Scintillator panel, x-ray photographing device, and method for manufacturing scintillator panel

Abstract

To provide a scintillator panel with good image characteristics, a digital X-ray inspection device using the scintillator panel, and a method for manufacturing the scintillator panel.SOLUTION: A scintillator panel is provided with a support substrate that transmits radiation, and a scintillator layer that is in contact with the support substrate and converts an X-ray incident from the outside into light. The scintillator layer is a phosphor containing Tl as an activator in CsI that is a halide. When an X-ray incident side of an area in a film thickness direction of the scintillator layer is defined as an A area and a non-incident side as a B area, and a center area forming the scintillator layer is defined as a center part and an outer peripheral area as a peripheral part, the A area and the B area has the following relationships. The A area does not contain the activator. The film thickness distribution in an in-plane direction of the A area is that the center part>the peripheral part. The film thickness distribution in the in-plane direction of the B area is that the center part<the peripheral part.SELECTED DRAWING: Figure 7

Description

[Technical field]

[0001] FIELD OF THE PRESENT EMBODIMENTS The present invention relates to a scintillator panel, an X-ray imaging device using the scintillator panel, and a method for manufacturing a scintillator panel. [Background technology]

[0002] 2. Description of the Related Art Generally, some X-ray imaging devices use a scintillator panel in which a scintillator layer is formed on a support substrate that transmits radiation. Such scintillator panels first convert X-rays into visible light using a scintillator layer, then convert the visible light into signal charges and guide them to a charge storage capacitor. However, optical diffusion and scattering on the way to reaching this charge storage capacitor can sometimes cause degradation of image characteristics such as sensitivity, resolution, and contrast. [Prior art documents] [Patent documents]

[0003] [Patent Document 1] JP 2015-38459 A Summary of the Invention [Problem to be solved by the invention]

[0004] Therefore, there has been a demand for improvements in the image characteristics of scintillator panels, such as sensitivity, resolution, and contrast.

[0005] An object of the present embodiment is to provide a scintillator panel with good image characteristics, an X-ray imaging device using the scintillator panel, and a method for manufacturing the scintillator panel. [Means for solving the problem]

[0006] One embodiment is a scintillator panel comprising a support substrate that transmits radiation, and a scintillator layer that is in contact with the support substrate and converts X-rays incident from the outside into light, the scintillator layer being a phosphor containing CsI, which is a halide, and Tl as an activator, and in which, when the regions in the thickness direction of the scintillator layer are defined as region A on the X-ray incident side and region B on the non-incident side, the central region forming the scintillator layer being the center portion, and the peripheral region being the peripheral portion, region A and region B have the following relationship:

[0007] The A region does not contain the activator, The thickness distribution in the in-plane direction of the region A is such that the central portion>the peripheral portion The thickness distribution in the in-plane direction of the region B is such that the central portion is smaller than the peripheral portion. [Brief description of the drawings]

