Radiation detector

Abstract

To provide a radiation detector capable of achieving synchronization with radiation ray on the basis of incident radiation ray and continuously detecting radiation ray.SOLUTION: A radiation detector includes a base material 2a, a plurality of scanning lines G, a plurality of signal lines U, a plurality of pixels PX, a drive circuit 11a, a detection circuit 11b, a detection part DT, and a control circuit. Each pixel PX includes a switching element, a pixel electrode, and a first conversion part. The detection part DT detects radiation ray and generates a radiation detection signal. The control circuit controls output of an image data signal on the basis of the radiation ray detection signal.SELECTED DRAWING: Figure 8

Description

[Technical Field]

[0001] FIELD OF THE INVENTION An embodiment of the present invention relates to a radiation detector. [Background technology]

[0002] As a radiation detector, for example, an X-ray detector (a flat panel X-ray detector) is known. Radiation detectors are classified into direct conversion type and indirect conversion type radiation detectors depending on the method of converting radiation into an electrical signal. Direct conversion type radiation detectors convert electric charges generated in a semiconductor by energy provided by radiation into an electrical signal by collecting them from electrodes provided at the edge of the semiconductor. Indirect conversion type radiation detectors convert incident radiation into light using a scintillator layer, and then convert the light from the scintillator layer into electric charges within the semiconductor.

[0003] An indirect conversion type radiation detector includes, for example, a conversion substrate having a photoelectric conversion unit that converts light into signal charges and a thin-film transistor that switches between accumulating and releasing the signal charges, and a scintillator layer that is provided on the conversion substrate and converts radiation into fluorescence. [Prior art documents] [Patent documents]

[0004] [Patent Document 1] Japanese Patent Application Laid-Open No. 2008-132216 Summary of the Invention [Problem to be solved by the invention]

[0005] This embodiment provides a radiation detector that can continuously detect radiation in synchronization with the incident radiation. [Means for solving the problem]

[0006] The radiation detector according to one embodiment includes: A substrate; a plurality of scanning lines disposed above the substrate; a plurality of signal lines provided above the substrate; a plurality of pixels provided above the substrate, each of the pixels including: a switching element electrically connected to a corresponding one of the plurality of scanning lines and a corresponding one of the plurality of signal lines; a pixel electrode electrically connected to the switching element; and a first conversion unit electrically connected to the pixel electrode, provided in a region overlapping the pixel electrode, and configured to convert incident radiation into an image data signal; a driving circuit electrically connected to the plurality of scanning lines and configured to sequentially scan the plurality of scanning lines; a detection circuit electrically connected to the plurality of signal lines and detecting the image data signals generated by the plurality of pixels; a detection unit that detects radiation and generates a radiation detection signal; and a control circuit for controlling the output of the image data signal based on the radiation detection signal. [Brief explanation of the drawings]

[0007] [Figure 1] FIG. 1 is a cross-sectional view showing an X-ray detector according to a comparative example. [Figure 2] FIG. 2 is a perspective view showing a support substrate, an X-ray detection panel, a circuit board, a plurality of FPCs, and an image transmission unit of the X-ray detector. [Figure 3] FIG. 3 is an enlarged cross-sectional view showing a part of the X-ray detection panel. [Figure 4] FIG. 4 is a circuit diagram showing the X-ray detection panel, the circuit board, and a plurality of FPCs. [Figure 5] FIG. 5 is a plan view showing the X-ray detection panel, and is a diagram for explaining the positional relationship between the effective area and the scintillator layer. [Figure 6] FIG. 6 is a cross-sectional view showing the X-ray detection panel of FIG. 5 taken along line VI-VI, and also showing an FPC. [Figure 7]FIG. 7 is a circuit diagram showing the conversion board, circuit board, image transmission unit, and power supply of the X-ray detector. [Figure 8] FIG. 8 is a circuit diagram showing the conversion board, circuit board, image transmission unit, power supply, and monitor unit of the X-ray detector according to the first embodiment. [Figure 9] FIG. 9 is a circuit diagram showing a conversion board, a circuit board, an image transmission unit, a power supply, a monitor unit, and a power supply of an X-ray detector according to a modified example of the first embodiment. [Figure 10] FIG. 10 is a flowchart for explaining an X-ray imaging method applied to the X-ray detector of the first embodiment. [Figure 11] FIG. 11 is a flowchart illustrating the X-ray imaging method, following FIG. [Figure 12] FIG. 12 is a timing chart showing the scanning signal, the X-ray detection signal, the image data signal to be output to the outside, the continuous pulse X-rays, and the current generated based on the continuous pulse X-rays before and after detecting the continuous pulse X-rays in the first embodiment. [Figure 13] FIG. 13 is a timing chart showing the scanning signal, the X-ray detection signal, the image data signal to be output to the outside, the single X-ray, and the current generated based on the single X-ray before and after detecting the single X-ray in the first embodiment. [Figure 14] FIG. 14 is a timing chart showing the scanning signal, the X-ray detection signal, the image data signal to be output to the outside, the continuous X-rays, and the current generated based on the continuous X-rays before and after detecting the continuous X-rays in the first embodiment. [Figure 15] FIG. 15 is a circuit diagram showing a part of the conversion board of the X-ray detector according to the second embodiment. [Figure 16] FIG. 16 is an enlarged cross-sectional view of a part of the conversion board shown in FIG. [Figure 17] FIG. 17 is an enlarged cross-sectional view showing a part of a conversion substrate of an X-ray detector according to a modified example of the second embodiment. DETAILED DESCRIPTION OF THE INVENTION

