Scintillator panel, and method for manufacturing a scintillator panel

Abstract

The objective is to provide a scintillator panel that can improve image characteristics and X-ray resistance, and a method for manufacturing the scintillator panel. [Solution] The scintillator panel according to the embodiment comprises a support substrate that transmits X-rays and a scintillator layer that is in contact with the support substrate and converts X-rays incident from the outside into light. The scintillator panel is provided with regions of different emission wavelengths in the direction of the thickness of the scintillator layer, and the peak wavelengths of the emission centers of each region of different emission wavelengths are α and β, respectively. The upper and lower limits of the full width at half maximum of the peak wavelength of the light receiving sensitivity of the photodetector in the detector into which the light from the scintillator layer is incident are γ1 and γ2, respectively, and the following equation is satisfied. α

Description

Technical Field

[0001] Embodiments of the present invention relate to a scintillator panel and a method for manufacturing the scintillator panel.

Background Art

[0002] There are X-ray flat panel imagers using active matrix or solid state imaging devices (such as CCD and CMOS). An X-ray imager is a solid state detector that irradiates a flat detector with X-rays to output an X-ray image or a real-time X-ray image as a digital signal. Therefore, many universities and manufacturers are engaged in research and development and commercialization because of its excellent image quality performance and stability. In addition, the main applications of X-ray imagers using active matrix are for chest and general radiography to collect still images with a relatively large dose, and for the cardiovascular and digestive fields where it is necessary to realize a real-time video of 30 Frame / sec or more under fluoroscopic dose. Development and commercialization are in progress. However, for such video applications, improvement of S / N and real-time processing technology for minute signals are important development items. Next, the main applications of X-ray imagers using solid state imaging devices (such as CCD and CMOS) include industrial non-destructive inspection to collect still images with a large dose, and dental use to insert into the oral cavity to collect still images. Also, including the correspondence to video applications, improvement of S / N, real-time processing of minute signals, miniaturization of the detector, improvement of reliability, etc. are important development items. Furthermore, X-ray imagers are roughly classified into two types: direct type and indirect type. The direct type is a method that directly converts X-rays into charge signals by a photoconductive film such as a-Se and guides them to a capacitor for charge accumulation. Since the photoconductive charges generated by X-rays are directly guided to the capacitor for charge accumulation by a high electric field, image characteristics defined by almost the pixel electrode pitch of the active matrix can be obtained. On the other hand, the indirect method converts X-rays into visible light using a scintillator, converts the visible light into signal charges using an a-Si photodiode, CCD, CMOS, etc., and guides the converted signal charges to a charge storage capacitor. Therefore, the optical diffusion and scattering of visible light from the scintillator to the photodiode, CCD, or CMOS degrades the image characteristics. Furthermore, indirect X-ray image detectors can be broadly classified into two types: one in which a semiconductor photodetector is formed on an active matrix substrate and a scintillator is directly formed on a circuit board; and another in which a scintillator panel, in which a scintillator is formed on an X-ray-transmitting support substrate, is bonded to a circuit board in which a semiconductor photodetector is formed on an active matrix substrate. In this case, since X-rays are incident on the scintillator, it is desirable to equip indirect X-ray image detectors with scintillators that have high X-ray resistance (i.e., less sensitivity reduction due to damage from X-ray irradiation). Therefore, there was a need for the development of technologies that could improve image characteristics and X-ray resistance in scintillator panels equipped with scintillators. [Prior art documents] [Patent Documents]

[0003] [Patent Document 1] Japanese Patent Publication No. 2019-174184 [Overview of the project] [Problems that the invention aims to solve]

[0004] The problem that this invention aims to solve is to provide a scintillator panel that can improve image characteristics and X-ray resistance, and a method for manufacturing a scintillator panel. [Means for solving the problem]

[0005] The scintillator panel according to the embodiment comprises an X-ray-transmitting support substrate and a scintillator layer in contact with the support substrate that converts X-rays incident from the outside into light. The scintillator panel is provided with regions of different emission wavelengths in the thickness direction of the scintillator layer, and the peak wavelengths of the emission centers in each of the regions of different emission wavelengths are denoted as α and β, respectively. The upper and lower limits of the full width at half maximum of the peak wavelength of the light receiving sensitivity of the photodetector in the detector into which the light from the scintillator layer is incident are denoted as γ1 and γ2, respectively, and the following equation is satisfied. α < γ2 ≤ β ≤ γ1 [Effects of the Invention]

[0006] According to embodiments of the present invention, it is possible to provide a scintillator panel that can improve image characteristics and X-ray resistance, and a method for manufacturing the scintillator panel. [Brief explanation of the drawing]

[0007] [Figure 1] This is a schematic perspective view illustrating the X-ray image detector according to this embodiment. [Figure 2] This is a schematic cross-sectional view illustrating an X-ray image detector. [Figure 3] This is a block diagram of an X-ray image detector. [Figure 4] This is a schematic cross-sectional view illustrating an X-ray image detector according to another embodiment. [Figure 5] This is a schematic diagram illustrating an X-ray image detector according to another embodiment. [Figure 6] This is a photograph illustrating an X-ray image obtained using a grid. [Figure 7] This is a photograph illustrating an X-ray image obtained using the scintillator layer according to this embodiment. [Figure 8] This table illustrates the types of objects that can be measured for image characteristics and afterimage characteristics. [Figure 9] This is a schematic diagram illustrating contrast shooting conditions. [Figure 10](a) is a schematic diagram for exemplifying the measurement conditions of contrast. (b) is a schematic diagram for exemplifying the method of calculating the contrast ratio. [Figure 11] It is a table for exemplifying the measurement results of image characteristics. [Figure 12] It is a table for exemplifying the sensitivity attenuation ratio. [Figure 13] (a) and (b) are schematic diagrams for exemplifying the method of measuring afterimage characteristics. [Figure 14] It is a table for exemplifying the measurement conditions of the afterimage ratio. [Figure 15] It is a table for exemplifying the measurement results of the afterimage ratio. [Figure 16] It is a schematic diagram for exemplifying the arrangement of the X-ray source and the X-ray image detector. [Figure 17] It is a photograph for exemplifying the sensitivity spots generated in the X-ray image. [Figure 18] It is a schematic cross-sectional view for exemplifying the scintillator layer. [Figure 19] It is a schematic cross-sectional view for exemplifying the scintillator layer according to another embodiment. [Figure 20] It is a table for exemplifying the measurement objects of image characteristics and afterimage characteristics. [Figure 21] It is a table for exemplifying the measurement results of image characteristics. [Figure 22] It is a table for exemplifying the sensitivity attenuation ratio. [Figure 23] It is a table for exemplifying the measurement results of the afterimage ratio. [Figure 24] It is a table for exemplifying the measurement results of BU(%). [Figure 25] (a) and (b) are schematic diagrams for exemplifying the manufacturing method of the scintillator layer. [Figure 26] (a) to (d) are schematic diagrams for exemplifying the aperture of the mask and the film formation state.

Embodiments for Carrying Out the Invention

[0008] The embodiments will be illustrated below with reference to the drawings. In each drawing, similar components are denoted by the same reference numerals, and detailed descriptions will be omitted as appropriate.

[0009] Furthermore, X-ray image detectors can be used, for example, in general medical care. However, the applications of X-ray image detectors are not limited to general medical care. For example, X-ray image detectors can also be used in non-destructive testing and in dental care where still images are taken. In addition, X-ray image detectors can be made capable of capturing at least one of both still images and moving images.

[0010] Figure 1 is a schematic perspective view illustrating an X-ray image detector 1 according to this embodiment. Figure 2 is a schematic cross-sectional view illustrating the X-ray image detector 1. To avoid complexity, circuit section 4 is omitted in Figure 2. Figure 3 is a block diagram of the X-ray image detector 1.

[0011] As shown in Figures 1 and 2, the X-ray image detector 1 includes, for example, a circuit board 2, a scintillator layer 3, a circuit section 4, a reflective layer 5, and a moisture-proof section 6.

[0012] The circuit board 2 may be provided with a substrate 2a, a photoelectric conversion unit 2b, a control line (or gate line) 2c1, a data line (or signal line) 2c2, wiring pads 2d1 and 2d2, and a protective layer 2f. Note that the number of photoelectric conversion units 2b, control lines 2c1, and data lines 2c2 are not limited to those exemplified in Figure 1.

[0013] The substrate 2a is plate-shaped and is formed from glass, such as alkali-free glass. The planar shape of the substrate 2a is, for example, a rectangle.

