阅读说明：本技术 放射线摄像装置和放射线摄像系统 (Radiation imaging apparatus and radiation imaging system ) 是由 石井孝昌 照井晃介 西部航太 保科智启 于 2020-11-02 设计创作，主要内容包括：本发明提供一种放射线摄像装置和放射线摄像系统。一种放射线摄像装置包括：第一闪烁体层,其被构造为将进入第一闪烁体层的放射线(R)转换成光；第二闪烁体层,其被构造为将透射过第一闪烁体层的放射线转换成光；配设在第一闪烁体层与第二闪烁体层之间的光纤板(FOP)；以及摄像部分,其被构造为将在第一闪烁体层中产生的光和在第二闪烁体层中产生的光转换为电信号。(The invention provides a radiation imaging apparatus and a radiation imaging system. A radiation imaging apparatus includes: a first scintillator layer configured to convert radiation (R) entering the first scintillator layer into light; a second scintillator layer configured to convert the radiation transmitted through the first scintillator layer into light; a Fiber Optic Plate (FOP) disposed between the first scintillator layer and the second scintillator layer; and an image pickup portion configured to convert the light generated in the first scintillator layer and the light generated in the second scintillator layer into electrical signals.)

1. A radiation imaging apparatus comprising:

a first scintillator layer configured to convert radiation entering the first scintillator layer into light;

a second scintillator layer configured to convert the radiation transmitted through the first scintillator layer into light;

a fiber plate disposed between the first scintillator layer and the second scintillator layer; and

an image pickup portion configured to convert light generated in the first scintillator layer and light generated in the second scintillator layer into electric signals.

2. The radiographic imaging apparatus according to claim 1, wherein the numerical aperture NA of the optical fiber plate is less than 1.0.

3. The radiographic imaging apparatus according to claim 1, wherein the optical fiber plate has a thickness of 1.0mm or less.

4. The radiation imaging apparatus according to claim 1, wherein the first scintillator layer and the second scintillator layer are each made of CsI: Tl.

5. The radiation imaging apparatus according to claim 1,

wherein the first scintillator layer is made of GOS, and

wherein the second scintillator layer is made of CsI: Tl.

6. The radiographic imaging apparatus according to claim 1, wherein the first scintillator layer and the second scintillator layer are different in thickness.

7. The radiation imaging apparatus according to claim 1, wherein the first scintillator layer, the optical fiber plate, the second scintillator layer, and the imaging portion are arranged in this order from an incident side of radiation.

8. The radiographic imaging apparatus according to claim 7, further comprising a moisture-proof resin between the optical fiber plate and the imaging portion and on a side surface of the second scintillator layer.

9. The radiation imaging apparatus according to claim 7, further comprising:

a support substrate configured to support the first scintillator layer on a radiation incident side of the first scintillator layer;

a first moisture-proof resin between the support substrate and the optical fiber plate and on a side surface of the first scintillator layer; and

a second moisture-proof resin between the optical fiber plate and the image pickup portion and on a side surface of the second scintillator layer.

10. A radiation imaging system comprising:

the radiation imaging apparatus according to claim 1;

a signal processing unit configured to process the electric signal obtained by the image pickup section;

a recording unit configured to record the electric signal processed by the signal processing unit;

a display unit configured to display the electric signal processed by the signal processing unit;

a transmission unit configured to transmit the electrical signal processed by the signal processing unit; and

a radiation generating unit configured to generate radiation.

Technical Field

The present disclosure relates to a radiation imaging apparatus and a radiation imaging system configured to perform imaging using radiation, and more particularly to a radiation imaging apparatus and a radiation imaging system suitable for use in, for example, a medical image diagnosis apparatus and an analysis apparatus.

Background

The radiation imaging apparatus generally includes: a scintillator (phosphor) layer configured to convert radiation entering the scintillator layer into light (for example, visible light) having a certain wavelength that can be detected by the photoelectric conversion elements; and an image pickup section including a photoelectric conversion element configured to convert light generated in the scintillator layer into an electric signal. For application to medical image diagnosis, such a radiation imaging apparatus with high sensitivity is desired to reduce radiation exposure of a patient. One example of a method for realizing a radiation imaging apparatus having high sensitivity is to increase the film thickness of a scintillator layer.

