signal processing section 103 is formed in the second wiring structure 410.

[0099] The photoelectric conversion substrate 11 and the signal processing substrate 21 are stacked so that their wiring structures 320 and 410 face each other, thereby producing the photoelectric conversion unit 100. The photoelectric conversion substrate 11 and the signal processing substrate 21 are electrically connected via a joint 328.

[0100] The scintillator section 500 is a pixel scintillator (or partition phosphor) that is separated to correspond to the pixels 901 of the photoelectric conversion section 100. In other words, the scintillator section 500 is separated for each pixel. By using a pixel scintillator, it is possible to reduce the incidence of visible light photons 700 on adjacent pixels, and to count the number of visible light photons 700 resulting from one radiation photon per pixel.

[0101] The following materials can be used for the scintillator section 500. Materials that have good radiation stopping properties and respond in 1 us or less include CsI:Tl, CsI, LuAlO:Ce, PbWO, GdAlGaO:Ce, (Lu,Y)SiO:Ce, and GdOS:Tb(Pr).

[0102] The pixel scintillator is formed by forming a metal reflective layer 955 as needed on a partition substrate 954, which is formed by penetrating a Si wafer using a through-silicon via (TSV) process, and then filling it with scintillator material. In addition to using a Si wafer, the partition substrate can also be formed by depositing polyimide resin (PI). Other methods for filling the scintillator material include heating and melting the scintillator material, or powdering the scintillator material and mixing it with a binder resin. When using a material that deliquesces in humidity, such as CsI, it is recommended to use ALPET, a composite material made by bonding aluminum foil and polyester film, as the moisture-proof layer 952. Furthermore, when bonding the moisture-proof layer 952 to the partition substrate 954, bonding it with a reflective layer 953 made of a hot-melt sheet containing TiO2 can achieve both moisture resistance and reflective performance.

[0103] Example 1 Example 1 of the fourth embodiment will be described with reference to the configuration diagrams of the radiation detection panel in FIGS. 17 and 18. FIG.

[0104] When radiation photon A in Fig. 17 enters scintillator 956, it is multiplied by scintillator 956. Normally, a scintillator multiplies radiation photons into visible light photons by about 1000 times, but here, for simplicity, a case where the radiation photons are multiplied by two is shown.

[0105] The amplified visible light photons C and B are incident on subpixels 3 and 5, respectively, and can be counted. Meanwhile, radiation photons D incident on the edge of the scintillator 956 are amplified into visible light photons E and visible light photons F. Visible light photon E is incident on subpixel 1, and visible light photon F is reflected by the metal reflective layer 955 and incident on subpixel 1, just like visible light photon E. Since visible light photons E and visible light photon F simultaneously enter subpixel 1, they pile up, and even though two visible light photons are incident, they are counted as one. When the scintillator 956 is separated by the partition substrate 954, the reflectivity of the metal reflective layer 955 causes differences in the count of visible light photons emitted at the edge and center of the scintillator 956. Specifically, accurate counting becomes impossible at the edge. Therefore, when estimating the energy of radiation photons from the count number of visible light photons, the energy of the radiation photons cannot be measured correctly.

[0106] FIG. 18 is a diagram showing the configuration of the radiation detection panel of the first embodiment.

[0107] In the radiation detection panel of Example 1, the size of the photoelectric conversion elements 102 in subpixels 1 and 9 located at the edges of the scintillator 956 is halved. That is, within the pixel 901, the size of the edge subpixels 101 is halved. As a result, the light-receiving area of ​​the photoelectric conversion elements 102 in the edge subpixels of the pixel 901 is also halved. Therefore, of the visible light photons E and visible light photons F that are multiplied from the radiation photons D that are incident on the edge of the scintillator 956, only the visible light photons E are incident on subpixel 1. Because the size of the photoelectric conversion elements 102 in subpixel 1 is half, the count is doubled and processed in the subsequent sub-signal processing unit 103. By performing such processing, it is possible to accurately estimate the visible light photons that are multiplied from the radiation photons D that are incident on the edge of the scintillator 956. Although the size of the photoelectric conversion elements 102 is changed in Example 1, the density of subpixels 1 may also be changed. Correction can be made by lowering the density of subpixel 1 only at the end of the scintillator 956 and multiplying the density by the reciprocal in the sub-signal processing unit 103 at the subsequent stage.