[0014] Multiple photoelectric conversion units 2b are provided on one side of the substrate 2a. Each photoelectric conversion unit 2b, for example, has a rectangular shape and is located in an area demarcated by control lines 2c1 and data lines 2c2. Multiple photoelectric conversion units 2b can be arranged in a matrix. Note that one photoelectric conversion unit 2b corresponds, for example, to one pixel in an X-ray image.

[0015] Each of the multiple photoelectric conversion units 2b is provided with, for example, a photoelectric conversion element 2b1 which is a semiconductor photodetector, a thin-film transistor (TFT) 2b2 which is a switching element, and a storage capacitor 2b3. The photoelectric conversion element 2b1 which is a semiconductor photodetector and the storage capacitor 2b3 are pixel electrodes. The photoelectric conversion element 2b1 is in contact with the circuit board 2 and converts light incident from the outside into an electrical signal. The photoelectric conversion element 2b1 is, for example, a photodiode. In other words, the circuit board 2 has multiple pixel electrodes and switching elements arranged in one or two dimensions on a planar substrate 2a. Note that the circuit board 2 illustrated in Figure 1 has multiple pixel electrodes and switching elements arranged in two dimensions on a planar substrate 2a.

[0016] The thin-film transistor 2b2 switches the storage and release of charge to the storage capacitor 2b3. The thin-film transistor 2b2 has, for example, a gate electrode 2b2a, a drain electrode 2b2b, and a source electrode 2b2c. The gate electrode 2b2a of the thin-film transistor 2b2 is electrically connected, for example, to the corresponding control line 2c1. The drain electrode 2b2b of the thin-film transistor 2b2 is electrically connected, for example, to the corresponding data line 2c2. The source electrode 2b2c of the thin-film transistor 2b2 is electrically connected, for example, to the corresponding photoelectric conversion element 2b1 and the storage capacitor 2b3. The anode side of the photoelectric conversion element 2b1 and the storage capacitor 2b3 can also be electrically connected to ground. In addition, the anode side of the photoelectric conversion element 2b1 and the storage capacitor 2b3 can also be electrically connected to a bias line.

[0017] The storage capacitor 2b3 stores the signal charge converted by the photoelectric conversion element 2b1 from the fluorescence from the scintillator layer 3. However, the storage capacitor 2b3 is not always necessary; depending on the capacitance of the photoelectric conversion element 2b1, the photoelectric conversion element 2b1 can also function as the storage capacitor 2b3.

[0018] The photoelectric conversion element 2b1, the thin-film transistor 2b2, and the storage capacitor 2b3 can be electrically connected, for example, through a through-hole 2b4 which is a contact hole.

[0019] As shown in Figure 1, multiple control lines 2c1 can be provided parallel to each other at predetermined intervals. The control lines 2c1 extend, for example, in the row direction. One control line 2c1 is electrically connected to one of a plurality of wiring pads 2d1 provided near the periphery of the substrate 2a. One wiring pad 2d1 is electrically connected to one of a plurality of wirings provided on the flexible printed circuit board 2e1. The other end of the plurality of wirings provided on the flexible printed circuit board 2e1 is electrically connected to a readout circuit 4a provided in the circuit section 4.

[0020] Multiple data lines 2c2 can be provided parallel to each other at predetermined intervals. The data lines 2c2 extend, for example, in a column direction perpendicular to the row direction. One data line 2c2 is electrically connected to one of a plurality of wiring pads 2d2 provided near the periphery of the substrate 2a. One wiring pad 2d2 is electrically connected to one of a plurality of wirings provided on the flexible printed circuit board 2e2. The other end of the plurality of wirings provided on the flexible printed circuit board 2e2 is electrically connected to a signal detection circuit 4b provided in the circuit section 4.

[0021] The control line 2c1 and data line 2c2 are formed using, for example, a low-resistance metal such as aluminum or chromium.

[0022] The protective layer 2f covers, for example, the photoelectric conversion unit 2b, the control line 2c1, and the data line 2c2. The protective layer 2f is formed from an insulating material such as an oxide or nitride.

[0023] The scintillator layer 3 is provided on top of the multiple photoelectric conversion units 2b and converts the incident X-rays 100 into fluorescence (visible light). The scintillator layer 3 is provided on the substrate 2a so as to cover the region (effective pixel region) on which the multiple photoelectric conversion units 2b are provided. In other words, the scintillator layer 3 is in contact with the photoelectric conversion element 2b1, which is a semiconductor photodetector, and converts the X-rays incident from the outside into light.

[0024] As will be described later, the scintillator layer 3 is formed, for example, by vapor phase growth (vacuum deposition). As shown in Figure 2, when the scintillator layer 3 is formed using vapor phase growth, a scintillator layer 3 containing multiple columnar crystals 3a (a scintillator layer 3 consisting of a phosphor having a strip-shaped columnar crystal structure) is formed. In this case, the thickness of the scintillator layer 3 is, for example, about 600 μm. The thickness of the columnar crystals 3a (the dimensions of the columnar crystals 3a in a direction intersecting the direction in which the columnar crystals 3a extend) is, for example, about 8 μm to 12 μm.

[0025] When forming the scintillator layer 3 using vapor phase growth, a mask with openings is used. In this case, as shown in Figure 1, the portion of the scintillator layer 3 formed at the openings of the mask (the portion of the scintillator layer 3 formed above the effective pixel region) has a nearly constant thickness. The portion of the scintillator layer 3 formed outside the openings of the mask (the portion of the scintillator layer 3 formed outside the effective pixel region) has a gradually decreasing thickness as it moves outwards.

[0026] Furthermore, as shown in Figure 2, each of the multiple columnar crystals 3a is provided with a first portion 3a1 (region A) and a second portion 3a2 (region B). In the direction in which the columnar crystal 3a extends, the second portion 3a2 is provided on one end side of the first portion 3a1. The first portion 3a1 is located on the incident side of the X-rays 100, and the second portion 3a2 is located on the opposite side from the incident side of the X-rays 100 (for example, on the photoelectric conversion unit 2b side). In each of the multiple columnar crystals 3a, the first portion 3a1 and the second portion 3a2 are provided integrally.

[0027] As will be described later, the first portion 3a1 contains Na (sodium) as an activator. The second portion 3a2 contains Tl (thallium) as an activator. Details of the scintillator layer 3, which contains multiple columnar crystals 3a having the first portion 3a1 and the second portion 3a2, will be described later.

[0028] As shown in Figure 1, the circuit section 4 is provided, for example, on the side of the circuit board 2 opposite to the side where the scintillator layer 3 is provided. For example, the circuit section 4 is electrically connected to the circuit board 2 via flexible printed circuit boards 2e1 and 2e2. As shown in Figure 3, the circuit section 4 includes, for example, a readout circuit 4a and a signal detection circuit 4b.

[0029] The readout circuit 4a switches, for example, the on and off states of the thin-film transistor 2b2. The readout circuit 4a includes, for example, a plurality of gate drivers 4aa and row selection circuits 4ab.

[0030] The row selection circuit 4ab receives a control signal S1 from, for example, an image processing unit located outside the X-ray image detector 1. The row selection circuit 4ab inputs the control signal S1 to the corresponding gate driver 4aa according to the scanning direction of the X-ray image. The gate driver 4aa inputs the control signal S1 to the corresponding control line 2c1. The control signal S1 input to the control line 2c1 turns on the thin-film transistor 2b2, enabling the reading of signal charge (image data signal S2) from the storage capacitor 2b3.

[0031] The signal detection circuit 4b includes, for example, multiple integrating amplifiers 4ba, multiple selection circuits 4bb, and multiple AD converters 4bc.

[0032] The integrating amplifier 4ba sequentially receives image data signals S2 from the photoelectric conversion unit 2b. The integrating amplifier 4ba integrates the current flowing within a certain time and outputs a voltage corresponding to the integrated value to the selection circuit 4bb. The integrating amplifier 4ba converts image data information corresponding to the fluorescence intensity distribution generated in the scintillator layer 3 into potential information.

[0033] The selection circuit 4bb selects the integral amplifier 4ba to perform the readout, and sequentially reads out the image data signal S2, which has been converted into potential information.

[0034] The AD converter 4bc sequentially converts the read-out image data signal S2 into a digital signal. For example, an image processing unit located outside the X-ray image detector 1 constructs an X-ray image based on the image data signal S2 converted into a digital signal.

[0035] The reflective layer 5 is provided to improve sensitivity characteristics by increasing the utilization efficiency of fluorescence generated in the scintillator layer 3. For example, the reflective layer 5 reflects the fluorescence generated in the scintillator layer 3 that is directed away from the side where the photoelectric conversion unit 2b is located, so that it is directed towards the photoelectric conversion unit 2b.