When the film thickness of the scintillator layer is increased, it is considered that light converted by the scintillator layer is scattered in the scintillator layer, thereby reducing the sharpness of the radiographic image. Therefore, the scintillator used for the scintillator layer is desired to be a columnar crystal having high light directivity. Materials for this type of scintillator include cesium iodide (CsI) obtained by doping CsI with thallium (Tl). In addition, a structure having a Fiber Optic Plate (FOP) is disposed between the image pickup portion and the scintillator layer to ensure the sharpness of a radiographic image while preventing radiation degradation of the photoelectric conversion element. For example, in japanese patent application laid-open No. 2011-158291, a scintillator plate having a scintillator with columnar crystals is formed on FOP formed by bundling a plurality of optical fibers. Japanese patent laid-open No. 2016-136094 describes a scintillator panel in which an FOP and a scintillator are bonded together.

In the scintillator plates disclosed in japanese patent application laid-open nos. 2011-158291 and 2016-136094, scintillators made of columnar crystals are used. Therefore, light scattering in the scintillator layer is small. However, there are gaps between the columnar crystals in the scintillator layer and the columnar crystals, i.e., an air layer, and it is therefore difficult to confine light completely in the columnar crystals. Therefore, even when a scintillator made of columnar crystals is used, as the film thickness of the scintillator layer becomes larger, the light scattering range in the scintillator layer becomes wider. That is, light converted by the scintillator layer whose film thickness is increased is scattered in the scintillator layer before reaching the FOP, and therefore, even when the FOP has a function of causing light to travel straight, the sharpness of the radiographic image is lowered.

Disclosure of Invention

An object of the present disclosure provided in view of such a problem is to provide a mechanism for realizing a radiation imaging apparatus having high sensitivity and suppressing a reduction in the sharpness of a radiation image.

According to the present disclosure, a radiation imaging apparatus is provided. The radiation imaging apparatus includes: a first scintillator layer configured to convert radiation entering the first scintillator layer into light; a second scintillator layer configured to convert the radiation transmitted through the first scintillator layer into light; a fiber plate disposed between the first scintillator layer and the second scintillator layer; and an image pickup portion configured to convert the light generated in the first scintillator layer and the light generated in the second scintillator layer into electrical signals. The present disclosure also provides a radiation imaging system including the above radiation imaging apparatus.

Additional features of the invention will become apparent from the following description of exemplary embodiments with reference to the accompanying drawings.

Drawings

Fig. 1 is a perspective view for illustrating an example of a schematic configuration of a radiation imaging apparatus according to a first embodiment of the present invention.

Fig. 2 is a cross-sectional view for illustrating an example of a detailed configuration of the radiation imaging apparatus according to the first embodiment of the present invention.

Fig. 3 is a perspective view for illustrating an example of a schematic configuration of a radiation imaging apparatus according to a second embodiment of the present invention.

Fig. 4 is a cross-sectional view for illustrating an example of a detailed configuration of a radiation imaging apparatus according to a second embodiment of the present invention.

Fig. 5 is a perspective view for illustrating an example of a schematic configuration of a radiation imaging apparatus according to a third embodiment of the present invention.

Fig. 6 is a cross-sectional view for illustrating an example of a detailed configuration of a radiation imaging apparatus according to a third embodiment of the present invention.

Fig. 7 is a conceptual diagram of an X-ray imaging system (radiation imaging system) according to a fourth embodiment of the present invention, which uses the radiation imaging apparatus according to any one of the first to third embodiments.

Detailed Description

A mode for carrying out the present embodiment is described with reference to the accompanying drawings. In addition, when the radiation imaging apparatus is used in, for example, a medical image diagnosis apparatus and an analysis apparatus, light includes visible light and infrared light, and radiation includes X-rays, α -rays, β -rays, and γ -rays.

First embodiment

Fig. 1 is a perspective view for illustrating an example of a schematic configuration of a radiation imaging apparatus 100 according to a first embodiment of the present invention. In fig. 1, an xyz coordinate system is shown in which the incident direction of the radiation ray R is set to the z direction, and the x direction and the y direction perpendicular to each other are perpendicular to the z direction.

