Radiation imaging device and radiation imaging system

Abstract

To provide a radiation imaging apparatus that improves image quality (e.g., image sharpness or energy resolution) by suppressing the direct conversion of X-rays absorbed by a semiconductor substrate into electrical signals. [Solution] The radiation imaging apparatus of the present disclosure is a radiation imaging apparatus 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, wherein the thickness of the semiconductor substrate is thinner than the thickness of the scintillator section.

Description

Technical Field

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

Background Art

[0002] As an imaging device used for medical imaging diagnosis and non-destructive inspection by radiation (such as X-rays), a radiation imaging device using a flat panel detector (Flat Panel Detector, hereinafter FPD) formed of a semiconductor material is known. Such a radiation imaging device is used, for example, in medical imaging diagnosis to acquire still images, moving images, and the like.

[0003] Examples of FPDs include an integrating type sensor and a photon counting type sensor. The integrating type sensor measures the total amount of charges generated by the incidence of radiation. In contrast, the photon counting type sensor discriminates the energy (wavelength) of the incident radiation and counts the number of detections of radiation for each of a plurality of energy levels. That is, the photon counting type sensor has energy resolution. Therefore, in each field, an improvement in diagnostic ability by the photon counting method is expected.

[0004] In addition, the photon counting type sensor is classified into a direct type and an indirect type. The direct type counts the number of detections of radiation by directly detecting the energy of radiation using CdTe or the like. The indirect type detects the intensity of visible light generated by a scintillator due to the incidence of radiation and counts the number of detections of the visible light.

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

Prior Art Documents

Patent Documents

[0006]

Patent Document 1

[0007] Here, a radiographic imaging system using an FPD requires images with a wider range and higher sharpness compared to an X-ray CT scanner using an FPD. Therefore, the scintillator area of ​​a radiographic imaging system is larger than that of an X-ray CT scanner.

[0008] However, forming a scintillator with the same thickness as that of an X-ray CT scanner's scintillator, and with a larger surface area, is not easy from a cost or manufacturing perspective. For this reason, the scintillator thickness of a radiation imaging device is thinner than that of a CT scanner. Consequently, not all X-rays incident on the scintillator may be absorbed. X-rays that are not absorbed by the scintillator and pass through are absorbed by the semiconductor substrate on which the photoelectric conversion element or counter circuit is formed. Since the X-rays absorbed by the semiconductor substrate are directly converted into electrical signals without being converted into visible light, image quality (e.g., image sharpness or energy resolution) may be reduced.

[0009] Therefore, the purpose of this 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 electrical signals. [Means for solving the problem]

[0010] The radiation imaging apparatus of this disclosure comprises 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, wherein the thickness of the semiconductor substrate is thinner than the thickness of the scintillator section. [Effects of the Invention]

[0011] According to this disclosure, it is possible to provide a radiographic imaging device with improved image quality (e.g., image sharpness or energy resolution). [Brief explanation of the drawing]

[0012] [Figure 1] Schematic diagram of a radiation imaging device according to the first embodiment of this disclosure [Figure 2] Cross-sectional view of a radiation imaging device according to the first embodiment of this disclosure [Figure 3] Schematic diagram of a radiation imaging device according to a second embodiment of this disclosure [Figure 4] Cross-sectional view of a radiation imaging device according to a second embodiment of the present disclosure [Figure 5] Schematic diagram of a radiation imaging device according to the third embodiment of this disclosure [Figure 6] Cross-sectional view of a radiation imaging device according to a third embodiment of the present disclosure [Figure 7] Schematic diagram illustrating an example of the application of the radiographic imaging device of this disclosure to an X-ray diagnostic system. [Modes for carrying out the invention]

[0013] Embodiments of this disclosure will be described below with reference to the attached drawings. Similar elements throughout the various embodiments will be given the same reference numerals, and redundant descriptions will be omitted. Hereinafter, each embodiment of this disclosure will be described in the context of a radiation imaging device used in medical imaging 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] A schematic configuration example of the radiation imaging device 100 according to the first embodiment of this disclosure will be described with reference to Figures 1 and 2. Figure 1 is a schematic configuration diagram of the radiation imaging device 100, and Figure 2 is a cross-sectional view.

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

[0016] The scintillator unit 200 has a plurality of scintillators 210 separated by a separation region 220. By separating in the separation region 220, the spread of the visible light photons generated by the scintillator 210 can be suppressed. As the scintillator 210, known materials such as CsI:Tl or Gd2O2S (GOS) can be used.