[0008] [Figure 1] FIG. 1 is a schematic cross-sectional view showing a first structural example of a scintillator panel according to an embodiment. [Diagram 2] FIG. 2 is a schematic cross-sectional view showing a second structural example of a scintillator panel according to an embodiment. [Diagram 3] FIG. 3 is a schematic cross-sectional view showing a third structural example of a scintillator panel according to an embodiment. [Figure 4] FIG. 4 is a schematic cross-sectional view showing a fourth structural example of a scintillator panel according to an embodiment. [Diagram 5] FIG. 5 is a schematic diagram of an X-ray imaging device using a scintillator panel according to an embodiment. [Figure 6] FIG. 6 shows a comparison of X-ray images with and without an X-ray grid, where (a) is an X-ray image without an X-ray grid and (b) is an X-ray image with an X-ray grid. [Figure 7]5(a) is a schematic diagram showing the relationship between regions A and B for a scintillator panel for the structural example shown in Figures 3 and 4, and (b) is a schematic diagram showing a conventional structure that does not have regions A and B for a scintillator panel for the structural example shown in Figures 3 and 4. [Figure 8] FIG. 8 shows X-ray images taken by a general X-ray imaging device, where (a) shows an X-ray image without flat-field correction, and (b) shows an X-ray image with flat-field correction. [Figure 9] Figure 9(a) is a schematic diagram showing the arrangement of an X-ray source and a detector (X-ray planar image detector) in a typical X-ray imaging device, and Figure 9(b) is a plan view of the X-ray planar image detector shown in Figure 9(a). [Figure 10] FIG. 10 is a schematic diagram illustrating a method for forming a scintillator layer according to an embodiment. [Figure 11] FIG. 11 is a diagram for explaining the contrast used in the specification, where (a) is a schematic diagram of a contrast measuring device, (b) is a diagram of an X-ray image detected by an X-ray imaging device, and (c) is a graph showing the contrast ratio in the X-ray image. [Figure 12] FIG. 12 is a diagram showing an X-ray image in an X-ray imaging device using a scintillator panel, where (a) is the scintillator panel according to this embodiment, and (b) is a scintillator panel according to a comparative example. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0009] Hereinafter, the present embodiment will be described with reference to the drawings. Note that the disclosure is merely an example, and those who are skilled in the art can easily conceive of appropriate modifications while maintaining the gist of the invention are naturally included in the scope of the present invention. In addition, the drawings may be schematic in terms of width, thickness, shape, etc. of each part compared to the actual embodiment in order to make the explanation clearer, but they are merely examples and do not limit the interpretation of the present invention. In addition, in this specification and each figure, components that perform the same or similar functions as those described above with respect to the previous figures are given the same reference numerals, and duplicate detailed explanations may be omitted as appropriate.

[0010] The scintillator panel and X-ray imaging device according to the present embodiment can be used in general medical treatment, but is not limited to general medical treatment and can also be used for industrial purposes such as non-destructive testing.

[0011] An embodiment will be described below with reference to Figures 1 to 12, but first, a structural example of a scintillator panel 1 will be described with reference to Figures 1 to 4. Figures 1 to 4 are vertical cross-sectional views of the scintillator panel 1. 1, the scintillator panel 1 has a structure in which a scintillator layer 5 that is in contact with the support substrate 3 and converts incident X-rays into visible light Y is formed on a support substrate 3 that transmits X-rays (radiation). Furthermore, the scintillator panel 1 has a protective layer 7 that blocks the scintillator layer 5 from the outside air, etc., and a reflective layer 9 formed on the support substrate 2 to increase the utilization efficiency of the visible light Y converted by the scintillator layer 5.

[0012] The support substrate 3 is preferably made of a material mainly composed of elements lighter than transition metal elements, such as aluminum (Al), carbon (C), carbon fiber reinforced plastic (CFRP), magnesium alloy, etc.

[0013] The reflective layer 9 is made of a metal material with high reflectivity, such as Al (aluminum), Ni (nickel), Cu (copper), Pd (palladium), Ag (silver), etc. The reflective layer 9 reflects the light generated in the scintillator layer 5 in the direction opposite to the supporting substrate 3, thereby improving the light utilization efficiency.

[0014] The scintillator layer 5 is made of a phosphor containing Tl (thallium) as an activator in a halogen compound of CsI (cesium iodide) in order to improve the output signal intensity with respect to incident X-rays. The scintillator layer 5 will be described later.

[0015] Protective layer 7 protects scintillator layer 5 from moisture in the air. When a halogen compound such as CsI, which is a high-brightness fluorescent material, is used for scintillator layer 5, it is possible to prevent the scintillator layer 5 from deliquescing due to a reaction with moisture in the air, which would otherwise cause deterioration of various characteristics and reliability. Furthermore, when the halide CsI is used in the scintillator layer 5, CsI alone cannot convert incident X-rays into visible light Y, and therefore an activator must be added to activate the optical excitation of incident X-rays, as in the case of general phosphors. The peak wavelength of light sensitivity of light receiving elements (TFT or CCD) used in general indirect type X-ray planar image detectors and DR devices (X-ray imaging devices) exists near the visible light region (400nm to 700nm), so when CsI is used for the scintillator layer 5, Tl (thallium), which emits light excited by incident X-rays with a wavelength of around 550nm, is used as an activator.

[0016] Fig. 2 shows a second structural example of the scintillator panel 1. The scintillator panel 1 of this second structural example differs from the first structural example shown in Fig. 1 in that it does not have a reflective layer 9. The other configurations are the same as those of the first structural example shown in Fig. 1.