Fig. 2 is a cross-sectional view for illustrating an example of a detailed configuration of the radiation imaging apparatus 100 according to the first embodiment of the present invention. In fig. 2, an xyz coordinate system corresponding to the xyz coordinate system shown in fig. 1 is shown; more specifically, fig. 2 illustrates a detailed configuration of the radiation imaging apparatus 100 according to the first embodiment in a plane defined by the x direction and the z direction illustrated in fig. 1. In fig. 2, the same components as those of the configuration shown in fig. 1 are denoted by the same reference numerals.

In the following description of the first embodiment, the radiation imaging apparatus 100 according to the first embodiment shown in fig. 1 and 2 is described as "radiation imaging apparatus 100-1".

As shown in fig. 1 and 2, the radiation imaging apparatus 100-1 includes a first scintillator plate (first phosphor plate) 130, a second scintillator plate (second phosphor plate) 120, and an imaging portion 110. In addition, as shown in fig. 2, the radiation imaging apparatus 100-1 further includes a first bonding member 141 and a second bonding member 142, and a first moisture-proof resin 151 and a second moisture-proof resin 152. In fig. 1, the components are shown to be spaced apart from each other for convenience of description, but as shown in fig. 2, the components are actually arranged by being stacked via the first and second engaging members 141 and 142.

As shown in fig. 2, the first scintillator plate 130 includes a first scintillator layer 131, a reflective layer 132, and a support substrate 133.

The first scintillator layer 131 is a phosphor configured to convert radiation R (including radiation R transmitted through the inspection object H) entering the first scintillator layer 131 through the support substrate 133 and the reflective layer 132 into light having a certain wavelength detectable by the photoelectric conversion elements 112 of the imaging section 110. The first scintillator layer 131 is made of, for example, CsI: Tl. In this case, for example, the first scintillator layer 131 is formed on the support substrate 133 provided with the reflective layer 132 by a vapor deposition method.

The reflective layer 132 shown in fig. 2 is a layer for reflecting light (which may include light generated in the second scintillator layer 121) generated in the first scintillator layer 131, which enters the reflective layer 132, in the z direction to guide the light to the photoelectric conversion elements 112 of the image pickup section 110. The reflective layer 132 is not always required in the first embodiment.

The support substrate 133 is a substrate that is arranged on a radiation incident side of the first scintillator layer 131, which the radiation R enters, and is configured to support the first scintillator layer 131. Examples of materials that can be used for the support substrate 133 include glass, amorphous carbon, CFRP, resin film, aluminum, and titanium.

In this case, when the support substrate 133 is made of aluminum, titanium, or other material having a function of reflecting light, the reflective layer 132 is not always required. The CsI: Tl has deliquescence, and therefore it is desirable that the first scintillator layer 131 made of the CsI: Tl is covered with the support substrate 133 and the moisture-proof protective film. In view of moisture resistance, it is also desirable to provide the first moisture-resistant resin 151 in the outer peripheral portion of the radiographic imaging apparatus 100-1. Specifically, the first moisture-proof resin 151 is disposed between the support substrate 133 and the fiber plate (FOP)122 of the second scintillator plate 120 on the side surface of the first scintillator layer 131. Fig. 2 shows an exemplary configuration in which the first scintillator layer 131 is in contact with the reflective layer 132. However, in order to protect the reflective layer 132 from CsI: Tl, which is a material for forming the first scintillator layer 131, an organic film may also be formed between the first scintillator layer 131 and the reflective layer 132.

As shown in fig. 2, the second scintillator plate 120 is connected to the first scintillator plate 130 via a first bonding member 141, and is also connected to the image pickup portion 110 via a second bonding member 142. As shown in fig. 1 and 2, the second scintillator panel 120 includes a second scintillator layer 121 and a fiber optic plate (hereinafter, simply referred to as "FOP") 122.