[0036] As shown in Figure 2, the reflective layer 5 is provided on the scintillator layer 3 on the side opposite to the photoelectric conversion section 2b. For example, the reflective layer 5 can be provided between the scintillator layer 3 and the moisture-proof section 6. For example, the reflective layer 5 can be provided so as to cover the area on the upper surface of the scintillator layer 3 that is opposite to the effective pixel area.

[0037] The reflective layer 5 can be formed by coating a material, for example, a mixture of light-scattering particles containing titanium dioxide (TiO2), a resin, and a solvent, onto the scintillator layer 3 and drying it. Alternatively, the reflective layer 5 may be formed by depositing or attaching a film containing a metal with high light reflectivity, such as a silver alloy or aluminum, onto the scintillator layer 3. The thickness of the reflective layer 5 can be, for example, about 100 μm. However, the reflective layer 5 is not necessarily required and should be provided as appropriate depending on the sensitivity characteristics required for the X-ray image detector 1.

[0038] As shown in Figure 2, the moisture-proof section 6 covers the scintillator layer 3 and the reflective layer 5. The peripheral region of the moisture-proof section 6 can be bonded to the circuit board 2, for example. The moisture-proof section 6 is provided to suppress the deterioration of the properties of the scintillator layer 3 due to moisture contained in the air. For this reason, the moisture-proof section 6 is formed from a material with a low moisture permeability coefficient. The moisture-proof section 6 is formed from, for example, a metal such as aluminum, or a resin film with a metal foil laminated on it.

[0039] There may be a gap between the moisture-proof section 6 and the reflective layer 5, or the moisture-proof section 6 and the reflective layer 5 may be in contact. For example, if the vicinity of the periphery of the moisture-proof section 6 is bonded to the circuit board 2 in an environment where the pressure is lower than atmospheric pressure, the area covered by the moisture-proof section 6 will be subjected to a pressure lower than atmospheric pressure. Therefore, atmospheric pressure can cause the moisture-proof section 6 and the reflective layer 5 to come into contact. The thickness of the moisture-proof section 6 is determined considering factors such as X-ray absorption and rigidity. The thickness of the moisture-proof section 6 is, for example, about 0.1 mm.

[0040] Next, we will further describe a scintillator layer 3 that contains multiple columnar crystals 3a, each having a first portion 3a1 and a second portion 3a2.

[0041] X-rays emitted from an X-ray source such as an X-ray tube pass through the subject and then enter the X-ray image detector 1. Here, when the X-rays 100 pass through the subject, so-called scattered X-rays 100a may be generated.

[0042] When scattered X-rays 100a are generated, as shown in Figure 2, the columnar crystal 3a may be incident on both the X-rays 100 that have passed through the subject and the scattered X-rays 100a that have not passed through the subject. In this case, the photoelectric conversion element 2b1 of the photoelectric conversion unit 2b may be incident on both the fluorescence converted from the X-rays 100 that have passed through the subject and the unintended fluorescence converted from the scattered X-rays 100a that have not passed through the subject. Therefore, when scattered X-rays 100a are generated, the image characteristics of the X-ray image may deteriorate.

[0043] In this case, if a grid is provided on the X-ray incident side of the scintillator layer 3, scattered X-rays 100a can be suppressed from entering the scintillator layer 3. However, using a grid may cause moiré patterns (interference fringes) due to fluorescence interference. When moiré patterns occur, the image quality of the X-ray image may deteriorate significantly, making it difficult to miniaturize the size and pitch of the photoelectric conversion unit. Furthermore, using a grid imposes many constraints on obtaining a suitable X-ray image (for example, constraints on the distance between the X-ray source and the X-ray image detector). Moreover, when a grid is used, it becomes structurally impossible to remove scattered X-rays 100a generated by structures existing between the grid and the scintillator layer 3.

[0044] Therefore, the scintillator layer 3 according to this embodiment contains a plurality of columnar crystals 3a having a first portion 3a1 and a second portion 3a2.

[0045] As shown in Figure 2, in each of the multiple columnar crystals 3a, the first portion 3a1 is located on the incident side of the X-rays 100 than the second portion 3a2. The first portion 3a1 is located in the scintillator layer 3, including the end on the incident side of the X-rays 100. Therefore, scattered X-rays 100a are incident on the second portion 3a2 via the first portion 3a1.

[0046] In this case, if the first portion 3a1 absorbs the scattered X-rays 100a, the first portion 3a1 can be given the function of a grid. Furthermore, since the first portion 3a1 absorbs the scattered X-rays 100a, the aforementioned problems that arise from using a grid can be solved.

[0047] In this case, the scintillator layer 3 is generally formed using the halide CsI (cesium iodide). Furthermore, an activator is added to the CsI to activate fluorescence excitation. In this case, the activator is appropriately selected according to the peak wavelength of the photoelectric conversion element 2b1's light-receiving sensitivity.

[0048] For example, if the photoelectric conversion element 2b1 is a photoderode, the peak wavelength of light detection sensitivity is often around 550 nm. In this case, by adding Tl as an activator to CsI, fluorescence with a peak wavelength of around 550 nm can be generated. Therefore, by using Tl as an activator, the fluorescence intensity can be efficiently detected by the photoelectric conversion element 2b1.

[0049] In this case, if an activator is not added to the halide such as CsI, even if the incident scattered X-rays 100a are absorbed, almost no fluorescence will be generated. Therefore, for example, by providing a first portion 3a1 containing a halide without an activator and a second portion 3a2 containing a halide with an activator, the first portion 3a1 can be used as an absorption layer for scattered X-rays 100a, and the second portion 3a2 can be used as a layer that converts the X-rays 100 transmitted through the subject into fluorescence.

[0050] In other words, the first portion 3a1, to which no activator is added, can be given the function of a grid, thereby absorbing and removing scattered X-rays 100a, and solving the aforementioned problems that arise from using a grid.

[0051] However, in recent years, there has been a demand for further improvements in image characteristics, such as for higher resolution X-ray images.

[0052] Therefore, in the scintillator layer 3 according to this embodiment, an activator (corresponding to an example of the first activator) is also added to the first portion 3a1. However, if the first portion 3a1 contains a halide to which an activator has been added, the incident scattered X-rays 100a are converted into fluorescence, and the converted fluorescence may be incident on the photoelectric conversion element 2b1. Therefore, as mentioned above, by utilizing the fact that there is a range in the peak wavelength of the photosensitive reception of the photoelectric conversion element 2b1, the peak wavelength of the emission center of the fluorescence generated in the first portion 3a1 is made to be outside the range of the peak wavelength of the photosensitive reception of the photoelectric conversion element 2b1.

[0053] According to the findings of the present inventors, when the peak wavelength of the fluorescence emission center of the first portion 3a1 is α (μm), the peak wavelength of the fluorescence emission center of the second portion 3a2 is β (μm), the upper limit of the full width at half maximum of the peak wavelength of the photoelectric conversion element 2b1's light-receiving sensitivity is γ1 (μm), and the lower limit of the full width at half maximum of the peak wavelength of the photoelectric conversion element 2b1's light-receiving sensitivity is γ2 (μm), “α(μm)<γ2(μm)≦β(μm)≦γ1(μm)” The activator to be added to the first part 3a1 should be selected such that the following relationship holds true.

[0054] For example, by adding Na as an activator to the halide CsI, fluorescence with a peak wavelength of around 420 nm can be generated. As mentioned above, the peak wavelength of the photoelectric conversion element 2b1's light-receiving sensitivity is around 550 nm, so even if fluorescence with a peak wavelength of around 420 nm is incident on the photoelectric conversion element 2b1, it is unlikely to be converted into a signal charge.

[0055] Therefore, by selecting an activator to be added to the first part 3a1 such that "α(μm) < γ2(μm)", the absorption amount of scattered X-rays 100a can be increased, and even if the fluorescence generated by the conversion of scattered X-rays 100a is incident on the photoelectric conversion element 2b1, it will not become a noise factor. As described above, the first part 3a comprises a halide and an activator, which converts incident X-rays into fluorescence having a peak wavelength α (μm). The second part 3a2 comprises a halide and an activator, which converts incident X-rays into fluorescence having a peak wavelength β (μm).

[0056] The following effects can be enjoyed by using the scintillator layer 3 according to this embodiment.

[0057] Firstly, when X-rays 100 are incident on the scintillator layer 3, short-wavelength fluorescence with higher energy than the peak wavelength of the fluorescence emission center generated in the second portion 3a2 is generated in the first portion 3a1. Then, the high-energy, short-wavelength fluorescence generated in the first portion 3a1 propagates through the second portion 3a2, thereby discharging the charge trapped within the crystal of the second portion 3a2. Since the charge trapped within the crystal of the second portion 3a2 is a factor in the deterioration of afterimage characteristics, discharging the trapped charge can improve the afterimage characteristics.