[0108] In this way, by changing the visible light detection sensitivity of the subpixels per unit area in the phosphor-formed region within one pixel, it becomes possible to measure visible light photons with high accuracy. Specifically, in Example 1, by lowering the visible light detection sensitivity per unit area of ​​the subpixels at the ends of the scintillator 956 compared to the central portion of the scintillator 956, it becomes possible to measure visible light photons with high accuracy. Note that in this example, the size of the subpixels 101 at the ends of one pixel 901 is reduced, but this is not limiting, and the size may be gradually reduced from the center to the ends. An example of this case is shown in FIG. 23. Note that FIG. 23 shows the top view of one pixel (as viewed from the Z-axis direction in FIG. 7).

[0109] Example 2 of the Fourth Embodiment Example 2 will be described with reference to the configuration diagram of the radiation detection panel in FIG.

[0110] In Example 1, the visible light detection sensitivity of the subpixel per unit area at the end of the scintillator 956 is reduced by reducing the size of the photoelectric conversion element 102. In Example 2, however, a light-shielding layer 970, which is a shielding member, is disposed on the photoelectric conversion element 102. The light-shielding layer 970 is disposed only at the end of the scintillator, thereby reducing the visible light detection sensitivity of the subpixel per unit area of ​​subpixel 1 and subpixel 9. By providing the light-shielding layer 970, the same effect as in Example 1 can be achieved without changing the photoelectric conversion element 102 of the subpixel. The light-shielding layer 970 can be formed by reducing the aperture ratio of the photoelectric conversion element 102 using metal wiring, or by inserting an organic or inorganic material that absorbs visible light between the scintillator 956 and the photoelectric conversion element 102. Note that, in this example as well, the aperture area of ​​the light-shielding layer may be reduced from the center to the end within one pixel 901 (see FIG. 24 ). The opening 971 is an opening provided in the light-shielding layer 970, and the size of the opening 971 indicates the size of the opening area.

[0111] Example 3 of the Fourth Embodiment Example 3 will be described with reference to the configuration diagram of the radiation detection panel in FIG.

[0112] In Examples 1 and 2, the size and light-shielding area of ​​the photoelectric conversion elements 102 at the ends of the scintillator 956 are changed, but in Example 3, the photoelectric conversion elements 102 at the ends of the scintillator 956 are divided. In other words, the photoelectric conversion elements 102 are divided so that the density of the sub-pixels at the ends is higher within a pixel. By dividing them, visible light photons emitted in the vicinity can be accurately measured, and pile-up becomes less likely to occur.

[0113] Dividing all of the subpixels in a pixel into two will increase power consumption accordingly, but by dividing only the end portion of the scintillator 956, it is possible to prevent pileup while suppressing the increase in power consumption.

[0114] Example 4 of the Fourth Embodiment Example 4 will be described with reference to the configuration diagram of the radiation detection panel in FIG.

[0115] In Example 4, the driving method of the sub-signal processing unit 103 is different between the end and center of the scintillator 956. Visible light photons E and visible light photons F may be emitted simultaneously or may emit at different times depending on the time response. Therefore, by speeding up the timing of resetting the active quench element of the sub-signal processing unit 103, visible light photons E and visible light photons F can be counted individually. For example, the reset clock signals pSetA and pSetI are reset at 20 MHz, and the clock signals pSetB to pSetH operate at 10 MHz. In other words, the sampling rate of the sub-signal processing unit is different between the end sub-pixels and the center sub-pixels of the pixel. By differentiating the processing capabilities of the sub-signal processing units in this way, pileup can be suppressed and various power consumption can be reduced. While speeding up the clock signal increases power consumption, the increase in power consumption is minimized by speeding up the clock signal partially. In this way, it becomes possible to accurately count visible light photons while reducing power consumption by using different driving methods for the ends and the center of the scintillator 956. Such driving effectively increases the visible light detection sensitivity of the subpixel per unit area.

[0116] Furthermore, by continuously inputting any one of the clock signals pSetA to pSetI (continuing to hold the photoelectric conversion elements 102 in a reset state), it is possible to stop the operation of a specific photoelectric conversion element 102. In Examples 1 to 3, the size of the photoelectric conversion elements 102 is physically changed or a light-shielding layer is provided, but by stopping the operation of a specific photoelectric conversion element 102, it is also possible to reduce the visible light detection sensitivity of the subpixel per unit area.