[0008] Hereinafter, embodiments, modifications, and comparative examples of the present invention will be described with reference to the drawings. (Comparative Example) First, the configuration of an X-ray detector 1 and an X-ray imaging method according to a comparative example will be described. FIG. 1 is a cross-sectional view showing the X-ray detector 1 according to this comparative example. The X-ray detector 1 is an X-ray image detector, and is an X-ray flat panel detector that uses an X-ray detection panel. The X-ray detector 1 is portable and can be used for medical purposes in hospital rooms or outdoors. Because the communication of the X-ray detector 1 is wireless, the X-ray detector 1 that detects X-rays can communicate wirelessly with an X-ray tube device that emits X-rays. The X-ray detector 1 of this modified example is an indirect conversion type X-ray detector.

[0009] 1, the X-ray detector 1 includes an X-ray detection module 10, a support substrate 12, spacers 9a, 9b, 9c, and 9d, a housing 51, an entrance window 52, etc. The X-ray detection module 10 includes an X-ray detection panel PNL, a circuit board 11, an FPC (flexible printed circuit board) 2e1, etc. The X-ray detection panel PNL is located between the support substrate 12 and the entrance window 52. The X-ray detection panel PNL includes a moisture-proof cover 7 facing the entrance window 52.

[0010] The entrance window 52 is attached to the opening of the housing 51. The entrance window 52 allows X-rays to pass through. Therefore, the X-rays pass through the entrance window 52 and enter the X-ray detection panel PNL. The entrance window 52 is formed in a plate shape and has the function of protecting the inside of the housing 51. It is desirable that the entrance window 52 be formed thin and made of a material with low X-ray absorption rate. This makes it possible to reduce scattering of X-rays and attenuation of the X-ray amount that occur at the entrance window 52. As a result, a thin and lightweight X-ray detector 1 can be realized. The X-ray detection module 10, the support substrate 12, etc. are housed inside a space surrounded by a housing 51 and an entrance window 52.

[0011] The X-ray detection panel PNL is made by laminating thin members, and is therefore light and has low mechanical strength. For this reason, the X-ray detection panel PNL is fixed to one flat surface of the support substrate 12 via an adhesive sheet. The support substrate 12 is formed into a plate shape from, for example, an aluminum alloy, and has the strength required to stably hold the X-ray detection panel PNL. This makes it possible to prevent damage to the X-ray detection panel PNL when external vibrations or impacts are applied to the X-ray detector 1.

[0012] A circuit board 11 is fixed to the other surface of the support board 12 via spacers 9a and 9b. By using the spacers 9a and 9b, an electrical insulation distance can be maintained between the support board 12, which is mainly made of metal, and the circuit board 11. The circuit board 11 is fixed to the inner surface of the housing 51 via spacers 9c and 9d. By using the spacers 9c and 9d, an electrical insulation distance can be maintained between the housing 51, which is mainly made of metal, and the circuit board 11. The housing 51 supports the support board 12 and the like via the circuit board 11 and the spacers 9a, 9b, 9c, and 9d.