[0017] Fig. 3 shows a third structural example of the scintillator panel 1. The scintillator panel 1 of this third structural example differs from the first structural example shown in Fig. 1 in that the scintillator layer 5 has a rectangular columnar crystal structure 11. The other configurations are the same as those of the first structural example shown in Fig. 1. The rectangular columnar crystal structure 11 can be formed by using a vacuum deposition method.

[0018] Fig. 4 shows a fourth structural example of the scintillator panel 1. The scintillator panel 1 of this fourth structural example differs from the third structural example shown in Fig. 3 in that it does not have a reflective layer 9. The other configurations are the same as those of the third structural example shown in Fig. 3.

[0019] Now, with reference to FIG. 5, a CCD-DR type X-ray imaging 13 using the scintillator panel 1 will be described. The X-ray imaging device 13 has a housing 15, and the scintillator panel 1 is installed at one end of the housing 15. A mirror reflector 17 and an optical lens 19 are installed inside the housing 15, and a light receiving element 21 such as a CCD is installed at the other end of the housing 15. X-rays (X in the figure) emitted from an X-ray tube (X-ray source) 23 enter the scintillator panel 1 via a grid 25, and then visible light (X-ray image) Y converted by the scintillator layer 5 is emitted from the surface of the scintillator layer 5. The X-ray image Y is projected on the surface of the scintillator layer 5, and the X-ray image is reflected by the reflector 17 and collected by the optical lens 19 to be irradiated to the light receiving element 21, which converts the X-ray image Y into an electrical signal and outputs it.

[0020] Next, the scintillator layer 5 according to this embodiment will be described in further detail with reference to FIG. 7(a). Scintillator layer 5 is a phosphor containing Tl as an activator in the halide CsI, and when the regions in the thickness direction of scintillator layer 5 are defined as region A on the X-ray incident side and region B on the non-incident side, with the central region forming scintillator layer 5 being the central portion and the outer circumferential region being the peripheral portion, region A and region B have the following relationships (1) to (4). (1) Region A does not contain the activator Tl. (2) The film thickness (arrow T) distribution in the in-plane direction (arrow M direction) of region A is center > peripheral (3) The thickness distribution in the in-plane direction in region B is center < peripheral (4) The scintillator layer has a rectangular columnar crystal structure. In this embodiment, a scintillator panel 1 having a rectangular columnar crystal structure 11 shown in FIG. 3 and FIG. 4 is used. In the scintillator panel 1 shown in FIG. 3, the base material of the scintillator layer 5 is CsI, the activator is Tl, the support substrate 3 is CFRP, and the reflective layer 9 is TiO 2、 The protective layer 7 was made of polyparaxylylene.

[0021] Next, the formation of the scintillator layer 5 will be described with reference to FIG. 10, a substrate 29 on which a scintillator layer 5 is formed is mounted facing downward in a vacuum chamber 27. A first evaporation source 31 and a second evaporation source 33 are provided in a lower portion of the vacuum chamber 27. The first evaporation source 31 is an evaporation source of CsI, and the second evaporation source 33 is an evaporation source of Tl.

[0022] In this way, in the vacuum deposition method using two evaporation sources 31, 33, the scintillator layer 5 is formed by the deposition of evaporated particles CsI and Tl from each evaporation source 31, 33. Therefore, during the process of forming the scintillator layer 5, by blocking or stopping the supply of evaporated particles Tl from the second evaporation source 33, it is possible to form a region that does not contain activator in any region in the film thickness direction of the scintillator layer 5. For this reason, in the present embodiment, in a vacuum deposition method using two evaporation sources 31, 33 as shown in FIG. 10, the supply of evaporated particles Tl from the second evaporation source 33 is blocked or stopped at the beginning of the formation process of the scintillator layer 5, and the supply of evaporated particles CsI from the first evaporation source 31 is controlled, thereby forming a region A that does not contain activator Tl, a region B that contains a predetermined concentration of activator Tl, a film thickness distribution in the in-plane direction of region A (central part > peripheral part), and a film thickness distribution in the in-plane direction of region B (central part < peripheral part).