[0117] Furthermore, the visible light detection sensitivity of the subpixel per unit area can be changed by halving the size of the photoelectric conversion element 102 or by stopping the operation of the photoelectric conversion element 102. In this case, the visible light detection sensitivity of the subpixel per unit area can also be adjusted by arranging the subpixels in a staggered arrangement (alternately) rather than arranging them all at the ends of the scintillator 956.

[0118] In the fourth embodiment, the following configuration is also possible.

[0119] If the partition substrate 954 is made of a material that absorbs visible light, such as a Si wafer, the visible light emitted by the scintillator 956 is absorbed by the partition substrate 954. As a result, the edges of the scintillator 956, which are closer to the partition substrate 954, may emit fewer visible photons than the center of the scintillator 956. When visible light is absorbed by the partition substrate 954 in this manner, the size of the photoelectric conversion elements 102 may be made smaller in the center of the scintillator 956 than in the edges of the scintillator 956. Alternatively, the density of the photoelectric conversion elements 102 in the center of the scintillator 956 may be made higher than in the edges of the scintillator 956. Correction can be achieved by changing the size and density of the photoelectric conversion elements 102 between the center and the edges of the scintillator 956, and multiplying the reciprocals of the size and density of the photoelectric conversion elements 102 in the downstream sub-signal processing unit 103.

[0120] Furthermore, when visible light is absorbed by the partition substrate 954, a light-shielding layer 970 for blocking visible light is disposed on the photoelectric conversion element 102, and the blocked area at the ends of the scintillator 956 is made smaller than that at the center of the scintillator 956. Even if the partition substrate 954 is a material that absorbs visible light, visible light can be detected uniformly at the center and ends of the scintillator 956, and pile-up can be suppressed.

[0121] Furthermore, when visible light is absorbed by the partition substrate 954 , the timing for resetting the active quench element of the sub-signal processing unit 103 may be earlier in the center of the scintillator 956 than in the end portion of the scintillator 956 .

[0122] For example, the clock signals pSetA and pSetI for resetting are reset at 10 MHz. The clock signals pSetB to pSetH can be operated at 20 MHz to suppress pile-up and also reduce power consumption. The partition substrate 954 can be made of a material that reflects visible light and a material that absorbs visible light, and pile-up can be suppressed by taking opposite approaches.

[0123] [Application] FIG. 25 is a conceptual diagram of an X-ray diagnostic system (radiation imaging system) using a radiation imaging device according to the present disclosure. X-rays 711, which are radiation generated by an X-ray tube 710 (radiation source), pass through a chest 721 of a patient or subject 720 and enter the radiation imaging device 100 according to the present disclosure, which includes a scintillator unit 200. The incident X-rays contain information about the inside of the patient's body. The scintillator unit 200 emits light in response to the incidence of the X-rays, which undergoes photoelectric conversion to obtain electrical information. This information is converted into a digital signal, which is image-processed by an image processor 730, which serves as a signal processing means, and can be observed on a display 740, which serves as a display means in a control room (X-ray room). The radiation imaging system includes at least the radiation imaging device 100 and the image processor 730, which processes signals from the radiation imaging device 100.

[0124] Furthermore, the image processed by the image processor 730 can be transmitted to a remote location (for example, a doctor's room) via a transmission processing means such as a telephone line 750. The transmitted image can be displayed on a display 741, which is a display means in the doctor's room, or can be stored on a recording means such as an optical disk, so that a doctor in a remote location can use the transmitted image to make a diagnosis. The transmitted image can also be recorded on a recording medium, film 761, by a film processor 760, which is a recording means.

[0125] The present invention is not limited to the above-described embodiments, and various modifications and variations are possible without departing from the spirit and scope of the invention. For example, although the above-described embodiments have been described using a radiation imaging device as an example, the present invention can also be applied to an imaging device having subpixels using photoelectric conversion elements that convert visible light into electric charges. Therefore, the following claims are appended to clarify the scope of the invention.

[0126] The disclosure of this specification includes the following imaging apparatus, radiation imaging apparatus, and radiation imaging system.