[0013] A connector corresponding to the FPC 2e1 is mounted on the circuit board 11, and the FPC 2e1 is electrically connected to the circuit board 11 via the connector. A thermocompression bonding method using an ACF (anisotropic conductive film) is used to connect the FPC 2e1 and the X-ray detection panel PNL. This method ensures electrical connection between the multiple fine pads of the X-ray detection panel PNL and the multiple fine pads of the FPC 2e1, and physically fixes the FPC 2e1 to the X-ray detection panel PNL. The pads of the X-ray detection panel PNL will be described later.

[0014] As described above, the circuit board 11 is electrically connected to the X-ray detection panel PNL via the connector, the FPC 2e1, etc. The circuit board 11 electrically drives the X-ray detection panel PNL and electrically processes output signals from the X-ray detection panel PNL.

[0015] Fig. 2 is a perspective view showing the support substrate 12, X-ray detection panel PNL, circuit board 11, multiple FPCs 2e1 and 2e2, and image transmission unit 4 of the X-ray detector 1 according to this comparative example. Note that Fig. 2 does not show all components of the X-ray detector 1. Some components of the X-ray detector 1, such as a bonded body described below, are omitted from Fig. 2. With regard to the conversion substrate 2, an effective area (effective area DA described below), which is the detection area, is shown, but a non-detection area (non-detection area NDA described below) is not shown.

[0016] 2, the X-ray detection panel PNL includes a conversion substrate 2, a scintillator layer 5, etc. The conversion substrate 2 includes a base material 2a, a plurality of scanning lines (or gate lines) G, a plurality of signal lines (or data lines) U, a plurality of thin-film transistors 13a as a plurality of switching elements, a plurality of thin-film photodiodes 15a as a plurality of conversion elements, etc. Note that the numbers, arrangements, etc. of the scanning lines G, signal lines U, thin-film transistors 13a, thin-film photodiodes 15a, and FPCs 2e1, 2e2 are not limited to the example in FIG.

[0017] The plurality of scanning lines G are provided above the base material 2a, extend in a row direction X as a first direction, and are arranged at predetermined intervals in a column direction Y intersecting the first direction. The plurality of signal lines U are provided above the base material 2a, extend in the column direction Y as a second direction, intersect with the plurality of scanning lines G, and are arranged at predetermined intervals in the row direction X.

[0018] A plurality of pixels PX are provided on the upper side of one surface of the base material 2a. The plurality of pixels PX are provided in a rectangular region defined by scanning lines G and signal lines U. One pixel PX corresponds to one pixel of an X-ray image. The plurality of pixels PX are arranged in a matrix in the row direction X and column direction Y. From the above, the conversion substrate 2 is an array substrate.

[0019] Each pixel PX has a thin-film photodiode 15a, a thin-film transistor 13a, etc. Hereinafter, the thin-film photodiode 15a will be referred to as a TFD 15a, and the thin-film transistor 13a will be referred to as a TFT 13a. The TFT 13a is provided above the substrate 2a and is electrically connected to a corresponding one of the multiple scanning lines G and a corresponding one of the multiple signal lines U. The TFD 15a is provided above the substrate 2a and is electrically connected to the TFT 13a.

[0020] The scanning lines G are electrically connected to the circuit board 11 via the FPC 2e1. The circuit board 11 applies scanning signals S1 to the scanning lines G via the FPC 2e1. The signal lines U are electrically connected to the circuit board 11 via the FPC 2e2. An image data signal S2 (charges accumulated in the TFD 15a) converted by the TFD 15a and a scintillator layer 5 (described later) is transmitted to the circuit board 11 via the TFT 13a, the signal lines U, and the FPC 2e2.

[0021] The X-ray detector 1 includes an image transmission unit 4. The image transmission unit 4 is connected to a circuit board 11 via wiring 4a. The image transmission unit 4 may be incorporated into the circuit board 11. The image transmission unit 4 generates an X-ray image based on image data signals that have been converted into digital signals by a plurality of analog-to-digital converters (not shown). The generated X-ray image data is output from the image transmission unit 4 to an external device.