The second scintillator layer 121 is a phosphor configured to convert radiation R transmitted through the first scintillator plate 130 including the first scintillator layer 131 and the FOP 122. The radiation R that is transmitted through the inspection object H and converted into light having a certain wavelength can be detected by the photoelectric conversion element 112 of the imaging portion 110. The second scintillator layer 121 is made of, for example, CsI: Tl. In this case, the second scintillator layer 121 is formed on the FOP 122 by, for example, a vapor deposition method. In addition, as described above, CsI: Tl has deliquescence, and therefore it is desirable that the second scintillator layer 121 made of CsI: Tl is covered with the FOP 122 and a moisture-proof protective film (not shown). Organic films, such as parylene, may be used as moisture-proof protective films. In view of moisture resistance, it is also desirable to provide the second moisture resistant resin 152 in the outer peripheral portion of the radiation imaging apparatus 100-1. Specifically, the second moisture-proof resin 152 is disposed between the FOP 122 and the imaging substrate 111 of the imaging section 110 and on the side surface of the second scintillator layer 121. Fig. 2 shows a configuration example in which the second scintillator layer 121 is in contact with the FOP 122. However, in order to secure adhesion strength or prevent the columnar crystals of the scintillator from being disturbed, an organic film may be formed between the second scintillator layer 121 and the FOP 122.

The FOP 122 may include a fiber plate formed by bundling a plurality of optical fibers between the first scintillator layer 131 and the second scintillator layer 121. As the numerical aperture NA of the FOP 122 decreases, oblique light entering the FOP 122 may be blocked. That is, the FOP 122 can limit the incident angle of incident light by the numerical aperture NA. In this embodiment, the numerical aperture NA of the FOP 122 can be less than about 1.0. The FOP 122 also has a function of blocking the radiation R, and provides a greater shielding effect as the thickness becomes larger. The radiation imaging apparatus 100-1 according to the present embodiment employs a mode in which the radiation R transmitted through the first scintillator layer 131 (not absorbed by the first scintillator 131) is absorbed by the second scintillator layer 121 to be converted into light. Thus, in this embodiment, the FOP 122 may be 1.0mm or less in thickness. In this case, the thickness of the FOP 122 may be set to 0.5mm in consideration of the function of the FOP 122 serving as the support substrate of the second scintillator layer 121.

The image pickup portion 110 converts light generated in the first scintillator layer 131 and light generated in the second scintillator layer 121 into electric signals. As shown in fig. 1 and 2, the image pickup section 110 includes an image pickup substrate 111 and a photoelectric conversion element 112.

A plurality of photoelectric conversion elements 112 arranged in a matrix are formed on the image pickup substrate 111. The photoelectric conversion element 112 detects incident light (light generated in the first scintillator layer 131 and light generated in the second scintillator layer 121), and converts the incident light into an electric signal. Examples of the photoelectric conversion element 112 that can be used include a PIN type sensor and a MIS type sensor each using amorphous silicon.

For the first joining member 141 and the second joining member 142, joining members that can be melted or softened by heating may be used. Each of the first joining member 141 and the second joining member 142 is formed of a sheet-like or liquid joining material (also referred to as "hot melt resin") containing a thermoplastic elastomer such as a styrene-based, olefin-based, vinyl chloride-based, urethane-based, or amide-based. For each of the first and second coupling members 141 and 142, an acrylic-based or silicon-based adhesive sheet having an adhesive function at room temperature, for example, may also be used.

In addition, the first moisture-proof resin 151 and the second moisture-proof resin 152 are disposed in the outer peripheral portion of the radiation imaging apparatus 100-1 in consideration of moisture-proof property. The first and second moisture-proof resins 151 and 152 may be made of, for example, silicone resin, acrylic resin, epoxy resin, urethane resin, or other resins.

As shown in fig. 1 and 2, the radiation imaging apparatus 100-1 includes a support substrate 133, a first scintillator layer 131, a FOP 122, a second scintillator layer 121, and an imaging portion 110, which are arranged in this order from the incident side of radiation R. In this case, the first scintillator layer 131 and the second scintillator layer 121 have different thicknesses.