[0127] (Item 1) A radiation imaging device having a scintillator section that converts incident radiation into visible light and a photoelectric conversion section that converts the visible light into an electrical signal, wherein the scintillator section has a first visible light absorbing member that separates a plurality of scintillators included in the scintillator section in a direction perpendicular to the incident direction of the radiation, and a second visible light absorbing member provided on the radiation incident surface.

[0128] (Item 2) The radiation imaging device according to item 1, wherein the first visible light absorbing member is either glass or silicon, and the second visible light absorbing member is either glass, amorphous carbon, CFRP, a resin film, aluminum, or titanium.

[0129] (Item 3) The radiation imaging device according to item 1, wherein a third visible light absorbing member is disposed between the scintillator section and the photoelectric conversion section.

[0130] (Item 4) The radiation imaging device according to item 3, wherein the third visible light absorbing member is an FOP.

[0131] (Item 5) The radiation imaging device according to item 1, wherein the photoelectric conversion unit has a fourth visible light absorbing member provided on a side of the photoelectric conversion unit on which the visible light is incident.

[0132] (Item 6) The radiation imaging device described in Item 5, characterized in that the photoelectric conversion unit has a photoelectric conversion element that converts the visible light into an electrical signal, and the fourth visible light absorbing member is a metal film or an organic film formed in a position that reduces the visible light that enters the photoelectric conversion element.

[0133] (Item 7) The radiation imaging device described in Item 1, characterized in that the first visible light absorbing member and the second visible light absorbing member are integrally formed, and the first visible light absorbing member and the second visible light absorbing member are silicon.

[0134] (Item 8) A radiation imaging system comprising: a radiation source that generates radiation; and a radiation imaging device according to any one of items 1 to 7 that detects the radiation generated by the radiation source.

[0135] (Item 9) A scintillator panel that converts incident radiation into visible light, comprising: a first visible light absorbing member that separates a plurality of scintillators included in the scintillator panel in a direction perpendicular to the incident direction of the radiation; and a second visible light absorbing member provided on the radiation incident surface.

[0136] (Item 10) A radiation imaging device comprising: a scintillator that generates visible light in response to irradiation with radiation; a plurality of sub-pixels each including a photoelectric conversion element that generates electric charges in response to the visible light generated by the scintillator and a sub-signal processing unit that generates electric signals in response to the electric charges generated by the photoelectric conversion element; and a signal processing unit that processes the electric signals generated by the sub-signal processing units included in the plurality of sub-pixels to generate pixel signals, wherein the visible light receiving area of ​​the photoelectric conversion element differs among the plurality of sub-pixels.

[0137] (Item 11) The radiation imaging apparatus according to item 10, wherein the scintillator is separated for each of the pixels.

[0138] (Item 12) The radiation imaging device according to item 10 or 11, wherein, of the plurality of sub-pixels included in the pixel, the sub-pixels located at the ends of the pixel and the sub-pixels located in the center of the pixel have different light receiving areas.

[0139] (Item 13) The radiation imaging device according to any one of items 10 to 12, wherein the plurality of sub-pixels included in the pixel are arranged so as to have different densities within the pixel.

[0140] (Item 14) The radiation imaging device according to any one of items 10 to 13, further comprising a visible light shielding member between the scintillator and the photoelectric conversion elements of a plurality of subpixels included in the pixel, wherein the opening areas of the shielding member are different for the plurality of subpixels.

[0141] (Item 15) A radiation imaging device comprising: a scintillator that generates visible light when irradiated with radiation; a plurality of sub-pixels each including a photoelectric conversion element that generates electric charges when irradiated with radiation by the scintillator and a sub-signal processing unit that generates electric signals when irradiated with the electric charges by the photoelectric conversion element; and a signal processing unit that processes the electric signals generated by the sub-signal processing units included in the plurality of sub-pixels to generate pixel signals, wherein the processing capability of the sub-signal processing units to process the electric charges differs among the plurality of sub-pixels, the imaging device described in any one of items 10 to 14.

[0142] (Item 16) The radiation imaging device according to any one of items 10 to 15, wherein the sub-signal processing units of the plurality of sub-pixels included in the pixel are driven by different methods for the plurality of sub-pixels.

[0143] (Item 17) The radiation imaging device according to any one of items 10 to 16, wherein the sub-signal processing units of the plurality of sub-pixels included in the pixel have different sampling rates for receiving the charges generated by the photoelectric conversion elements in the plurality of sub-pixels.