[0017] The scintillator unit 200 and the photoelectric conversion unit 300 are arranged overlapping each other via an adhesive member 201 as shown in FIG. 2. As the adhesive member 201, an adhesive member having the property of melting or softening by heating can be used. The adhesive member 201 can use, for example, a sheet-like or liquid adhesive material (also called a hot melt resin) containing thermoplastic elastomers such as styrene-based, olefin-based, vinyl chloride-based, urethane-based, and amide-based. Also, as the adhesive member 201, an adhesive sheet such as an acrylic-based or silicone-based adhesive sheet having an adhesive function at room temperature can be used.

[0018] The photodiode substrate 310 and the counter circuit substrate 320 are laminated and electrically connected. The photodiode substrate 310 is provided with a pixel region 311 in which pixels 301 having a sensor function are arranged in a matrix. Therefore, the photodiode substrate 310 may also be referred to as a sensor member, a sensor substrate, a sensor chip, etc. The counter circuit substrate 320 is provided with a circuit region 321 that processes the signals detected in the pixel region 311. The counter circuit substrate 320 may also be referred to as a circuit member, a circuit substrate, a circuit chip, etc. Also, each substrate is a semiconductor substrate using a silicon wafer as a material.

[0019] The photodiode substrate 310 has a first semiconductor layer 311 that forms a photodiode, which is a photoelectric conversion element, and a first wiring structure 312, and from these, the pixel 301 is constituted. As described above, in the scintillator unit 200 of the present embodiment, since the scintillator 210 is separated by the separation region 220, the volume of the scintillator 210 becomes smaller than that of the scintillator unit where the separation region 220 is not arranged. Thereby, since the amount of light emission in the scintillator 210 becomes smaller, an avalanche photodiode having a function of amplifying a signal is suitable as the photoelectric conversion element. The counter circuit substrate 320 has a second semiconductor layer 321 and a second wiring structure 322 that constitute a circuit such as a signal processing unit so as to correspond to the pixel 301.

[0020] The photoelectric conversion unit 300 is a back-illuminated type photoelectric conversion unit in which visible light photons 500 are incident from a second surface facing a first surface on which the first wiring structure 312 is formed, and the counter circuit substrate 320 is arranged on the first surface side. Note that the invention according to each embodiment can also be applied to a photoelectric conversion unit having a front-illuminated type structure.

[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; 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, but if the thickness of the scintillator cannot be sufficiently secured, the amount of X-ray photons 400 that pass through the scintillator increases. In addition, although the energy of the X-rays is set to be suitable for imaging, high-energy X-ray photons may be mixed in due to variations in the X-ray generator, and these may pass through the scintillator section 200. X-ray photons 400 that pass through the scintillator section 200 are incident on the photoelectric conversion section 300 and then incident on the photodiode substrate 310 or the counter circuit board 320. When the substrate material is silicon, X-ray photons 400 are directly converted into electrical signals, and this can lead to image degradation due to false detections when the signals are read out. For example, if X-ray photons 400 are detected in an adjacent pixel to the incident pixel, it can cause image blurring, and if they are detected in the same pixel while measuring the emission of a scintillator, errors may occur in the output value or energy resolution.

[0022] In this embodiment, the thickness of the photodiode substrate 310 and the counter circuit board 320 is thinner than the thickness of the scintillator 210. Since the X-ray absorption coefficient of silicon is normally smaller than that of the scintillator material, the probability of X-ray photons 400 being absorbed by each substrate can be reduced by making the photodiode substrate 310 and the counter circuit board 320 thinner. In particular, a greater effect can be expected by making the thickness of the photodiode substrate 310, which has a photoelectric conversion function, 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 board 320 is thinner than the thickness of the scintillator 210.

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

[0024] [Second Embodiment] A schematic configuration example of the radiation imaging device 100 according to the second embodiment of this disclosure will be described with reference to Figures 3 and 4. Figure 3 is a schematic configuration diagram of the radiation imaging device 100, and Figure 4 is a cross-sectional view. The operating principle of the radiation imaging device 100 is the same as that of the first embodiment, so a description will be omitted. The differences from the first embodiment will be described in detail below.