Radiation R emitted for exposure toward the inspection object H in the direction indicated by the arrow in fig. 1 and 2 is attenuated by the inspection object H and then enters the first scintillator layer 131 and the second scintillator layer 121. The first scintillator layer 131 and the second scintillator layer 121 each convert incident radiation R into light (for example, visible light) having a certain wavelength that can be detected by the photoelectric conversion elements 112. Then, the light converted by each of the first scintillator layer 131 and the second scintillator layer 121 enters the photoelectric conversion element 112 formed on the imaging substrate 111 to be converted into an electric signal, and a radiographic image is generated based on the electric signal. By repeating this operation, the radiation imaging apparatus 100-1 can also obtain a moving image relating to a radiation image. The sharpness of the radiographic image is described below. When only one scintillator layer 131 is provided (the scintillator layer 121 is not provided), the radiation R transmitted through the inspection object H is converted into visible light by the scintillator layer 131. When the converted light is made to travel straight to the photoelectric conversion element 112, a radiographic image with high definition can be obtained. Tl is a columnar crystal; further, since light scattering in the scintillator layer 131 is small, it has high light directivity. However, the void in the form of an air layer between the columnar crystals (the void is an air layer) makes it difficult to completely confine light in the columnar crystals. Therefore, the light converted in the vicinity of the surface of the scintillator layer 131 into which the radiation R enters is diffused and propagated toward the surface of the scintillator layer 131 located on the photoelectric conversion element 112 side from which the light exits, while being repeatedly scattered. Meanwhile, a part of the light converted in the vicinity of the surface of the scintillator layer 131 that emits the light diffuses and propagates toward the surface side of the scintillator layer 131 that the radiation R enters while being repeatedly scattered, and is reflected by the reflection layer 132 to return to the emission surface of the scintillator layer 131 while being further diffused. As the thickness of the scintillator layer 131 becomes larger, the range of light diffusion becomes wider. Therefore, even when the FOP 122 is used, as the thickness of the scintillator layer 131 becomes larger, the proportion of diffused light included in light entering the FOP 122 increases, and the sharpness of the radiographic image decreases.

In view of this, in the first embodiment, in two scintillator layers as scintillator layers: the FOP 122 is disposed between the first scintillator layer 131 and the second scintillator layer 121. In the first embodiment, two scintillator layers are provided in this way, and therefore the thickness of each of the first scintillator layer 131 and the second scintillator layer 121 can be set so that the total thickness of the first scintillator layer 131 and the thickness of the second scintillator layer 121 is equal to or larger than the thickness of one scintillator layer required so far. The thickness of each of the first scintillator layer 131 and the second scintillator layer 121 can be made smaller (thinner) than the thickness of one scintillator layer required hitherto, and therefore the light diffusion range in each of the first scintillator layer 131 and the second scintillator layer 121 can be reduced. The thicknesses of the first scintillator layer 131 and the second scintillator layer 121 may be set to be unequal to each other (may be set to be different). In this case, the first scintillator layer 131 located on the incident side of the radiation R mainly converts low-energy radiation R into light, and the second scintillator layer 121 located on the photoelectric conversion element 112 side mainly converts high-energy radiation R transmitted through the first scintillator layer 131 and the FOP 122 into light. Therefore, it is desirable to determine the thickness of each of the first scintillator layer 131 and the second scintillator layer 121 in accordance with the property of the radiation R to be used. Now, the sharpness of the radiographic image is described in more detail as follows. The radiation R absorbed by the first scintillator layer 131 is converted into light that enters the second scintillator layer 121 through the FOP 122. The thickness of the first scintillator layer 131 is smaller (thinner) than that of the scintillator in the single scintillator layer structure. Therefore, the proportion of diffused light included in light entering the FOP 122 is small. In addition, some of the light converted by the first scintillator layer 131 propagates toward the reflective layer 132. This light is reflected by the reflective layer 132 to enter the FOP 122, but since the thickness of the first scintillator layer 131 is small (thin), the diffusion range is smaller. That is, light transmitted through the FOP 122 to enter the second scintillator layer 121 has high definition. The radiation R that has not been absorbed by the first scintillator layer 131 transmits the FOP 122 to enter the second scintillator layer 121. The radiation R is converted into light entering the photoelectric conversion elements 112 by the second scintillator layer 121. The thickness of the second scintillator layer 121 is also smaller (thinner) than that of the scintillator layer in the single scintillator layer structure, and therefore the proportion of diffused light included in light entering the photoelectric conversion elements 112 is small. In addition, some of the light converted by the second scintillator layer 121 travels toward the FOP 122. A part of the light is reflected by the FOP 122 to enter the photoelectric conversion element 112. Since the thickness of the second scintillator layer 121 is small (thin), the diffusion range is small. In addition, a part of the light entering the FOP 122 passes through the first scintillator layer 131 to be reflected by the reflective layer 132 to return along the same path as described above.