[0058] Secondly, since the first portion 3a1 contains a halide to which Na is added as an activator, it has higher X-ray resistance (less sensitivity reduction due to damage from X-ray irradiation) compared to the second portion 3a2 which contains a halide to which Tl is added as an activator (corresponding to an example of the second activator). However, when Tl is added as an activator, there is a problem that the higher the concentration of Tl, the lower the X-ray resistance of the second portion 3a2 becomes. As mentioned above, if the first portion 3a1 is provided, the afterimage characteristics of the scintillator layer 3 are improved, so it is possible to lower the concentration of Tl in the second portion 3a2. Therefore, if the first portion 3a1 is provided, the X-ray resistance of the second portion 3a2 can also be improved by lowering the concentration of Tl, thus improving the X-ray resistance of the entire scintillator layer 3.

[0059] Thirdly, the fluorescence generated in the first portion 3a1 can be at a wavelength shorter than the lower limit (γ2) of the full width at half maximum of the peak wavelength of the photoelectric conversion element 2b1, or at a wavelength longer than the upper limit γ1 (μm). Therefore, even if the fluorescence generated in the first portion 3a1 is incident on the photoelectric conversion element 2b1, it is hardly converted into a signal charge. Thus, the first portion 3a1 can perform a function similar to that of a grid, and even if fluorescence is generated in the first portion 3a1, it does not become a noise factor.

[0060] Fourth, the first portion 3a1, which is a columnar crystal, and the second portion 3a2, which is also a columnar crystal, can be easily formed by vapor phase growth. In this case, if the halides contained in the first portion 3a1 and the second portion 3a2 are the same, the crystal structure (cubic) and refractive index can be made identical, and the following two effects can be enjoyed. (1) Since fluorescence propagation is not inhibited at the interface between the first portion 3a1 and the second portion 3a2, a decrease in the sensitivity of the scintillator 3 caused by the provision of the first portion 3a1 is unlikely to occur. (2) Since the first portion 3a1 and the second portion 3a2 are each individual columnar crystals and are formed continuously, the presence of the first portion 3a1 does not easily cause a decrease in the resolution (MTF) of the scintillator 3.

[0061] Furthermore, according to the findings of the present inventors, the concentration of the activator (e.g., Na) in the first portion 3a1 is 1.0 × 10 -3 mass% or more, 10.0×10 -3 It is preferable that the concentration be less than or equal to mass%. The concentration of the activator (e.g., Tl) in the second portion 3a2 is preferably between 0.1 mass% and 2.0 mass%. By doing so, you can more reliably enjoy the effects mentioned above.

[0062] Furthermore, it is preferable that the ratio of the length of the first portion 3a1 to the length of the columnar crystal 3a in the direction in which the columnar crystal 3a extends is 5% or more and 50% or less. In this way, the aforementioned effects of the first portion 3a1 can be enjoyed, and the amount of fluorescence generated in the second portion 3a2 can be ensured.

[0063] As described above, the scintillator layer 3 according to this embodiment makes it possible to improve the characteristics of the scintillator 3 (for example, achieving both improved afterimage characteristics and improved X-ray resistance). Furthermore, it is possible to further improve the image characteristics of the X-ray image detector 1 (resolution (MTF), contrast ratio, and moiré).

[0064] In the above examples, the scintillator layer 3 is formed directly on the circuit board 2, but a scintillator panel 30 having the scintillator layer 3 and the substrate 31 may be bonded to the circuit board 2.

[0065] Figure 4 is a schematic cross-sectional view illustrating an X-ray image detector 1a according to another embodiment. To avoid complexity, circuit section 4 is omitted in Figure 4. As shown in Figure 4, the X-ray image detector 1a includes, for example, a circuit board 2, a scintillator panel 30, and a junction 33. Although not shown in the figure, the X-ray image detector 1a can also include a circuit section 4, similar to the X-ray image detector 1 described above.

[0066] The scintillator panel 30 comprises a scintillator layer 3, a reflective layer 5, a substrate 31, and a moisture-proof section 32. The substrate 31 is plate-shaped and X-ray transparent. The substrate 31 is formed from a material with high X-ray transmittance. For example, the substrate 31 can be formed from aluminum alloy, carbon fiber reinforced plastic (CFRP), carbon, etc. A scintillator layer 3 is provided on one side of the substrate 31 via a reflective layer 5.

[0067] As mentioned above, the reflective layer 5 is not necessarily required and should be provided as appropriate depending on the sensitivity characteristics required for the X-ray image detector 1a.

[0068] As mentioned above, the first portion 3a1 is located on the incident side of the X-rays 100 than the second portion 3a2. In this case, as shown in Figure 4, the X-rays 100 are incident on the scintillator layer 3 via the substrate 31. Therefore, the first portion 3a1 is located on the substrate 31 side than the second portion 3a2.

[0069] In other words, in the case of the X-ray image detector 1 described above, the second portion 3a2 of the multiple columnar crystals 3a contained in the scintillator layer 3 is located on the side of the multiple photoelectric conversion units 2b. In contrast, in the case of the scintillator panel 30, a scintillator layer 3 containing a plurality of columnar crystals 3a is provided on one side of the substrate 31. The first portion 3a1 is located between the substrate 31 and the second portion 3a2.

[0070] The moisture-proof section 32 covers the scintillator layer 3, the reflective layer 5, and the substrate 31. The moisture-proof section 32 is provided to suppress the deterioration of the properties of the scintillator layer 3 due to water vapor contained in the air. The moisture-proof section 32 is formed from a material that is translucent and has a low moisture permeability coefficient. The moisture-proof section 32 can be formed from, for example, polyparaxylylene, polymonochloroparaxylylene, polyfluoroparaxylylene, polydimethylparaxylylene, polydiethylparaxylylene, etc. The moisture-proof section 32 can be formed using, for example, a thermal CVD method.

[0071] The joint 33 is provided between the circuit board 2 and the scintillator panel 30. The joint 33 is light-transmitting and joins the circuit board 2 and the scintillator panel 30. In this case, the scintillator layer 3 of the scintillator panel 30 is joined to the region of the circuit board 2 where the multiple photoelectric conversion units 2b are provided.

[0072] The joint 33 can be formed, for example, by curing an optical adhesive or optical gel. Alternatively, the joint 33 may be made of, for example, optical double-sided tape (OCA tape (Optical Clear Adhesive Tape)).

[0073] Even when the scintillator layer 3 according to this embodiment is provided on the scintillator panel 30, the effects and benefits of the scintillator layer 3 described above can still be enjoyed.

[0074] Figure 5 is a schematic diagram illustrating an X-ray image detector 1b according to another embodiment. In the aforementioned X-ray image detector 1a, the example given was that the scintillator panel 30 is bonded to the circuit board 2, but the scintillator panel 30 can also be provided at a distance from the fluorescence detection unit. The X-ray image detector 1b can be, for example, a DR (Digital Radiography) device.

[0075] As shown in Figure 5, the X-ray image detector 1b includes, for example, a scintillator panel 30, a reflector 34, an optical lens 35, a detection unit 36, and a housing 37. The housing 37 is box-shaped, and a scintillator panel 30 is provided at one end. The detection unit 36 ​​is provided at a distance from the scintillator panel 30 and detects fluorescence 103 from the scintillator panel 30. The second portion 3a2 of the multiple columnar crystals 3a contained in the scintillator layer 3 is located on the side of the detection unit 36. The detection unit 36 ​​is provided at the end of the housing 37 that intersects with the end on which the scintillator panel 30 is provided. The detection unit 36 ​​may be the circuit board 2 described above, or it may be a CCD camera or the like. The optical lens 35 is provided on the incident side of the detection unit 36. The reflector 34 is provided at the inclined end of the housing 37 and faces the scintillator panel 30 and the detection unit 36. The reflector 34 is not strictly necessary; the scintillator panel 30 and the detection unit 36 ​​may face each other via an optical lens 35.

[0076] X-rays 100 emitted from the X-ray source 101 pass through the subject 102 and then enter the scintillator panel 30. The X-rays 100 that enter the scintillator panel 30 are converted into fluorescence 103. The fluorescence 103 emitted from the scintillator panel 30 is reflected by the reflector 34 and enters the detection unit 36 ​​via the optical lens 35. The detection unit 36 ​​converts image data information corresponding to the intensity distribution of the incident fluorescence 103 into potential information.

[0077] As described above, the scintillator panel 30 may be bonded to the circuit board 2, or it may be provided at a distance from the fluorescence detection unit 36.