[0144] (Item 18) A radiation imaging device comprising: a scintillator that generates visible light in response to irradiation with radiation; a photoelectric conversion unit that receives the visible light generated by the scintillator and generates an electrical signal; and a signal processing unit that classifies and outputs the electrical signal generated by the photoelectric conversion unit, wherein the photoelectric conversion unit comprises a plurality of sub-pixels, each of which includes a photoelectric conversion element that receives the visible light and generates an electric charge, and a sub-signal processing unit that receives the electric charge generated by the photoelectric conversion element and generates the electric signal, and the plurality of sub-pixels are arranged so that the visible light detection sensitivity per unit area of ​​the photoelectric conversion unit is different from one another.

[0145] (Item 19) A radiation imaging device comprising a plurality of sub-pixels, each of which includes a photoelectric conversion element that receives visible light and generates an electric charge, and a sub-signal processing unit that receives the electric charge generated by the photoelectric conversion element and generates an electric signal, and a signal processing unit that processes the electric signal generated by the sub-signal processing units included in the plurality of sub-pixels to generate a pixel signal, wherein the visible light receiving area of ​​the photoelectric conversion element differs among the plurality of sub-pixels.

[0146] (Item 20) A radiation imaging device comprising a plurality of sub-pixels, each of which has a photoelectric conversion element that receives visible light to generate an electric charge and a sub-signal processing unit that receives the electric charge generated by the photoelectric conversion element and generates an electric signal, and a signal processing unit that processes the electric signal generated by the sub-signal processing units provided in the plurality of sub-pixels to generate a pixel signal, characterized in that the processing capability of the sub-signal processing units to process the electric charges differs among the plurality of sub-pixels.

[0147] (Item 21) A radiation imaging system comprising: the radiation imaging device according to any one of items 10 to 18; and a control device that acquires image data from the radiation imaging device and processes the image data.

[0148] (Item 22) An imaging system including the imaging device according to item 19 or 20, and a control device that acquires image data from the imaging device and processes the image data.

[0149] The present invention is not limited to the above-described embodiments, and various modifications and variations can be made without departing from the spirit and scope of the present invention. Therefore, to apprise the public of the scope of the present invention, the following claims are appended.

[0150] This application claims priority based on Japanese Patent Application Nos. 2024-152991 and 2024-152990 filed on September 5, 2024, and Japanese Patent Application No. 2025-108932 filed on June 27, 2025, the entire contents of which are incorporated herein by reference.

[0151] 980 Radiation imaging system 990 Radiation imaging device 984 Radiation generating device 985 Radiation 500 Scintillator section 11 Photoelectric conversion board 21 Signal processing board 102 Photoelectric conversion element 956 Scintillator

Claims

1. A radiation imaging device having a scintillator section that converts incident radiation into visible light and a photoelectric conversion section that converts the visible light into an electrical signal, wherein the scintillator section has a first visible light absorbing member that separates multiple scintillators included in the scintillator section in a direction perpendicular to the incident direction of the radiation, and a second visible light absorbing member provided on the radiation incident surface.

2. The radiation imaging device according to claim 1, wherein the first visible light absorbing member is either glass or silicon, and the second visible light absorbing member is either glass, amorphous carbon, CFRP, a resin film, aluminum, or titanium.

3. The radiation imaging device according to claim 1, wherein a third visible light absorbing member is disposed between the scintillator section and the photoelectric conversion section.

4. The radiation imaging apparatus according to claim 3, wherein the third visible light absorbing member is an FOP.

5. A radiation imaging device according to claim 1, wherein the photoelectric conversion section has a fourth visible light absorbing member provided on the side of the photoelectric conversion section where the visible light is incident.

6. The radiation imaging device according to claim 5, wherein the photoelectric conversion unit has a photoelectric conversion element that converts the visible light into an electrical signal, and the fourth visible light absorbing member is a metal film or an organic film formed in a position that reduces the visible light incident on the photoelectric conversion element.

7. A radiation imaging device as described in claim 1, characterized in that the first visible light absorbing member and the second visible light absorbing member are formed integrally, and the first visible light absorbing member and the second visible light absorbing member are made of silicon.

8. A radiation imaging system comprising: a radiation source that generates radiation; and a radiation imaging device according to any one of claims 1 to 7 that detects the radiation generated by the radiation source.