As described above, in the radiation imaging apparatus 100-1, two scintillator layers are disposed: the first scintillator layer 131 and the second scintillator layer 121 with the FOP 122 sandwiched therebetween. The thickness of each of the first scintillator layer 131 and the second scintillator layer 121 can be made smaller (thinner) than the thickness of the single scintillator layer. In this way, by setting the total thickness of the first scintillator layer 131 and the thickness of the second scintillator layer 121 to be equal to or greater than a predetermined thickness, a radiation imaging apparatus having high sensitivity can be realized, and it is also possible to reduce the light diffusion range in each of the first scintillator layer 131 and the second scintillator layer 121 and to reduce the proportion of diffused light included in light entering the photoelectric conversion element 112, so that it is possible to suppress a decrease in the sharpness of a radiographic image.

Further, the FOP 122 having a numerical aperture NA smaller than 1.0 is disposed between the first scintillator layer 131 and the second scintillator layer 121, and thus the incident angle of light entering both surfaces (upper surface and lower surface) of the FOP 122 can be restricted. That is, by disposing the FOP 122 between the first scintillator layer 131 and the second scintillator layer 121, diffused light of diffused light close to the linearly propagating light in the middle of the scintillator layer group (the first scintillator layer 131 and the second scintillator layer 121) can be corrected.

Second embodiment

In the following description of the second embodiment, the description of the same matters as those of the above-described first embodiment is omitted, and matters different from those of the above-described first embodiment are described.

Fig. 3 is a perspective view for illustrating an example of a schematic configuration of the radiation imaging apparatus 100 according to the second embodiment. In fig. 3, an xyz coordinate system corresponding to the xyz coordinate system shown in fig. 1 and 2 is shown. Further, in fig. 3, the same components as those of the configuration shown in fig. 1 and 2 are denoted by the same reference numerals, and detailed description thereof is omitted.

Fig. 4 is a cross-sectional view for illustrating an example of a detailed configuration of the radiation imaging apparatus 100 according to the second embodiment. In fig. 4, an xyz coordinate system corresponding to the xyz coordinate system shown in fig. 3 is shown, and more specifically, fig. 4 is an illustration of a detailed configuration of the radiation imaging apparatus 100 according to the second embodiment in a plane defined by the x direction and the z direction shown in fig. 3. In fig. 4, the same components as those of the configuration shown in fig. 3 are denoted by the same reference numerals.

In the following description of the second embodiment, the radiation imaging apparatus 100 shown in fig. 3 and 4 is described as "radiation imaging apparatus 100-2". Differences from the above-described first embodiment are described below.

As shown in fig. 3 and 4, the radiation imaging apparatus 100-2 includes a first scintillator plate 130, a second scintillator plate 120, an imaging portion 110, and a fixed substrate 160. In addition, as shown in fig. 4, the radiation imaging apparatus 100-2 further includes first and second bonding members 141 and 142, first and second moisture-proof resins 151 and 152, and a fixing member 170. In fig. 3, these components are shown to be spaced apart from each other for convenience of description, but as shown in fig. 4, these components are actually arranged by being stacked via the first and second engagement members 141 and 142 and the fixing member 170.

As shown in fig. 3 and 4, the radiation imaging apparatus 100-2 includes a plurality of imaging sections 110. Each image pickup portion 110 includes one of a plurality of photoelectric conversion elements 112 arranged in a matrix on an image pickup substrate 111, and is configured to detect light generated in the first scintillator layer 131 and light generated in the second scintillator layer 121 to convert the light into an electric signal. For example, a CMOS sensor using crystalline silicon can be used as the photoelectric conversion element 112.

The plurality of image pickup portions 110 are fixed to the fixing substrate 160 via a fixing member 170. In a CMOS sensor used as the photoelectric conversion element 112, the size of a crystalline silicon wafer is limited, and thus a desired large-sized image pickup substrate may not be manufactured by using a single image pickup substrate 111. In view of this, as shown in fig. 3, unlike the first embodiment, an image pickup section 110 including an image pickup substrate 111 is formed by being arranged in a 2 × 4 matrix. However, the number of the arranged image pickup substrates 111 is not limited to the 2 × 4 matrix array.