[0078] Next, the effects of the scintillator layer 3 according to this embodiment will be further explained. Figure 6 is a photograph illustrating an X-ray image obtained using a grid. Figure 7 is a photograph illustrating an X-ray image obtained using the scintillator layer 3 according to this embodiment. As shown in Figure 6, using a grid can sometimes cause moiré patterns. When moiré patterns occur, the image quality of the X-ray image deteriorates significantly. As shown in Figure 7, using the scintillator layer 3 according to this embodiment can suppress the occurrence of moiré patterns.

[0079] Figure 8 is a table illustrating the types of images and afterimages that can be measured. Comparative Example 1 is the case of a normal scintillator layer, that is, the entire columnar crystal is formed from CsI with Tl added. For example, the first portion 3a1 is not provided, and the columnar crystal is composed only of the second portion 3a2. Comparative Example 2 is a case where a grid is provided in a normal scintillator layer. Comparative Example 3 is a case where the concentration of Tl in the normal scintillator layer of Comparative Example 1 is different. The scintillator layer (1) and the scintillator layer (2) according to this embodiment differ in the concentration of Tl in the second portion 3a2.

[0080] For the measurement targets exemplified in Figure 8, the subjects were photographed under predetermined shooting conditions, and the image characteristics (sensitivity ratio, MTF ratio (resolution), contrast ratio, moiré) of each X-ray image were determined.

[0081] Figure 9 is a schematic diagram illustrating contrast shooting conditions. As shown in Figure 9, the object to be photographed was a rectangular acrylic plate 104 with a lead object 102a in the center. The thickness of the acrylic plate 104 was 60 mm. The acrylic plate 104 with the object 102a was placed between the X-ray source 102 and the X-ray image detector 1. The X-ray irradiation conditions were 70 kV - 0.0087 mGy.

[0082] Figure 10(a) is a schematic diagram illustrating the conditions for measuring contrast. Figure 10(b) is a schematic diagram illustrating the method for calculating the contrast ratio. As shown in Figure 10(a), in contrast measurement, the luminance on the luminance measurement line 105 passing through the center of the subject 102a was measured. Then, as shown in Figure 10(b), the ratio (H / L) of the luminance level H when the subject 102a is absent to the luminance level L when the subject 102a is present was determined, with a luminance level of 0 as the reference, and the obtained ratio was defined as the contrast ratio.

[0083] Figure 11 is a table illustrating the measurement results of image characteristics. Note that the sensitivity ratio, MTF ratio, and contrast ratio are relative values ​​with the image characteristics of Comparative Example 1 as the baseline (1.00).

[0084] As can be seen from Figure 11, the scintillator layers (1) and (2) according to this embodiment do not cause a significant decrease in sensitivity ratio as in the case of using a grid, and it is possible to improve the MTF ratio and contrast ratio in the same way as when using a grid. In addition, unlike when using a grid, the occurrence of moiré patterns can be suppressed.

[0085] Figure 12 is a table illustrating the sensitivity attenuation ratio. Note that Comparative Example 2 is not included as a comparison because it involves the use of grid. The sensitivity attenuation ratio is a relative value with the sensitivity attenuation ratio of Comparative Example 1 as the reference (1.00).

[0086] As can be seen from Figure 12, with the scintillator layers (1) and (2) according to this embodiment, the sensitivity attenuation ratio corresponding to the concentration of Tl in the second portion 3a2 is greater than the sensitivity attenuation ratio of Comparative Examples 1 and 3 which have similar concentrations of Tl, thus improving X-ray resistance.

[0087] Next, we will explain the afterimage characteristics. Due to the high energy of the incident X-rays 100, the emission characteristics of the scintillator 3 change, and there is a risk that the influence of the previously acquired X-ray image may remain in the currently acquired X-ray image. As a result, the image of the previously acquired subject may appear as an afterimage in the currently acquired X-ray image. If an afterimage is present, it can lead to misjudgment in diagnostics using X-ray images. Therefore, it is necessary to improve the afterimage characteristics by lowering the afterimage ratio (making it less likely for afterimages to appear), which will be discussed later.

[0088] Figures 13(a) and (b) are schematic diagrams illustrating methods for measuring afterimage characteristics. First, as shown in Figure 13(a), the subject 102b is photographed. Next, as shown in Figure 13(b), subject 102c is photographed. The X-ray image of subject 102c may show the afterimage 102b1 of subject 102b. In this case, if we let A be the output value (density) of the region in the X-ray image where the afterimage 102b1 is not present, and B be the output value (density) of the region in the X-ray image where the afterimage 102b1 is present, then the "afterimage ratio (%) = {(AB) / A} × 100" can be expressed as follows.

[0089] Figure 14 is a table illustrating the measurement conditions for the afterimage ratio. Figure 15 is a table illustrating the measurement results of the afterimage ratio. The afterimage ratio is a relative value with the afterimage ratio of Comparative Example 1 as the baseline (1.00).

[0090] As can be seen from Figure 15, the afterimage ratio corresponding to the Tl concentration in the second portion 3a2 can be made significantly lower than the afterimage ratios of Comparative Examples 1 and 3, which have similar Tl concentrations. For example, the afterimage ratio of the scintillator layer (1) according to this embodiment could be set to 65% of the afterimage ratio of Comparative Example 1, which has a similar Tl concentration. For example, the afterimage ratio of the scintillator layer (2) according to this embodiment could be set to 50% of the afterimage ratio of Comparative Example 3, which has a similar Tl concentration.

[0091] As mentioned above, the scintillator layer 3 according to this embodiment also improves the resistance of the scintillator layer 3 to X-rays, thus solving the problems that arise when adding Tl as an activator to the halide CsI (improvement of afterimage characteristics and improvement of resistance to X-rays).

[0092] As described above, the scintillator layer 3 according to this embodiment can solve both the problems of improving afterimage characteristics and improving resistance to X-rays, and it is also possible to improve the image characteristics of X-ray images (sensitivity ratio, MTF ratio (resolution), contrast ratio, moiré).

[0093] Next, we will illustrate another example of a scintillator layer 13 according to a different embodiment. In an X-ray image detector, the intensity of scattered X-rays produced when X-rays pass through an object differs between the central region, which is the center of the scintillator layer formation area, and the peripheral region, which is the outer edge of the scintillator layer formation area. Generally, the intensity of scattered X-rays in the central region is higher than that in the peripheral region.

[0094] In this case, generally, the lengths of the multiple columnar crystals contained in the scintillator layer (the dimensions of the columnar crystals in the direction in which they extend) and the composition of the multiple columnar crystals are approximately the same, resulting in an emission distribution that corresponds to the intensity distribution of scattered X-rays.

[0095] Furthermore, due to the arrangement of the X-ray source and the X-ray image detector, the intensity of the X-rays incident on the scintillator layer differs between the central and peripheral parts of the scintillator layer. Figure 16 is a schematic diagram illustrating the arrangement of an X-ray source and an X-ray image detector. As can be seen from Figure 16, the distance from the X-ray source and the angle of incidence of X-rays entering the scintillator layer differ between the central and peripheral parts of the scintillator layer. For example, if L is the distance from the X-ray source to the center of the scintillator layer, L' is the distance from the X-ray source to the periphery of the scintillator layer, θ is the angle of incidence of the X-rays at the periphery, M is the distance from the center to the periphery, E1 is the intensity of the incident X-rays at the center, and E2 is the intensity of the incident X-rays at the periphery, then the intensity of the incident X-rays can be expressed by the following formula.

[0096]

number

number

[0097] As can be seen from the above formula, the intensity of incident X-rays decreases in proportion to at least the square of the distance from the X-ray source. Thus, the intensity of the incident X-rays changes depending on the distance from the X-ray source and the angle at which the X-rays are incident. As a result, the intensity of X-rays incident on the center of the scintillator layer is higher than that of X-rays incident on the periphery.

[0098] Due to the intensity distribution of scattered X-rays and incident X-rays, concentric sensitivity patches (shading) may appear in X-ray images. Figure 17 is a photograph illustrating sensitivity variations in an X-ray image. As shown in Figure 17, when sensitivity spots occur in an X-ray image, the quality of the X-ray image may be significantly reduced. Therefore, a correction called Flat Field correction is generally applied to reduce sensitivity variations.

[0099] However, flat-field correction also amplifies the noise components contained in the X-ray image. Therefore, the larger the sensitivity variation before correction, the lower the signal-to-noise ratio of the corrected X-ray image will be, as will the dynamic range of the X-ray image.