[0023] Table 1 shows the film thickness (μm) of each of the A and B regions, the activator concentration (mass%) of this embodiment, conventional examples 1 and 2, and the comparative example, and the respective conditions. Table 2 shows the results. [Table 1] In Table 1, the X-ray grid is either provided with the parallel grid 25 in FIG. 5 or not provided with the parallel grid 25 (X-ray grid: none). Conventional examples 1 and 2 are both cases where region A and region B are not provided, and the comparative example shows a case where region A and region B are provided but the conditions are different from those of the embodiment.

[0024] FIG. 6 shows conventional examples 1 and 2 with and without the parallel grid 25, with FIG. 6(a) showing the case without the parallel grid 25 (conventional example 1) and FIG. 6(b) showing the case with the parallel grid 25 (conventional example 2). As shown in Figures 6(a) and 6(b), when an X-ray grid 25 is used (conventional example 2), it is generally possible to improve the contrast and resolution of an X-ray image. However, it becomes difficult to accommodate the miniaturization of the pixel electrode size and pixel electrode pitch of the imaging element (active matrix substrate, CCD, CMOS, etc.) in order to improve image characteristics, and it becomes impossible to prevent image quality degradation due to the occurrence of moire, as shown in Figure 6(b). Furthermore, when the X-ray grid 25 is used in a general X-ray planar image detector or DR device, there are many restrictions (SID, etc.) for obtaining an appropriate X-ray image, and it is structurally difficult to eliminate the effects of scattered X-rays caused by structures existing between the X-ray grid 25 and the scintillator layer 5.

[0025] Next, with reference to Table 2, the image characteristics of this embodiment, conventional examples 1 and 2, and a comparative example will be described. Table 2 shows the image characteristics such as sensitivity ratio, MTF (resolution) ratio, contrast ratio, and the presence or absence of moire. In Table 2, the sensitivity ratio, MTF (resolution) ratio, and contrast ratio are each set with the image characteristics of Conventional Example 1 as the reference value (1.00). The X-ray irradiation conditions for measuring the sensitivity ratio, MTF (resolution) ratio, and contrast ratio were 70KV-0.087mGy. The MTF (resolution) ratio @2.0LP / mm is the spatial frequency, and refers to the case where two pairs of black and white stripes are placed within a width of 1 mm.

[0026] The intensity distribution of scattered X-rays will be described with reference to Fig. 9. Fig. 9(a) is a configuration diagram of a general X-ray imaging device, and Fig. 9(b) is a plan view of a detector (X-ray planar image detector) 24. Reference numeral 23 denotes an X-ray generation source (X-ray tube). In a typical X-ray planar image detector or DR device using a scintillator panel 1, the intensity distribution of scattered X-rays generated when incident X-rays pass through a subject is such that, if the central region C of the detector 24 is the center and the peripheral region P is the periphery, the central region is greater than the periphery. Therefore, if an absorption layer for scattered X-rays corresponding to the intensity of the scattered X-rays in the center is uniformly formed in the formation region of the scintillator layer 5 of the detector 24, concentric sensitivity spots (shading) will be generated due to the intensity distribution of scattered X-rays as shown in Figure 8(a).

[0027] Furthermore, in a typical X-ray planar image detector or DR device using a scintillator panel 1, the X-ray source (X-ray tube) 23 is usually installed on the central axis of the detector 24, and therefore, as shown in FIG. 9, there is a difference in the distance M from the X-ray source 23 and the incident angle θ of the incident X-ray (symbol X) between the central portion C and the peripheral portion P of the detector 24. Here, as shown in Figure 9, if the distances from the X-ray source at the center (C) and peripheral (P) of detector 24 are L and L', the angle of the incident X-ray at the peripheral (P) is θ, the distance from the center (C) to the peripheral (P) is M, and the intensity (energy) of the incident X-ray at the center (C) and peripheral (P) is α and β, respectively, the intensity of the incident X-ray is given by the following formula 1.

[0028]

number

[0029] As shown in formula 1, the intensity of the incident X-rays attenuates in proportion to the square of the distance from the X-ray generation source 23. For this reason, in a typical X-ray planar image detector or DR device using the scintillator panel 1, if the film thickness distribution in the in-plane direction of the scintillator layer formation region that contributes to light emission is uniform, concentric sensitivity spots (shading) as shown in Fig. 8(a) will appear in the X-ray image due to the difference between the distance from the X-ray generation source 23 and the incidence angle of the incident X-rays.

