Radiation imaging device and radiation imaging system

Abstract

A radiation imaging device according to the present disclosure comprises a scintillator unit having a plurality of scintillators that convert radiation into visible light, and a semiconductor substrate having a photoelectric conversion element that converts the visible light into an electric signal, the radiation imaging device being characterized in that the semiconductor substrate is less thick than the scintillator unit.

Description

Radiation imaging device and radiation imaging system

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

[0002] A radiation imaging device using a flat panel detector (hereinafter referred to as FPD) formed from a semiconductor material is known as an imaging device used for medical image diagnosis and non-destructive testing using radiation (such as X-rays). Such a radiation imaging device is used, for example, in medical image diagnosis to acquire still images, moving images, etc.

[0003] FPDs include, for example, integral sensors and photon-counting sensors. Integral sensors measure the total amount of charge generated by incident radiation. In contrast, photon-counting sensors identify the energy (wavelength) of incident radiation and count the number of times radiation is detected for each of multiple energy levels. In other words, photon-counting sensors have energy resolution. For this reason, there are hopes that the photon-counting method will improve diagnostic capabilities in various fields.

[0004] Photon-counting sensors are further classified into direct and indirect types. Direct types count the number of times radiation is detected by directly detecting the energy of radiation using CdTe or the like. Indirect types detect the intensity of visible light generated in a scintillator when radiation is incident on the sensor and count the number of times the visible light is detected.

[0005] Patent Document 1 discloses an X-ray CT apparatus using an indirect photon-counting sensor.

[0006] JP 2019-86443 A

[0007] Here, a radiation imaging device using an FPD is required to produce images with higher sharpness over a wider range than an X-ray CT device using an FPD, and therefore the area of ​​the scintillator in the radiation imaging device is larger than the area of ​​the scintillator in the X-ray CT device.

[0008] However, it is not easy in terms of cost and manufacturing technology to form a scintillator having a thickness equal to that of a scintillator in an X-ray CT device and a larger area than the scintillator. Therefore, the thickness of the scintillator in a radiation imaging device is thinner than that of a scintillator in a CT device. As a result, not all X-rays incident on the scintillator may be absorbed by the scintillator. X-rays that are not absorbed by the scintillator and that pass through the scintillator are absorbed by a semiconductor substrate on which a photoelectric conversion element and a counter circuit are formed. X-rays absorbed by the semiconductor substrate are directly converted into an electrical signal without being converted into visible light, which may result in a decrease in image quality (e.g., image sharpness or energy resolution).

[0009] Therefore, an object of the present disclosure is to provide a radiation imaging device that improves image quality by suppressing the direct conversion of X-rays absorbed by a semiconductor substrate into an electrical signal.

[0010] The radiation imaging device disclosed herein is a radiation imaging device having a scintillator section having a plurality of scintillators that convert radiation into visible light, and a semiconductor substrate having a photoelectric conversion element that converts the visible light into an electrical signal, and is characterized in that the thickness of the semiconductor substrate is thinner than the thickness of the scintillator section.

[0011] According to the present disclosure, it is possible to provide a radiation imaging device with improved image quality (for example, image sharpness or energy resolution).

[0012] Schematic configuration diagram of a radiation imaging device according to a first embodiment of the present disclosure. Cross-sectional view of a radiation imaging device according to the first embodiment of the present disclosure. Schematic configuration diagram of a radiation imaging device according to a second embodiment of the present disclosure. Cross-sectional view of a radiation imaging device according to the second embodiment of the present disclosure. Schematic configuration diagram of a radiation imaging device according to a third embodiment of the present disclosure. Cross-sectional view of a radiation imaging device according to the third embodiment of the present disclosure. Schematic view illustrating an application example of the radiation imaging device according to the present disclosure to an X-ray diagnostic system.