Examples of materials that may be used for the fixed substrate 160 include glass, amorphous carbon, CFRP, and aluminum.

For the fixing member 170, a sheet-like bonding material obtained by, for example, arranging bonding layers above and below a foam body having a void may be used. Such a bonding material has large elasticity due to voids in the foam, and is therefore effective for absorbing height variations of the plurality of imaging substrates 111 and planarizing the imaging surface. It is also possible to use a sheet-like or liquid bonding material containing, for example, silicone resin, acrylic resin, epoxy resin, urethane resin, or hot-melt resin.

In addition, as shown in fig. 3 and 4, the second scintillator panel 120 includes a second scintillator layer 121 and a plurality of FOPs 122. Each FOP 122 is configured by bundling a plurality of optical fibers, and thus a desired large-sized FOP may not be manufactured by using a plurality of FOPs. In view of this, as shown in fig. 3, a manner of arranging the FOPs 122 in a 3 × 3 matrix is adopted. However, the number of array-like FOPs 122 is not limited to the 3 × 3 matrix array.

The radiation imaging apparatus 100-2 further includes two scintillator layers: the first scintillator layer 131 and the second scintillator layer 121 with the FOP 122 interposed therebetween, and therefore the same effects as those of the first embodiment described above can be produced. That is, with the radiation imaging apparatus 100-2, a radiation imaging apparatus with high sensitivity can be realized, and also the proportion of diffused light included in light entering the photoelectric conversion element 112 can be reduced, so that a decrease in the sharpness of a radiographic image can be suppressed.

Third embodiment

Next, a third embodiment is described. In the following description of the third embodiment, the description of the same matters as those of the above-described first and second embodiments is omitted, and matters different from those of the above-described first and second embodiments are described.

Fig. 5 is a perspective view for illustrating an example of a schematic configuration of a radiation imaging apparatus 100 according to a third embodiment of the present invention. In fig. 5, an xyz coordinate system corresponding to the xyz coordinate system shown in fig. 1 and 2 is shown. Further, in fig. 5, the same components as those of the configuration shown in fig. 1 and 2 are denoted by the same reference numerals, and detailed description thereof is omitted.

Fig. 6 is a cross-sectional view for illustrating an example of a detailed configuration of the radiation imaging apparatus 100 according to the third embodiment of the present invention. In fig. 6, an xyz coordinate system corresponding to the xyz coordinate system shown in fig. 5 is shown, and more specifically, fig. 6 is an illustration of a detailed configuration of the radiation imaging apparatus 100 in a plane defined by the x direction and the z direction shown in fig. 5. In fig. 6, the same components as those of the configuration shown in fig. 5 are denoted by the same reference numerals.

In the following description of the third embodiment, the radiation imaging apparatus 100 shown in fig. 5 and 6 is described as "radiation imaging apparatus 100-3". Differences from the above-described first and second embodiments are described below.

As shown in fig. 5 and 6, the radiation imaging apparatus 100-3 includes a first scintillator plate 130, a second scintillator plate 120, and an imaging portion 110. In addition, as shown in fig. 6, the radiation imaging apparatus 100-3 further includes a first bonding member 141, a third bonding member 143, and a second moisture-proof resin 152. In fig. 5, the components are shown to be spaced apart from each other for convenience of description, but as shown in fig. 6, the components may be arranged by being stacked via the first and third engaging members 141 and 143.

The second scintillator layer 121 is made of CsI: Tl, for example. In this case, the second scintillator layer 121 is formed on the image pickup substrate 111 by a vapor deposition method. In addition, as described above, CsI: Tl has deliquescence, and therefore it is desirable that the second scintillator layer 121 made of CsI: Tl be covered with the image pickup substrate 111 and a moisture-proof protective film (not shown). The moisture-proof protective film formed on the second scintillator layer 121 or the second scintillator layer 121 may be connected to the FOP 122 via the third bonding member 143.