[0100] Therefore, if sensitivity variations can be reduced without performing flat-field correction, the uniformity of the sensitivity distribution of the X-ray image (BU (Brightness Uniformity)) will improve, resulting in high-quality X-ray images (X-ray images with a high signal-to-noise ratio and a wide dynamic range).

[0101] Therefore, the scintillator layer 13 has the following configuration. Figure 18 is a schematic cross-sectional view illustrating the scintillator layer 13. The scintillator layer 13 can be formed directly on the circuit board 2, for example, similar to the scintillator layer 3 shown in Figure 2 above. The scintillator layer 13 is provided, for example, to cover the region on the substrate 2a where the multiple photoelectric conversion units 2b are provided (effective pixel region). Note that, to avoid complexity, elements other than the scintillator layer 13 are omitted in Figure 18. For example, the X-ray image detector 11 according to this embodiment can be modified by replacing the scintillator layer 3 of the X-ray image detector 1 above with the scintillator layer 13.

[0102] As shown in Figure 18, the scintillator layer 13 contains a plurality of columnar crystals 13a (corresponding to an example of a first columnar crystal) and a plurality of columnar crystals 13b (corresponding to an example of a second columnar crystal). Multiple columnar crystals 13a are located in the central part of the scintillator layer 13. Multiple columnar crystals 13b are located in the peripheral area outside the central part of the scintillator layer 13. When the scintillator layer 13 is viewed from the direction from which the X-rays are incident, the central part of the scintillator layer 13 can be a circular or rectangular region centered on the center of the scintillator layer 13. Furthermore, when the scintillator layer 13 is viewed from the direction from which the X-rays are incident, the ratio of the projected area of ​​the central part to the projected area of ​​the scintillator layer 13 can be approximately 0.5 to 0.75.

[0103] Each of the multiple columnar crystals 13a has a first portion 13a1 and a second portion 13a2. The first portion 13a1 and the second portion 13a2 are integrally formed. The first portion 13a1 is located on the incident side of the X-rays 100 than the second portion 13a2. The first portion 13a1 is positioned in the scintillator layer 13, including the end on the incident side of the X-rays 100. Therefore, scattered X-rays 100a are incident on the second portion 13a2 via the first portion 13a1.

[0104] Similar to the first portion 3a1 described above, the first portion 13a1 contains, for example, a halide CsI and an activator (corresponding to an example of the first activator) Na. In this way, when scattered X-rays 100a are incident on the first portion 13a1, fluorescence with a peak wavelength of around 420 nm is generated. As described above, the peak wavelength of the photoelectric conversion element 2b1's photoreception sensitivity is around 550 nm, so even if fluorescence with a peak wavelength of around 420 nm is incident on the photoelectric conversion element 2b1, it is unlikely to be converted into a signal charge.

[0105] Furthermore, the activator added to the first portion 13a1 should satisfy the same relationship as the activator added to the first portion 3a1 mentioned above: "α(μm) < γ2(μm) ≤ β(μm) ≤ γ1(μm)".

[0106] Similar to the second portion 3a2 described above, the second portion 13a2 contains, for example, a halide CsI and an activator (corresponding to an example of a second activator) Tl. Therefore, when X-rays 100 are incident on the second portion 13a2 via the first portion 13a1, fluorescence with a peak wavelength of around 550 nm can be generated. As a result, the fluorescence intensity can be efficiently detected by the power conversion element 2b1.

[0107] Each of the multiple columnar crystals 13b has a first portion 13b1 and a second portion 13b2. The first portion 13b1 and the second portion 13b2 are integrally formed. The first portion 13b1 is located on the incident side of the X-rays 100 than the second portion 13b2. The first portion 13b1 is located in the scintillator layer 13, including the end on the incident side of the X-rays 100. Therefore, scattered X-rays 100a are incident on the second portion 13b2 via the first portion 13b1.

[0108] The materials and function of the first part 13b1 may be the same as those of the first part 13a1 described above. The materials and function of the second part 13b2 may be the same as those of the second part 13a2 described above.

[0109] Here, as shown in Figure 18, in the direction in which the columnar crystal extends, the length of the second portion 13b2 of the columnar crystal 13b is longer than the length of the second portion 13a2 of the columnar crystal 13a. Therefore, when X-rays 100 of the same intensity are incident, the fluorescence intensity generated in the second portion 13b2 is higher than the fluorescence intensity generated in the second portion 13a2.

[0110] Furthermore, in the direction in which the columnar crystal extends, the length of the columnar crystal 13b (the sum of the length of the first portion 13b1 and the second portion 13b2) can be made approximately the same as the length of the columnar crystal 13a (the sum of the length of the first portion 13a1 and the second portion 13a2).

[0111] As mentioned above, the intensity of incident X-rays in the central part of the scintillator layer 13 is higher than that of incident X-rays in the peripheral part. However, the high-intensity X-rays are incident on the shorter second part 13a2, and the low-intensity X-rays are incident on the longer second part 13b2. Therefore, the difference between the fluorescence intensity generated in the second part 13a2 and the fluorescence intensity generated in the second part 13b2 can be reduced. Furthermore, if the first portion 13a1 and the first portion 13b1 are provided, scattered X-rays 100a can be absorbed.

[0112] Therefore, by using a scintillator layer 13, it is possible to suppress the occurrence of sensitivity variations in the X-ray image due to the intensity distribution of the incident X-rays 100 and the intensity distribution of the scattered X-rays 100a. Furthermore, by omitting or reducing flat field correction, the uniformity of the sensitivity distribution of X-ray images can be improved, ultimately resulting in high-quality X-ray images (X-ray images with a high signal-to-noise ratio and a wide dynamic range).

[0113] Figure 19 is a schematic cross-sectional view illustrating a scintillator layer 13 according to another embodiment. The scintillator layer 13 can be provided on the substrate 31, for example, in the same way as the scintillator panel 30 shown in Figure 4 above. Note that, to avoid complexity, elements other than the scintillator layer 13 are omitted in Figure 19. For example, the X-ray image detector 11a (scintillator panel 30a) according to this embodiment can be modified by replacing the scintillator layer 3 of the scintillator panel 30 above with the scintillator layer 13.

[0114] As shown in Figure 19, the scintillator layer 13 contains a plurality of columnar crystals 13a and a plurality of columnar crystals 13b. Multiple columnar crystals 13a are located in the central part of the scintillator layer 13. Multiple columnar crystals 13b are located in the peripheral area outside the central part of the scintillator layer 13.

[0115] In the case of the scintillator layer 13 shown in Figure 18, X-rays are incident on the scintillator layer 13 from the end opposite to the circuit board 2. Therefore, a first portion 13a1 and a first portion 13b1 are provided on the end of the scintillator layer 13 opposite to the circuit board 2.

[0116] In contrast, in the case of the scintillator layer 13 shown in Figure 19, X-rays are incident from the end of the scintillator layer 13 on the substrate 31 side. For this reason, a first portion 13a1 and a first portion 13b1 are provided on the end of the scintillator layer 13 on the substrate 31 side.

[0117] Even when the scintillator layer 13 is provided on the scintillator panel 30a, the effects of the scintillator layer 13 described above can still be enjoyed.

[0118] As explained above, if the central region of the scintillator layer 13 formation area is defined as the central part and the outer peripheral region as the peripheral part, then, in the scintillator layer 3 shown in Figure 2, if the following features (1) and (2) are added to the scintillator layer 3 as shown in Figures 18 and 19, the sensitivity of the peripheral part of the scintillator layer 3 will increase compared to the case where the film thickness distribution in the in-plane direction of the formation area of ​​the scintillator layer 3 is uniform. This reduces the concentric sensitivity spots that occur in the X-ray image as shown in Figure 17, making it possible to obtain a higher quality X-ray image. (1) In-plane film thickness distribution of region A including the first portion 3a1: center > periphery (2) In-plane film thickness distribution of region B including the second portion 3a2: center < peripheral

[0119] In addition, similar to the case of the X-ray image detector 1b illustrated in Figure 5, the scintillator panel 30a may be provided at a distance from the fluorescence detection unit 36.

[0120] Next, we will further explain the effect of the scintillator layer 13. Figure 20 is a table illustrating the types of objects for which image characteristics and afterimage characteristics are measured. Comparative Examples 1 to 3 are the same as those described in Figure 8. Furthermore, the contrast shooting conditions, contrast measurement conditions, and contrast ratio calculation method are the same as those described in Figures 9, 10(a), and 10(b).

[0121] Figure 21 is a table illustrating the measurement results of image characteristics. Note that the sensitivity ratio, MTF ratio, and contrast ratio are relative values ​​with the image characteristics of Comparative Example 1 as the baseline (1.00).