[0030] Therefore, in general X-ray planar image detectors and DR devices 13 using the scintillator panel 1, Flat Field correction is usually applied to the X-ray image to correct the sensitivity spots as shown in Figure 8 (b). However, when Flat Field correction is applied to an X-ray image, the noise components contained in the image are also amplified due to the correction, so the larger the sensitivity spots before correction, the lower the S / N of the image after correction and the smaller the dynamic range of the image. Therefore, the better the uniformity of the sensitivity distribution {BU (Brightness Uniformity)} of the X-ray image, the higher the quality of the X-ray image (high S / N, wide dynamic range) that can be obtained. Therefore, by imparting the above characteristics (1) to (4) to the scintillator layer 5 as shown in Figure 7(a), it is possible to reduce the concentric sensitivity spots associated with the intensity distribution of scattered X-rays as shown in Figure 8(a), making it possible to obtain X-ray images of higher image quality.

[0031] As shown in Figures 3 and 4, the scintillator panel 1 of this embodiment is a general X-ray planar image detector or DR device in which the above-mentioned features (1) to (4) are imparted to the scintillator layer 5. Therefore, region A is formed independently in each columnar crystal structure 11 of the scintillator layer 5 which has the rectangular columnar crystal structure 11, and therefore the following effects A, B, and C can be obtained compared to the case where an X-ray grid 25 (see Figure 5) is used. A: Moire does not occur regardless of the pixel electrode size and pixel electrode pitch of the light receiving element (TFT, CCD, CMOS, etc.), so there is no degradation in the image quality of the X-ray image. B: There are fewer constraints (SID, etc.) in obtaining appropriate X-ray images. C: Since there are no structures between the layer that absorbs scattered X-rays and the scintillator layer 5 that contributes to light emission, the image characteristics (resolution and contrast) are efficiently improved.

[0032] Furthermore, when the radiation quality of the incident X-rays is soft (when the tube voltage of the X-ray generation source (X-ray tube) is low), the absorption of the incident X-rays by the subject becomes large, and the effect of scattered X-rays becomes relatively significant, so that minute differences in the X-ray absorption rate of structures present on the X-ray incident side of the scintillator layer 5 become easily observable in the X-ray image. However, in a general X-ray planar image detector or DR device using the scintillator panel 1 according to the embodiment, when the above features (1) to (4) are imparted to the scintillator layer 5, unlike the case where the X-ray grid 25 (see Figure 5) is used, region A is formed on the non-incident side of the scintillator layer 5 relative to the structures present on the X-ray incident side, so that even when the incident X-rays have a soft radiation quality, it is possible to suppress minute X-ray image spots caused by the influence of scattered X-rays. This makes it possible to further improve the performance of the scintillator panel 1 and to improve the image characteristics (sensitivity, resolution, contrast, etc.) in indirect type X-ray planar image detectors and DR devices using the scintillator panel 1.

[0033] Furthermore, in the case of the scintillator panel 1, when the above characteristics (1) to (4) are imparted to the scintillator layer 5, region A does not contain the activator Tl and therefore becomes an X-ray absorption layer that does not contribute to light emission. Therefore, if the concentration range of the activator Tl in the film thickness direction at least in region B is 0.1 mass% to 2.0 mass%, good image characteristics (sensitivity, resolution, etc.) as shown in the patent document previously filed by the inventors (Patent Application No. 2013-252420) can be obtained. In particular, if the concentration range of the activator Tl in the film thickness direction in region B is 1.6 mass% ±0.4 mass%, it is possible to improve the afterimage characteristics without degrading the image characteristics (sensitivity, resolution, etc.).

[0034] Next, the definitions of the contrast ratios shown in Table 2 will be described with reference to FIG. As shown in FIG. 11(a), incident X-rays (symbol X) generated from an X-ray tube (X-ray source) 23 are irradiated toward a subject 39 embedded in an acrylic plate 37, and an X-ray image is detected by a detector (X-ray planar image detector) 24. Then, as shown in FIG. 11(b), the brightness between the subject 39 and its surroundings in the X-ray image was measured (brightness measurement line), and as shown in FIG. 11(c), a brightness level A from 0 level to the area with the subject 39 and a brightness level B from 0 level to the area without the subject 39 were obtained, which was determined as the contrast ratio B / A. The thickness of the acrylic plate 37 when the contrast ratio shown in FIG. 11 was measured was 60 mm.