9. A scintillator panel that converts incident radiation into visible light, comprising: a first visible light absorbing member that separates multiple scintillators contained in the scintillator panel in a direction perpendicular to the incident direction of the radiation; and a second visible light absorbing member provided on the radiation incident surface.

10. A radiation imaging device comprising: a scintillator that generates visible light in response to irradiation with radiation; a plurality of sub-pixels each including a photoelectric conversion element that generates electric charges in response to the visible light generated by the scintillator and a sub-signal processing unit that generates electric signals in response to the charges generated by the photoelectric conversion element; and a signal processing unit that processes the electric signals generated by the sub-signal processing units included in the plurality of sub-pixels to generate pixel signals, wherein the visible light receiving area of ​​the photoelectric conversion element differs among the plurality of sub-pixels.

11. A radiation imaging apparatus according to claim 10, wherein said scintillator is separated for each of said pixels.

12. A radiation imaging device according to claim 10, wherein, of the plurality of sub-pixels included in the pixel, the light receiving areas of the sub-pixels located at the edges of the pixel are different from those of the sub-pixels located in the center.

13. The radiation imaging device according to claim 10, wherein the plurality of sub-pixels included in the pixel are arranged so as to have different densities within the pixel.

14. The radiation imaging device according to claim 10, further comprising a visible light shielding member between the scintillator and the photoelectric conversion elements of the plurality of subpixels included in the pixel, the opening area of ​​the shielding member being different for each of the plurality of subpixels.

15. A radiation imaging device comprising: a scintillator that generates visible light in response to irradiation with radiation; a plurality of sub-pixels each including a photoelectric conversion element that generates electric charges in response to the visible light generated by the scintillator and a sub-signal processing unit that generates electric signals in response to the electric charges generated by the photoelectric conversion element; and a signal processing unit that processes the electric signals generated by the sub-signal processing units included in the plurality of sub-pixels to generate pixel signals, wherein the processing capability of the sub-signal processing units to process the electric charges differs for each of the plurality of sub-pixels.

16. A radiation imaging apparatus according to claim 15, wherein said scintillator is separated for each of said pixels.

17. The radiation imaging device according to claim 15, wherein the sub-signal processing units of the plurality of sub-pixels included in the pixel are driven in different ways for the plurality of sub-pixels.

18. The radiation imaging device according to claim 14, wherein the sub-signal processing units of the plurality of sub-pixels included in the pixel have different sampling rates for receiving the charges generated by the photoelectric conversion elements in the plurality of sub-pixels.

19. A radiation imaging device comprising: a scintillator that generates visible light in response to irradiation with radiation; a photoelectric conversion unit that receives the visible light generated by the scintillator and generates an electrical signal; and a signal processing unit that classifies and outputs the electrical signal generated by the photoelectric conversion unit, wherein the photoelectric conversion unit comprises a plurality of sub-pixels, each of which has a photoelectric conversion element that receives the visible light and generates an electric charge, and a sub-signal processing unit that receives the electric charge generated by the photoelectric conversion element and generates the electric signal, and the plurality of sub-pixels are arranged so that the visible light detection sensitivity per unit area of ​​the photoelectric conversion unit is different from one another.

20. A radiation imaging device comprising a plurality of sub-pixels, each of which has a photoelectric conversion element that receives visible light and generates an electric charge, and a sub-signal processing unit that receives the electric charge generated by the photoelectric conversion element and generates an electric signal, and a signal processing unit that processes the electric signal generated by the sub-signal processing units of the plurality of sub-pixels to generate a pixel signal, wherein the visible light receiving area of ​​the photoelectric conversion element differs among the plurality of sub-pixels.

21. A radiation imaging device comprising a plurality of sub-pixels, each of which has a photoelectric conversion element that receives visible light and generates an electric charge, and a sub-signal processing unit that receives the electric charge generated by the photoelectric conversion element and generates an electric signal, and a signal processing unit that processes the electric signal generated by the sub-signal processing units of the plurality of sub-pixels to generate a pixel signal, wherein the processing capability of the sub-signal processing unit to process the electric charge differs among the plurality of sub-pixels.

22. A radiation imaging system comprising the radiation imaging device according to claim 10 and a control device that acquires image data from the radiation imaging device and processes the image data.

Images