[0013] Embodiments of the present disclosure will be described below with reference to the accompanying drawings. Similar elements throughout the various embodiments will be designated by the same reference numerals, and redundant description will be omitted. Below, each embodiment of the present disclosure will be described in the context of a radiation imaging device used in medical imaging diagnostic devices, analytical devices, etc. In this disclosure, light includes visible light and infrared light, and radiation includes X-rays, alpha rays, beta rays, and gamma rays.

[0014] First Embodiment An example of the schematic configuration of a radiation imaging apparatus 100 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 100, and Fig. 2 is a cross-sectional view.

[0015] 1 and 2 , the radiation imaging device 100 includes a scintillator unit 200 and a photoelectric conversion unit 300. X-ray photons 400 are incident on the scintillator unit 200 and are converted and amplified into visible light photons 500 by the scintillator unit 200. The visible light photons 500 are incident on the photoelectric conversion unit 300 and detected as an electrical signal. The photoelectric conversion unit 300 includes a photodiode substrate 310 and a counter circuit substrate 320. The photoelectric conversion unit 300 is an example of a semiconductor substrate that converts the visible light into an electrical signal.

[0016] The scintillator section 200 has a plurality of scintillators 210 separated by separation regions 220. Separation by the separation regions 220 can suppress the spread of visible light photons generated in the scintillators 210. Known materials such as CsI:Tl and Gd2O2S (GOS) can be used as the scintillators 210. The thickness of the scintillators 210 is preferably 300 μm to 1000 μm.

[0017] The scintillator section 200 and the photoelectric conversion section 300 are disposed on top of each other with an adhesive member 201 interposed therebetween, as shown in FIG. 2 . The adhesive member 201 can be an adhesive member that melts or softens when heated. The adhesive member 201 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 201 can also be an acrylic-based, silicone-based, or other adhesive sheet that has adhesive properties at room temperature.

[0018] The photodiode substrate 310 and the counter circuit substrate 320 are stacked and electrically connected. The photodiode substrate 310 has a pixel region 311 in which pixels 301 having a sensor function are arranged in a matrix. Therefore, the photodiode substrate 310 may be called a sensor member, a sensor substrate, a sensor chip, etc. The counter circuit substrate 320 has a circuit region 321 that processes signals detected in the pixel region 311. The counter circuit substrate 320 may be called a circuit member, a circuit substrate, a circuit chip, etc. Furthermore, both substrates are semiconductor substrates made of silicon wafers.

[0019] The photodiode substrate 310 has a first semiconductor layer 311 and a first wiring structure 312 that form photodiodes, which are photoelectric conversion elements, and these constitute the pixels 301. As described above, in the scintillator section 200 of this embodiment, the scintillator 210 is separated by the separation region 220, so the volume of the scintillator 210 is smaller than in a scintillator section that does not have the separation region 220. This reduces the amount of light emitted by the scintillator 210, so an avalanche photodiode that has a function of amplifying a signal is suitable for the photoelectric conversion element. The counter circuit substrate 320 has a second semiconductor layer 321 and a second wiring structure 322 that constitute circuits such as a signal processing section, corresponding to the pixels 301.

[0020] The photoelectric conversion unit 300 is a back-illuminated photoelectric conversion unit in which visible light photons 500 are incident from a second surface opposite to the first surface on which the first wiring structure 312 is formed, and a counter circuit board 320 is disposed on the first surface side. When the photoelectric conversion unit has a back-illuminated structure, the thickness of the photoelectric conversion unit 300 (semiconductor substrate) is preferably 10 μm or less. Note that the inventions according to each embodiment can also be applied to photoelectric conversion units having a front-illuminated structure. When the photoelectric conversion unit has a front-illuminated structure, the thickness of the photoelectric conversion unit 300 (semiconductor substrate) is 300 μm to 500 μm. In order to make the photoelectric conversion unit 300 thinner than the scintillator 210, a back-illuminated structure is more preferable than a front-illuminated structure. The back-illuminated structure provides a larger difference in thickness between the photoelectric conversion unit 300 and the scintillator 210 than the front-illuminated structure, thereby improving image quality.