[0025] As shown in Figures 3 and 4, the radiation imaging device 100 comprises a scintillator unit 200 and a photoelectric conversion unit 300. The photoelectric conversion unit 300 is composed of a single photoelectric conversion substrate 330, and the photoelectric conversion substrate 330 has semiconductor layers that constitute the photoelectric conversion element and circuits such as the signal processing unit 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. Since one pixel is formed by the photoelectric conversion element and the signal processing circuit, the area of ​​the photoelectric conversion element in this embodiment is smaller than that of the photoelectric conversion element in the first embodiment. For this reason, an avalanche photodiode having the function of amplifying signals is suitable for the photoelectric conversion element.

[0026] In Figure 4, each semiconductor layer is arranged to correspond to the isolated scintillator 210 (placed within the compartment S), but the second semiconductor layer 321, which does not contribute to photoelectric conversion, does not necessarily need to be placed within the compartment S of the scintillator 210. The second semiconductor layer 321 may be arranged to overlap, for example, the isolation region 220. Furthermore, the scintillator section 200 and the photoelectric conversion section 300 may be offset and overlapped so that the second semiconductor layer 321 corresponds to the adjacent compartment.

[0027] As mentioned 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 X-ray photons 400 being absorbed by the photoelectric conversion substrate 330 can be reduced. It is desirable that the thickness of the photoelectric conversion substrate 330 be 1 / 2 times or less the thickness of the scintillator 210. Also, 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 X-ray photons 400 being absorbed by the photoelectric conversion substrate 330 can be reduced.

[0028] As described above, the radiation imaging apparatus of this disclosure reduces the probability that X-ray photons 400 that have passed through the scintillator 210 are absorbed by the photoelectric conversion substrate 330 by making the thickness of the photoelectric conversion substrate 330 thinner than the thickness of the scintillator 210, thereby preventing image degradation as in the first embodiment. In other words, it becomes possible to provide a radiation imaging apparatus with high image quality.

[0029] [Third Embodiment] A schematic configuration example of the radiation imaging apparatus 100 according to the third embodiment of this disclosure will be described with reference to Figures 5 and 6. Figure 5 is a schematic configuration diagram of the radiation imaging apparatus 100, and Figure 6 is a cross-sectional view. The operating principle of the radiation imaging apparatus 100 is the same as that of the first embodiment, so a description will be omitted. The differences from the above embodiment will be described in detail below.

[0030] As shown in Figures 5 and 6, the radiation imaging device 100 comprises a scintillator unit 200, a photoelectric conversion unit 300, and a radiation absorption unit 600. The photoelectric conversion unit 300 is composed of a photoelectric conversion substrate 330 in which a semiconductor layer having a photoelectric conversion element and a circuit such as a signal processing unit is formed on the same substrate, similar to the second embodiment.

[0031] In this embodiment, a radiation absorbing section 600 is positioned 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 FOP (Fiber Optic Plate). Furthermore, the radiation absorbing section 600 and the photoelectric conversion section 300 are arranged on top of each other via an adhesive member 601, as shown in Figure 6. The adhesive member 601 can be made of the same material as the adhesive member 201.

[0032] As mentioned above, some of the X-ray photons 400 may pass through the scintillator section 200, but the X-ray photons 400 that pass through the scintillator section 200 are absorbed by the FOP, which is the radiation absorption 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] Similar to the above embodiment, by making the thickness of the photoelectric conversion substrate 330 thinner than the thickness of the scintillator 210, the probability of X-ray photons 400 being absorbed by the photoelectric conversion substrate 330 can be reduced. Furthermore, since the X-ray photons 400 are absorbed by the radiation absorption section 600 before they enter the photoelectric conversion substrate 330, the probability of X-ray photons 400 being absorbed by the photoelectric conversion substrate 330 can be further reduced. As the amount of X-ray photons 400 absorbed increases with the thickness of the radiation absorption section 600, it is desirable that the thicknesses be in the order of photoelectric conversion substrate 330, scintillator 210, and radiation absorption section 600. Typically, the X-ray absorption coefficient of FOP is smaller than that of the scintillator material, so a greater effect can be expected by making the thickness of the radiation absorption section 600 more than twice the thickness of the scintillator 210.

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

[0035] [Fourth Embodiment] Figure 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, as radiation generated in the X-ray tube 710 (radiation source), pass through the chest 221 of the patient or subject 720 and enter the radiation imaging device 100 of the present disclosure, which includes a scintillator unit 200. These incident X-rays contain information about the inside of the patient's body. The scintillator unit 200 emits light in response to the incident X-rays, and this is photoelectrically converted to obtain electrical information. This information is converted into a digital signal and processed as an image by an image processor 730, which is a signal processing means, and can be observed on a display 740, which is a display means in the control room (X-ray room). The radiation imaging system comprises at least a radiation imaging device 100 and an image processor 730 that processes signals from the radiation imaging device 100.