[0035] The sensitivity ratios shown in Table 2 will be described with reference to FIG. FIG. 12(a) shows this embodiment, and (b) shows a comparative example. C is the sensitivity of the central area, and P is the sensitivity of the peripheral area (the average of the four points of the peripheral area P shown in the figure). The concentric circles shown by dashed lines in Figures 12(a) and (b) are circles that are 9% of the diagonal length of the square X-ray image. As shown in FIG. 12, the sensitivity BU % is P / C×100, which is 80% in the present embodiment shown in FIG. 12(a) and 62% in the comparative example shown in FIG. 12(b).

[0036] No flat field correction was performed on the X-ray image shown in Fig. 12. The conditions for taking the X-ray image and the image processing conditions are as follows. Incident X-ray: 70KV-0.087mGy, Window processing: Yes (image histogram average value ±50%) As is clear from Table 2, the contrast ratio of Conventional Example 2 is 2.5 compared to Conventional Example 1, and 1.8 in this embodiment, but moire occurs in Conventional Example 2. In contrast, this embodiment is superior to Conventional Examples 1 and 2 and the comparative example in that it obtains predetermined values ​​for the sensitivity ratio, MTF ratio (resolution), and contrast ratio and does not have moire.

[0037] As shown in Fig. 12, according to this embodiment, it is possible to reduce concentric sensitivity spots accompanying the intensity distribution of scattered X-rays as shown in Fig. 8(a), and therefore it is possible to obtain an X-ray image with higher image quality. This makes it possible to obtain a further improvement in the performance of the scintillator panel 1 and an improvement in the image characteristics (sensitivity, resolution, contrast, etc.) in an indirect type X-ray planar image detector or DR device using the scintillator panel.

[0038] It should be noted that the present invention is not limited to the above-described embodiment itself, and that the components can be modified and embodied at the implementation stage without departing from the gist of the invention. In addition, various inventions can be formed by appropriately combining multiple components disclosed in the above-described embodiment. For example, some components may be deleted from all the components shown in the embodiment. Furthermore, components from different embodiments may be appropriately combined. [Explanation of symbols]

[0039] 1...scintillator panel, 3...support substrate, 5...scintillator layer, A: Region of the scintillator layer on the X-ray incidence side; B: Region of the scintillator layer on the side where X-rays are not incident.

Claims

1. A support substrate that is transparent to radiation; a scintillator layer in contact with the support substrate for converting X-rays incident from the outside into light; The scintillator layer is a phosphor containing Tl as an activator in CsI, which is a halide, and when the regions in the thickness direction of the scintillator layer are defined as region A on the X-ray incident side and region B on the non-incident side, the central region forming the scintillator layer is defined as the center portion, and the outer peripheral region is defined as the peripheral portion, the region A and the region B have the following relationship: The A region does not contain the activator, The thickness distribution in the in-plane direction of the region A is such that the central portion is greater than the peripheral portion. The thickness distribution in the in-plane direction of the region B is such that the central portion is smaller than the peripheral portion.

2. 2. The scintillator panel in accordance with claim 1, wherein the scintillator layer has a rectangular columnar crystal structure, and the A region occupies at least 5% to 50% of the film thickness of the scintillator layer.

3. 2. The scintillator panel according to claim 1, wherein the central portion of the scintillator layer occupies 50% or more of a concentric or rectangular region formed on the scintillator layer with the central portion as a reference.

4. 2. The scintillator panel according to claim 1, wherein the concentration of the activator in at least the region B in the film thickness direction is in the range of 0.1 mass % to 2.0 mass %.

5. 2. The scintillator panel according to claim 1, wherein the supporting substrate is made of a material containing an element having a lighter mass than a transition metal element.

6. 6. An X-ray imaging device comprising: the scintillator panel according to claim 1; and a photoelectric conversion section that converts fluorescence emitted by the scintillator panel into an electric charge signal.

7. A method for producing a scintillator panel according to any one of claims 1 to 5, comprising the steps of: The method for producing a scintillator panel is characterized in that the step of forming the scintillator layer on one surface of the support substrate is performed by a vapor phase growth method using CsI and Tl as material sources.

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