[0021] As described above, X-ray photons 400 irradiated onto the radiation imaging device 100 are absorbed by the scintillator section 200 and converted into visible light photons 500. However, not all X-ray photons 400 are absorbed by the scintillator section 200, and some X-ray photons 400 pass through the scintillator section 200. For example, the absorption rate of X-rays is determined by the absorption coefficient and thickness of the scintillator material. If the scintillator is not thick enough, the amount of X-ray photons 400 passing through the scintillator increases. Furthermore, although the X-ray energy is set to an appropriate level for imaging, high-energy X-ray photons may be present due to variations in the X-ray generator, and these may pass through the scintillator section 200. The X-ray photons 400 that pass through the scintillator section 200 enter the photoelectric conversion section 300 and then enter the photodiode substrate 310 or the counter circuit substrate 320. When the material of each substrate is silicon, the X-ray photons 400 are directly converted into electrical signals, which may be read out as signals, resulting in a false detection, which may degrade the image quality. For example, if the X-ray photons 400 are detected in a pixel adjacent to the pixel where they are incident, this may result in a blurred image, and if they are detected while measuring the light emission of the scintillator in the same pixel, errors may occur in the output value or energy resolution.

[0022] In this embodiment, the thicknesses of the photodiode substrate 310 and the counter circuit substrate 320 are thinner than the thickness of the scintillator 210. Since the X-ray absorption coefficient of silicon is typically smaller than that of the scintillator material, thinning the photodiode substrate 310 and the counter circuit substrate 320 can reduce the probability of X-ray photons 400 being absorbed by each substrate. In particular, greater effects can be expected by making the thickness of the photodiode substrate 310, which has a photoelectric conversion function, equal to or less than half the thickness of the scintillator 210. Furthermore, it is even better if the sum of the thicknesses of the photodiode substrate 310 and the counter circuit substrate 320 is thinner than the thickness of the scintillator 210.

[0023] As described above, in the radiation imaging device of the present disclosure, by making the thickness of the photodiode substrate 310 and the thickness of the counter circuit substrate 320 thinner than the thickness of the scintillator 210, it is possible to reduce the probability of absorption of X-ray photons 400 that have passed through the scintillator 210 and prevent image degradation. In other words, it is possible to provide a radiation imaging device with high image quality.

[0024] Second Embodiment An example of the schematic configuration of a radiation imaging apparatus 100 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 100, and Fig. 4 is a cross-sectional view. The operating principle of the radiation imaging apparatus 100 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.

[0025] As shown in Figures 3 and 4, the radiation imaging device 100 includes a scintillator section 200 and a photoelectric conversion section 300. The photoelectric conversion section 300 is composed of a single photoelectric conversion substrate 330, which has semiconductor layers constituting photoelectric conversion elements and circuits such as a signal processing circuit on the same substrate. That is, a pixel 302 has a first semiconductor layer 311, a first wiring structure 312, a second semiconductor layer 321, and a second wiring structure 322. 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.

[0026] 4, each semiconductor layer is arranged to correspond to the separated scintillator 210 (arranged in section S), but the second semiconductor layer 321, which does not contribute to photoelectric conversion, does not necessarily have to be arranged in section S of the scintillator 210. The second semiconductor layer 321 may be arranged to overlap, for example, the separation region 220. Furthermore, the scintillator section 200 and the photoelectric conversion section 300 may be overlapped with a shift so that the second semiconductor layer 321 corresponds to an adjacent section.

[0027] As described above, some of the X-ray photons 400 may pass through the scintillator section 200, but by making the thickness of the photoelectric conversion substrate 330 thinner than the thickness of the scintillator 210, the probability of the X-ray photons 400 being absorbed by the photoelectric conversion substrate 330 can be reduced. It is desirable to set the thickness of the photoelectric conversion substrate 330 to be equal to or less than half the thickness of the scintillator 210. Furthermore, since there is only one photoelectric conversion substrate 330, the thickness of the photoelectric conversion section 300 can be made thinner than in the first embodiment. As a result, the probability of the X-ray photons 400 being absorbed by the photoelectric conversion substrate 330 can be reduced.