[0122] As can be seen from Figure 21, the scintillator layers (3) and (4) according to this embodiment do not cause a significant decrease in sensitivity ratio as in the case of using a grid, and it is possible to improve the MTF ratio and contrast ratio in the same way as when using a grid. In addition, unlike when using a grid, the occurrence of moiré patterns can be suppressed.

[0123] Figure 22 is a table illustrating the sensitivity attenuation ratio. Note that Comparative Example 2 is not included as a comparison because it involves the use of grid. The sensitivity attenuation ratio is a relative value with the sensitivity attenuation ratio of Comparative Example 1 as the reference (1.00).

[0124] As can be seen from Figure 22, with the scintillator layers (3) and (4) according to this embodiment, the sensitivity attenuation ratio corresponding to the concentration of Tl in the second portions 13a2 and 13b2 is greater than the sensitivity attenuation ratio of Comparative Examples 1 and 3 which have similar concentrations of Tl, thus improving X-ray resistance.

[0125] Next, I will explain the afterimage characteristics. The measurement method for afterimage characteristics is the same as that described in Figures 13(a) and (b). The measurement conditions for the afterimage ratio are the same as those described in Figure 14. Figure 23 is a table illustrating the measurement results of the afterimage ratio. The afterimage ratio is a relative value with the afterimage ratio of Comparative Example 1 as the baseline (1.00).

[0126] As can be seen from Figure 23, the afterimage ratio corresponding to the Tl concentration in the second portions 13a2 and 13b2 can be made significantly lower than the afterimage ratios of Comparative Examples 1 and 3, which have similar Tl concentrations. For example, the afterimage ratio of the scintillator layer (3) according to this embodiment could be set to 65% of the afterimage ratio of Comparative Example 1, which has a similar Tl concentration. For example, the afterimage ratio of the scintillator layer (4) according to this embodiment could be set to 50% of the afterimage ratio of Comparative Example 3, which has a similar Tl concentration.

[0127] As described above, the scintillator layer 13 according to this embodiment solves both the problem of improving afterimage characteristics and improving resistance to X-rays, similar to the scintillator layer 3 described above, and also makes it possible to improve the image characteristics of X-ray images (sensitivity ratio, MTF ratio (resolution), contrast ratio, moiré).

[0128] Figure 24 is a table illustrating the measurement results of BU(%). Furthermore, if the sensitivity of the central part of the X-ray image is a and the sensitivity of the peripheral part of the X-ray image is b, BU(%) can be expressed as "BU(%) = (b / a) × 100". As can be seen from Figure 24, the scintillator layer 13 according to this embodiment has an improved BU compared to Comparative Example 1, which reduces sensitivity variations in the X-ray image and improves the uniformity of the sensitivity distribution of the X-ray image, resulting in a high-quality X-ray image (an X-ray image with a high signal-to-noise ratio and a wide dynamic range).

[0129] Next, we will illustrate the method for manufacturing an X-ray image detector according to this embodiment. Since known techniques can be applied to methods other than the manufacturing method of the scintillator layer, the manufacturing method of the scintillator layer will be described below. Figures 25(a) and (b) are schematic diagrams illustrating the manufacturing method of the scintillator layer 3. To avoid complexity, the illustration of the mask defining the formation range of the scintillator layer 3 has been omitted. As shown in Figures 25(a) and (b), the scintillator manufacturing apparatus 200 is provided with a chamber 201, a first crucible 202, a second crucible 203, a third crucible 204, and a rotating section 205. The first crucible 202, the second crucible 203, the third crucible 204, and the rotating section 205 can be installed inside the chamber 201. The second crucible 203 and the third crucible 204 can be installed side by side with the first crucible 202 in between. The openings of the first crucible 202, the second crucible 203, and the third crucible 204 each face the rotating section 205.

[0130] The first crucible 202 side of the rotating part 205 is detachably held by the circuit board 2 or the substrate 31 of the scintillator panel 30. The rotating part 205 also rotates the held circuit board 2 or substrate 31. The first crucible 202 can contain a halide (e.g., CsI). The second crucible 203 can contain an activating material (for example, TlI). The third crucible 204 can contain an activating material (for example, NaI).

[0131] First, the gas inside chamber 201 is evacuated using a pump so that the internal pressure is 1.0 Pa or less. Next, the rotating part 205 rotates the held circuit board 2 or board 31. At this time, a heater is used to raise the temperature of the circuit board 2 or board 31 to approximately 150°C.

[0132] To form the first portion 3a1, the first crucible 202 and the third crucible 204 are heated with a heater, as shown in Figure 25(b). At this time, the temperature of the first crucible 202 is set to approximately 700°C, and the temperature of the third crucible 204 is set to approximately 600°C. The vapor of the halide contained in the first crucible 202 and the vapor of the activator contained in the third crucible 204 are mixed in the space inside the chamber 201, and the first portion 3a1 of the columnar crystal 3a is formed by the mixed vapor. In other words, the first portion 3a1 of the columnar crystal is formed by vapor phase growth. The concentration of the activator in the first portion 3a1 can be controlled, for example, by the temperature of the third crucible 204.

[0133] To form the second portion 3a2, the first crucible 202 and the second crucible 203 are heated with a heater, as shown in Figure 25(a). At this time, the temperature of the first crucible 202 is set to approximately 700°C, and the temperature of the second crucible 203 is set to approximately 400°C. The vapor of the halide contained in the first crucible 202 and the vapor of the activator contained in the second crucible 203 are mixed in the space inside the chamber 201, and the second portion 3a1 of the columnar crystal 3a is formed by the mixed vapor. That is, the second portion 3a2, which is a columnar crystal, is formed by vapor phase growth. Furthermore, the concentration of the activator in the second portion 3a2 can be controlled, for example, by the temperature of the second crucible 203.

[0134] As shown in Figures 2 and 4, in each of the multiple columnar crystals 3a, the first portion 3a1 is located closer to the X-ray 100 incidence than the second portion 3a2. Therefore, when forming the scintillator layer 3 on the circuit board 2, first, the second portion 3a2 is formed on the multiple photoelectric conversion units 2b, and then the first portion 3a1 is formed on the second portion 3a2.

[0135] Furthermore, when forming a scintillator layer 3 on the substrate 31 (when forming a scintillator panel 30), first, a first portion 3a1 is formed on the substrate 31, and then, a second portion 3a2 is formed on top of the first portion 3a1.

[0136] In these cases, the first portion 3a1 and the second portion 3a2 contain the same halide, so their crystal structure and refractive index can be made identical.

[0137] When the first portion 3a1 and the second portion 3a2 are formed, heating of the first crucible 202 and the second crucible 203 or third crucible 204 is stopped, and the temperature inside the chamber 201 is returned to room temperature. Subsequently, nitrogen gas is supplied into the chamber 201 to return the pressure inside the chamber 201 to atmospheric pressure. Next, the circuit board 2 on which the scintillator layer 3 is formed, or the substrate 31 on which the scintillator layer 3 is formed, is removed from the chamber 201. As described above, the scintillator layer 3 according to this embodiment can be manufactured.

[0138] The above examples illustrate methods for manufacturing the scintillator layer 3, but the method for manufacturing the scintillator layer 13 can be similar.

[0139] However, in the case of the scintillator layer 3, the lengths of the multiple first portions 3a1 in the direction in which the columnar crystals 3a extend were approximately the same in the central and peripheral parts of the scintillator layer 3, and the lengths of the multiple second portions 3a2 in the direction in which the columnar crystals 3a extend were approximately the same.

[0140] In contrast, in the case of the scintillator layer 13, the lengths of the first portion 13a1 and the first portion 13b1 in the direction in which the columnar crystals 13a extend are different in the central and peripheral portions of the scintillator layer 13. Also, the lengths of the second portion 13a2 and the second portion 13b2 in the direction in which the columnar crystals 13a extend are different.

[0141] Therefore, when manufacturing the scintillator layer 13, the scintillator manufacturing apparatus 200 may also be further equipped with a mask 206 or the like to control the scattering of steam. For example, as shown in Figures 25(a) and (b), a mask 206 can be provided inside the chamber 201 between the rotating part 205 and the first crucible 202, the second crucible 203, and the third crucible 204. The mask 206 can be made movable, for example, in a direction intersecting the direction from the rotating part 205 toward the second crucible 203.

[0142] Furthermore, masks 206a can be provided at the ends of the first crucible 202, the second crucible 203, and the third crucible 204. The masks 206a can, for example, have multiple holes with different opening diameters, and the scattering of steam can be controlled by the opening diameter of the holes, the number of holes, the arrangement of the holes, etc.