[0036] Furthermore, images processed by the image processor 730 can be transmitted to a remote location (for example, a doctor's room) via a transmission means such as a telephone line 750. The transmitted images can be displayed on a display 741, which is a display means in the doctor's room, or saved on a recording means such as an optical disc, so that a doctor in a remote location can make a diagnosis using the transmitted images. In addition, the transmitted images can also be recorded on a recording medium, film 761, by a recording means, film processor 760.

[0037] (Other embodiments) This embodiment includes the following configurations, methods, and programs.

[0038] (Composition 1) A radiation imaging apparatus comprising a scintillator section having multiple 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, A radiation imaging apparatus characterized in that the thickness of the semiconductor substrate is thinner than the thickness of the scintillator portion.

[0039] (Configuration 2) The radiation imaging apparatus according to configuration 1, characterized in that the thickness of the semiconductor substrate is 1 / 2 times or less the thickness of the scintillator portion.

[0040] (Composition 3) The scintillator portion and the semiconductor substrate further have a radiation absorbing portion, The radiation imaging apparatus according to configuration 1 or 2, characterized in that the thickness of the radiation absorbing portion is greater than the thickness of the scintillator portion and greater than the thickness of the semiconductor substrate.

[0041] (Composition 4) The radiation imaging apparatus according to configuration 3, characterized in that the thickness of the radiation absorbing portion is twice or more the thickness of the scintillator portion.

[0042] (Composition 5) The radiation imaging apparatus according to configuration 3 or 4, characterized in that the radiation absorbing part is an FOP.

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

[0044] (Composition 7) The semiconductor substrate includes a photodiode substrate having a photoelectric conversion element that converts visible light into an electrical signal, and a counter circuit substrate that processes the electrical signal. The radiation imaging apparatus according to any one of configurations 1 to 6, characterized in that the thickness of the photodiode substrate is 1 / 2 times or less the thickness of the scintillator portion.

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

[0046] (Composition 9) A radiation source that emits radiation, A radiation imaging system characterized by comprising a radiation imaging device according to any one of configurations 1 to 8 for detecting radiation generated by the aforementioned radiation source. [Explanation of Symbols]

[0047] 100 Radiation imaging device 200 Scintillator section 201,601 Adhesive material 210 Scintillator 220 Separation area 300 Photoelectric conversion unit 301 pixels 310 Photodiode board 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-rays 500 visible light 600 Radiation absorption section 710 X-ray tube 720 patients or subjects 721 The patient's chest or the subject's chest 730 Image processor (signal processing means) 740,741 Display (display means) 750 Telephone line (transmission processing means) 760 Film Processor (Recording Method) 761 Film (recording medium)

Claims

1. A radiation imaging apparatus comprising a scintillator section having multiple 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, A radiation imaging apparatus characterized in that the thickness of the semiconductor substrate is thinner than the thickness of the scintillator portion.

2. The radiation imaging apparatus according to claim 1, characterized in that the thickness of the semiconductor substrate is 1 / 2 times or less the thickness of the scintillator portion.

3. The scintillator portion and the semiconductor substrate further have a radiation absorbing portion, The radiation imaging apparatus according to claim 1, characterized in that the thickness of the radiation absorbing portion is greater than the thickness of the scintillator portion and greater than the thickness of the semiconductor substrate.

4. The radiation imaging apparatus according to claim 3, characterized in that the thickness of the radiation absorbing portion is twice or more the thickness of the scintillator portion.

5. The radiation imaging apparatus according to claim 3, characterized in that the radiation absorbing part is an FOP.

6. The radiation imaging apparatus according to claim 1, characterized in that the photoelectric conversion element is an avalanche photodiode.

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

8. The plurality of scintillators are separated by separation regions, The radiation imaging apparatus according to claim 1, characterized in that the semiconductor substrate includes a plurality of pixel regions having photoelectric conversion elements, each corresponding to one of the plurality of scintillators.

9. A radiation source that emits radiation, A radiation imaging system comprising a radiation imaging device according to any one of claims 1 to 8, which detects radiation generated by the aforementioned radiation source.

Images