[0028] As described above, in the radiation imaging device of the present disclosure, by making the thickness of the photoelectric conversion substrate 330 thinner than the thickness of the scintillator 210, the probability that the X-ray photons 400 that have passed through the scintillator 210 will be absorbed by the photoelectric conversion substrate 330 can be reduced, and image degradation can be prevented as in the first embodiment. In other words, it is possible to provide a radiation imaging device with high image quality.

[0029] Third Embodiment An example of the schematic configuration of a radiation imaging apparatus 100 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 100, and Fig. 6 is a cross-sectional view. The operating principle of the radiation imaging apparatus 100 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.

[0030] 5 and 6, the radiation imaging device 100 includes a scintillator section 200, a photoelectric conversion section 300, and a radiation absorption section 600. Similar to the second embodiment, the photoelectric conversion section 300 is configured from a photoelectric conversion substrate 330 in which a semiconductor layer having photoelectric conversion elements and circuits such as a signal processing section is formed on the same substrate.

[0031] In this embodiment, a radiation absorbing section 600 is disposed between the scintillator section 200 and the photoelectric conversion section 300. The radiation absorbing section 600 is preferably made of a material that absorbs X-rays and transmits visible light, such as a fiber optic plate (FOP). The radiation absorbing section 600 and the photoelectric conversion section 300 are disposed to overlap each other with an adhesive member 601 interposed therebetween, as shown in FIG. 6 . The adhesive member 601 may be made of the same material as the adhesive member 201.

[0032] As described above, some of the X-ray photons 400 may pass through the scintillator section 200, but the X-ray photons 400 that have passed through the scintillator section 200 are absorbed by the FOP, which is the radiation absorbing section 600. On the other hand, the visible light photons 500 generated in the scintillator section 200 pass through the FOP and are incident on the photoelectric conversion substrate 330.

[0033] As in the above embodiment, by making the thickness of the photoelectric conversion substrate 330 thinner than the thickness of the scintillator 210, the probability that the X-ray photons 400 will be absorbed by the photoelectric conversion substrate 330 can be reduced. Furthermore, because the X-ray photons 400 are absorbed by the radiation absorbing section 600 before they enter the photoelectric conversion substrate 330, the probability that the X-ray photons 400 will be absorbed by the photoelectric conversion substrate 330 can be further reduced. Since the thicker the radiation absorbing section 600, the greater the amount of X-ray photons 400 absorbed, it is desirable that the thicknesses of the photoelectric conversion substrate 330, the scintillator 210, and the radiation absorbing section 600 increase in this order. Typically, the X-ray absorption coefficient of an FOP is smaller than that of a scintillator material, so a greater effect can be expected by making the thickness of the radiation absorbing section 600 at least twice the thickness of the scintillator 210.

[0034] As described above, in the radiation imaging device of the present disclosure, by making the thickness of the photoelectric conversion substrate 330 thinner than the thickness of the scintillator 210, it is possible to reduce the probability that the X-ray photons 400 that have passed through the scintillator 210 will be absorbed by the photoelectric conversion substrate 330, and it is possible to prevent image degradation as in the first and second embodiments. In other words, it is possible to provide a radiation imaging device with high image quality.

[0035] Fourth Embodiment FIG. 7 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 the chest 221 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 viewed 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.

[0036] 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.

[0037] (Other Embodiments) The disclosure of this embodiment includes the following configurations, methods, and programs.

[0038] (Configuration 1) A radiation imaging device having a scintillator section having a plurality of scintillators that convert radiation into visible light, and a semiconductor substrate having photoelectric conversion elements that convert the visible light into an electrical signal, wherein the thickness of the semiconductor substrate is thinner than the thickness of the scintillator section.

[0039] (Configuration 2) The radiation imaging device according to configuration 1, wherein the thickness of the semiconductor substrate is equal to or less than half the thickness of the scintillator section.