[0143] Figure 26(a) is a schematic bottom view of mask 206a1. Figure 26(b) is a schematic diagram illustrating the scintillator layer 13 formed using the mask 206a1. As shown in Figure 26(a), the mask 206a1 has an opening 206a1a. When the scintillator layer 13 is deposited using such a mask 206a1, the film thickness distribution is as shown in Figure 26(b).

[0144] Figure 26(c) is a schematic bottom view of mask 206a2. Figure 26(d) is a schematic diagram illustrating the scintillator layer 13 formed using mask 206a2. As shown in Figure 26(c), the mask 206a2 has multiple openings 206a2a with different opening dimensions. When the scintillator layer 13 is deposited using such a mask 206a2, the film thickness distribution is as shown in Figure 26(d).

[0145] The morphology and arrangement of masks 206, 206a, 206a1, and 206a2 can be appropriately determined through experiments and simulations.

[0146] As described above, the method for manufacturing the scintillator layer 3 according to this embodiment may include the following steps. A first step involves growing a first portion of a columnar crystal in the vapor phase by mixing a vapor of a halide with a vapor of a first activator in an atmosphere reduced to atmospheric pressure. A second step involves mixing a halogen vapor with a second activator vapor in an atmosphere reduced to atmospheric pressure to grow a second portion of a columnar crystal in the vapor phase. In this case, the second step is performed after the first step, or the first step is performed after the second step.

[0147] As described above, the method for manufacturing an X-ray image detector according to this embodiment may include the following steps. A process for forming a circuit board in which multiple pixel electrodes and switching elements are arranged in one or two dimensions on a planar substrate. A process for forming a semiconductor photodetector that is in contact with a circuit board and converts light incident from the outside into an electrical signal. A process for forming a scintillator layer that is in contact with a semiconductor photodetector and converts X-rays incident from the outside into light. Furthermore, in the process of forming the scintillator layer, the scintillator layer is formed by a multi-component vapor phase growth method using at least three different materials as material sources.

[0148] As described above, the method for manufacturing the scintillator panel 30 according to this embodiment may include the following steps. A first step involves mixing a halogen vapor with a first activator vapor in an atmosphere reduced to atmospheric pressure, and growing a first portion in the vapor phase on one side of a substrate. A second step involves mixing a halogenated vapor with a second activator vapor in an atmosphere reduced to atmospheric pressure, and vapor-depositing the second portion onto the end of the first portion opposite to the substrate side. In other words, in the process of forming the scintillator layer, the scintillator layer is formed by a multi-component vapor phase growth method using at least three different materials as material sources.

[0149] The embodiments described above are illustrative examples. However, the present invention is not limited to these descriptions. With respect to the embodiments described above, those in which a person skilled in the art has appropriately added, deleted, or modified components, or added, omitted, or modified processes, are also included within the scope of the present invention, as long as they retain the features of the present invention. Furthermore, the elements of each embodiment described above can be combined as much as possible, and these combinations are also included within the scope of the present invention insofar as they include the features of the present invention.

[0150] The following are additional notes regarding the embodiments described above.

[0151] (Note 1) A scintillator panel comprising an X-ray-transmitting support substrate and a scintillator layer in contact with the support substrate that converts incident X-rays from the outside into light, The scintillator layer is provided with regions having different emission wavelengths in the direction of its film thickness, and A scintillator panel that satisfies the following equation, where α and β are the peak wavelengths of the light emission centers in each region with different emission wavelengths, respectively, and γ1 and γ2 are the upper and lower limits of the full width at half maximum of the light receiving sensitivity of the photodetector element in the detector into which the light from the scintillator layer is incident. α < γ2 ≤ β ≤ γ1

[0152] (Note 2) The scintillator panel according to Appendix 1, wherein, in the film thickness direction of the scintillator layer, the peak wavelength of the emission center in region A on the X-ray incident side is α, and the peak wavelength of the emission center in region B on the X-ray non-incident side is β.

[0153] (Note 3) The scintillator layer consists of a phosphor having a strip-shaped columnar crystalline structure containing an activator in the halogen CsI (cesium iodide), and region A contains Na (sodium) as an activator, and region B contains Tl (thallium) as an activator. The concentration of Na in the film thickness direction is 1.0 × 10 -3 mass% or more, 10.0×10 -3 It is less than or equal to mass%. The scintillator panel according to Appendix 1 or 2, wherein the concentration of Tl in the film thickness direction is 0.1 mass% or more and 2.0 mass% or less.

[0154] (Note 4) A scintillator panel according to any one of the appendices 1 to 3, wherein the in-plane film thickness distribution of region A and region B are inversely related.

[0155] (Note 5) When the central region of the scintillator layer formation area is defined as the central part, and the area surrounding the scintillator layer formation area is defined as the peripheral part, The in-plane film thickness distribution of region A is such that the central portion is thicker than the peripheral portion. The film thickness distribution in the in-plane direction of region B is such that the peripheral portion is thicker than the central portion, as described in any one of the scintillator panels described in Appendix 1 to 4.

[0156] (Note 6) The region A is a scintillator panel according to any one of the appendices 1 to 5, comprising at least 5% and 50% of the film thickness of the scintillator layer.

[0157] (Note 7) The central portion of the scintillator layer occupies 50% or more of a concentric or rectangular area based on the center of the scintillator layer formation region, as described in Appendix 5.

[0158] (Note 8) The scintillator panel according to any one of the appendices 1 to 7, wherein the support substrate is made of a material consisting of elements lighter than transition metal elements.

[0159] (Note 9) A method for manufacturing a scintillator panel, comprising the step of forming a scintillator layer on an X-ray-transmitting support substrate that converts externally incident X-rays into light, wherein the scintillator layer is formed by a multi-component vapor phase growth method using at least three or more different materials as material sources. [Explanation of Symbols]

[0160] 1 X-ray image detector, 2 circuit board, 2a substrate, 2b photoelectric conversion unit, 3 scintillator, 3a columnar crystal, 3a1 first part, 3a2 second part, 13 scintillator, 13a columnar crystal, 13a1 first part, 13a2 second part, 13b columnar crystal, 13b1 first part, 13b2 second part, 30 scintillator panel, 31 substrate, 100 X-rays, 100a scattered X-rays

Claims

1. In a scintillator panel comprising an X-ray-transmitting support substrate and a scintillator layer in contact with the support substrate that converts X-rays incident from the outside into light, The scintillator layer is provided with regions having different emission wavelengths in the direction of its film thickness, and A scintillator panel that satisfies the following equation, where α and β are the peak wavelengths of the light emission centers in each region with different emission wavelengths, respectively, and γ1 and γ2 are the upper and lower limits of the full width at half maximum of the light receiving sensitivity of the photodetector element in the detector into which the light from the scintillator layer is incident, respectively. α < γ2 ≤ β ≤ γ1

2. The scintillator panel according to claim 1, wherein, in the film thickness direction of the scintillator layer, the peak wavelength of the emission center in the X-ray incident region A is α, and the peak wavelength of the emission center in the X-ray non-incident region B is β.

3. The scintillator layer consists of a phosphor having a strip-shaped columnar crystalline structure containing an activator in CsI (cesium iodide), which is a halide, and region A contains Na (sodium) as an activator, and region B contains Tl (thallium) as an activator. The concentration of Na in the film thickness direction is 1.0 × 10 －３ mass% or more, 10.0×10 －３ It is less than or equal to mass%, The scintillator panel according to claim 1 or 2, wherein the concentration of Tl in the film thickness direction is 0.1 mass% or more and 2.0 mass% or less.

4. The scintillator panel according to claim 1 or 2, wherein the in-plane film thickness distributions of region A and region B are opposite.

5. When the central region of the scintillator layer formation area is defined as the central part, and the area surrounding the scintillator layer formation area is defined as the peripheral part, The film thickness distribution in the in-plane direction of region A is such that the central portion is thicker than the peripheral portion. The scintillator panel according to claim 1 or 2, wherein the film thickness distribution in the in-plane direction of region B is such that the peripheral portion is thicker than the central portion.

6. The scintillator panel according to claim 1 or 2, wherein region A comprises at least 5% and 50% of the film thickness of the scintillator layer.

7. The scintillator panel according to claim 5, wherein the central portion of the scintillator layer occupies 50% or more of a concentric or rectangular region based on the center of the scintillator layer formation region.

8. The scintillator panel according to claim 1 or 2, wherein the support substrate is made of a material composed of elements lighter than transition metal elements.

9. A method for manufacturing a scintillator panel, comprising the step of forming a scintillator layer on an X-ray-transmitting support substrate that converts externally incident X-rays into light, wherein the scintillator layer is formed by a multi-component vapor phase growth method using at least three or more different materials as material sources.

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