As shown in fig. 6, the first scintillator plate 130 includes a first scintillator layer 131 and a support substrate 133. Fig. 6 shows an example in which a member corresponding to the reflective layer 132 shown in fig. 2 is not provided. The first scintillator plate 130 is a non-columnar scintillator plate in which a resin and granular Gd forming the first scintillator layer 131 are formed on a support substrate 133 by a coating method₂O₂S (GOS) phosphor. In the present embodiment, it is desirable that the support substrate 133 be made of a material having a function of reflecting light. For the support substrate 133, not only a metal material but also, for example, a material containing titanium oxide (TiO) can be used₂) A pelletized PET resin sheet. In the case where the resin is disposed around the GOS particles, the GOS forming the first scintillator layer 131 also exhibits less deterioration due to humidity, and thus the first moisture-proof resin 151 is not disposed in fig. 6. The non-columnar scintillator plate can be manufactured by a coating method, and thus the production cost can be reduced in some cases. The GOS forming the first scintillator layer 131 is granular, and thus light scattering and diffusion in the scintillator layer is larger than that in the columnar CsI: Tl. However, the thickness of the first scintillator layer 131 made of GOS can be reduced by arranging the second scintillator layer 121 made of CsI: Tl on the side of the imaging substrate 111 and increasing the thickness of the second scintillator layer 121. That is, the thickness of the first scintillator layer 131 is smaller (thinner) than the thickness of the second scintillator layer 121. The diffused light of the first scintillator layer 131 can also be corrected to diffused light close to the linearly propagating light by disposing the FOP 122 between the first scintillator layer 131 and the second scintillator layer 121.

The radiation imaging apparatus 100-3 may further include two scintillator layers: the first scintillator layer 131 and the second scintillator layer 121 with the FOP 122 interposed therebetween, and therefore the same effects as those of the first embodiment described above can be produced. That is, with the radiation imaging apparatus 100-3 according to the present embodiment, a radiation imaging apparatus having high sensitivity can be realized, and also the proportion of diffused light included in light entering the photoelectric conversion element 112 can be reduced, so that a decrease in the sharpness of a radiation image can be suppressed.

Fourth embodiment

Next, a fourth embodiment is described. In the following description of the fourth embodiment, the description of the same matters as those of the above-described first to third embodiments is omitted, and matters different from those of the above-described first to third embodiments are described.

Fig. 7 is a conceptual diagram of an X-ray imaging system (radiation imaging system) according to a fourth embodiment, which uses the radiation imaging apparatus 100 according to any one of the first to third embodiments.

The X-rays 211 as the radiation R generated by the X-ray tube 210 (radiation generating unit) are transmitted through the chest 221 of the person to be examined 220, such as the examination subject H, to enter the radiation imaging apparatus 100 according to any one of the first to third embodiments. The X-ray 211 entering the radiation imaging apparatus 100 includes information on the inside of the body of the person to be examined 220.

In the radiation imaging apparatus 100, the first scintillator layer 131 and the second scintillator layer 121 emit light in response to the X-rays 211 entering the radiation imaging apparatus 100. The light generated in these scintillator layers is photoelectrically converted into an electric signal by the image pickup portion 110, thereby obtaining electric information about the inside of the body of the person to be examined 220. The electrical information is converted into a digital signal and image-processed by an image processor 230 serving as a signal processing unit, and thus can be observed on a display 240 serving as a display unit of the control room.

Further, the electrical information obtained by the radiation imaging apparatus 100 and processed by the image processor 230 may be forwarded to a remote site through a transmission unit 250 (e.g., a telephone line), and may be transmitted to, for example, a doctor's room located at another place. In a doctor room located at another place, the electrical information received via the transmission unit 250 may be displayed on the display 241 serving as a display unit, or may be stored in a recording unit such as an optical disk, so that a doctor at a remote site may also make a diagnosis. The electrical information can also be recorded on a film 261 serving as a recording medium by the film processor 260 serving as a recording unit.

All of the above embodiments describe only specific examples for implementing the invention. Therefore, the technical scope of the present invention should not be limited by the above-described embodiments. Specifically, the present invention may be embodied in various forms without departing from the technical idea or main features of the present invention.

It is to be understood that the invention is not limited to the disclosed exemplary embodiments. The scope of the following claims is to be accorded the broadest interpretation so as to encompass all such modifications and equivalent structures and functions.