[0040] (Configuration 3) The radiation imaging device according to Configuration 1 or 2, further comprising a radiation absorbing section between the scintillator section and the semiconductor substrate, wherein the thickness of the radiation absorbing section is greater than the thickness of the scintillator section and greater than the thickness of the semiconductor substrate.

[0041] (Configuration 4) The radiation imaging device according to configuration 3, wherein the thickness of the radiation absorbing section is at least twice the thickness of the scintillator section.

[0042] (Configuration 5) The radiation imaging apparatus according to configuration 3 or 4, wherein the radiation absorbing portion is an FOP.

[0043] (Configuration 6) The radiation imaging apparatus according to any one of configurations 1 to 5, wherein the photoelectric conversion element is an avalanche photodiode.

[0044] (Structure 7) The radiation imaging device described in any one of Structures 1 to 6, wherein the semiconductor substrate includes a photodiode substrate having a photoelectric conversion element that converts the visible light into an electrical signal, and a counter circuit substrate that processes the electrical signal, and the thickness of the photodiode substrate is 1 / 2 or less of the thickness of the scintillator section.

[0045] (Structure 8) A radiation imaging device according to any one of structures 1 to 7, characterized in that the plurality of scintillators are separated by separation regions, and the semiconductor substrate includes a plurality of pixel regions having photoelectric conversion elements corresponding to each of the plurality of scintillators.

[0046] (Configuration 9) A radiation imaging system comprising: a radiation source that generates radiation; and a radiation imaging device according to any one of configurations 1 to 8 that detects the radiation generated by the radiation source.

[0047] 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, the following claims are appended to apprise the public of the scope of the present invention.

[0048] This application claims priority based on Japanese Patent Application No. 2024-152992, filed September 5, 2024, the entire contents of which are incorporated herein by reference.

[0049] REFERENCE SIGNS LIST 100 Radiation imaging device 200 Scintillator section 201, 601 Adhesive member 210 Scintillator 220 Separation region 300 Photoelectric conversion section 301 Pixel 310 Photodiode substrate 311 First semiconductor layer 312 First wiring structure 320 Counter circuit board 321 Second semiconductor layer 322 Second wiring structure 330 Photoelectric conversion substrate 400, 711 Radiation or X-ray 500 Visible light 600 Radiation absorbing section 710 X-ray tube 720 Patient or subject 721 Chest of patient or chest of subject 730 Image processor (signal processing means) 740, 741 Display (display means) 750 Telephone line (transmission processing means) 760 Film processor (recording means) 761 Film (recording medium)

Claims

1. A radiation imaging device having a scintillator section having a plurality of scintillators that convert radiation into visible light, and a semiconductor substrate having photoelectric conversion elements that convert the visible light into an electrical signal, wherein the thickness of the semiconductor substrate is thinner than the thickness of the scintillator section.

2. The radiation imaging device according to claim 1, wherein the thickness of said semiconductor substrate is equal to or less than half the thickness of said scintillator portion.

3. The radiation imaging device according to claim 1, further comprising a radiation absorbing section between the scintillator section and the semiconductor substrate, wherein the thickness of the radiation absorbing section is greater than the thickness of the scintillator section and greater than the thickness of the semiconductor substrate.

4. The radiation imaging device according to claim 3, wherein the thickness of said radiation absorbing portion is at least twice the thickness of said scintillator portion.

5. The radiation imaging apparatus according to claim 3, wherein the radiation absorbing portion is an FOP.

6. The radiation imaging apparatus according to claim 1, wherein said photoelectric conversion element is an avalanche photodiode.

7. The radiation imaging device according to claim 1, characterized in that the semiconductor substrate includes a photodiode substrate having a photoelectric conversion element that converts the visible light into an electrical signal, and a counter circuit board that processes the electrical signal, and the thickness of the photodiode substrate is 1 / 2 or less the thickness of the scintillator section.

8. The radiation imaging device according to claim 1, wherein the plurality of scintillators are separated by separation regions, and the semiconductor substrate includes a plurality of pixel regions having photoelectric conversion elements corresponding to each of the plurality of scintillators.

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

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