[0022] Fig. 3 is an enlarged cross-sectional view showing a portion of the X-ray detection panel PNL according to this comparative example. As shown in Fig. 3, the conversion substrate 2 has a base material 2a, a plurality of pixels PX, and insulating layers 21, 22, 23, 24, and 25. The plurality of pixels PX are arranged above the base material 2a and located in an effective area DA that is effective for detecting radiation (e.g., X-rays). Each pixel PX has a first conversion unit 2b including a TFT 13a, a conductive layer CL as a pixel electrode, and a TFD 15a.

[0023] The TFT 13a has a gate electrode GE, a semiconductor layer SC, a source electrode SE, and a drain electrode DE. The TFD 15a is composed of a thin-film photodiode.

[0024] The substrate 2a has a plate-like shape and is made of an insulating material. Examples of the insulating material include glass such as alkali-free glass. In this comparative example, the substrate 2a is made of glass, but it may also be made of an organic insulating material such as resin. The planar shape of the substrate 2a is, for example, rectangular. The thickness of the substrate 2a is, for example, 0.5 to 0.7 mm. The insulating layer 21 is provided on the substrate 2a.

[0025] A gate electrode GE is formed on the insulating layer 21. The gate electrode GE is electrically connected to the scanning line G. An insulating layer 22 is provided on the insulating layer 21 and the gate electrode GE. A semiconductor layer SC is provided on the insulating layer 22 and faces the gate electrode GE. The semiconductor layer SC is formed of a semiconductor material such as amorphous silicon as an amorphous semiconductor or polycrystalline silicon as a polycrystalline semiconductor.

[0026] A source electrode SE and a drain electrode DE are provided on the insulating layer 22 and the semiconductor layer SC. The gate electrode GE, the source electrode SE, the drain electrode DE, the scanning line G, and the signal line U are formed using a low-resistance metal such as aluminum or chromium.

[0027] The source electrode SE is electrically connected to the source region of the semiconductor layer SC. The source electrode SE is also electrically connected to the signal line U. The drain electrode DE is electrically connected to the drain region of the semiconductor layer SC.

[0028] The insulating layer 23 is provided on the insulating layer 22, the semiconductor layer SC, the source electrode SE, and the drain electrode DE. The conductive layer CL is formed above the insulating layer 23 and is electrically connected to the drain electrode DE of the TFT 13a. The TFD 15a is formed on the conductive layer CL and is electrically connected to the conductive layer CL. The TFD 15a is formed through a film formation process and a patterning process using a dry etching method.

[0029] The insulating layer 24 is disposed on the insulating layer 23, the conductive layer CL, and the TFD 15a. The bias line BL1 is disposed on the insulating layer 24 and is connected to the TFD 15a through a contact hole formed in the insulating layer 24. The insulating layer 25 is disposed on the insulating layer 24 and the bias line BL1.

[0030] The insulating layers 21, 22, 23, 24, and 25 are made of insulating materials such as inorganic insulating materials and organic insulating materials. Examples of inorganic insulating materials include oxide insulating materials, nitride insulating materials, and oxynitride insulating materials. Examples of organic insulating materials include resins.

[0031] The scintillator layer 5 is provided on the conversion substrate 2 (multiple TFDs 15a). The scintillator layer 5 is located at least in the active area DA and covers the upper parts of the multiple TFDs 15a. The scintillator layer 5 is configured to convert incident X-rays (radiation) into visible light (fluorescence).

[0032] The TFD 15a converts visible light incident from the scintillator layer 5 side into electric charges whose magnitude corresponds to the intensity of the visible light. The converted electric charges are stored in the TFD 15a. One of the TFDs 15a electrically connected to the conductive layer CL and a portion of the scintillator layer 5 overlapping the TFD 15a constitute the first conversion unit 2b. The first conversion unit 2b is electrically connected to the conductive layer CL and is located in the area overlapping the conductive layer CL, and is configured to convert incident X-rays into an image data signal S2.

[0033] The TFT 13a can switch between charging and discharging the TFD 15a. If the self-capacitance of the TFD 15a is insufficient, the conversion substrate 2 may further include a capacitor (storage capacitor) to store the charge converted by the TFD 15a.

[0034] The scintillator layer 5 is made of thallium-activated cesium iodide (CsI:Tl). When the scintillator layer 5 is formed using a vacuum deposition method, the scintillator layer 5 is made up of an aggregate of a plurality of columnar crystals. The thickness of the scintillator layer 5 is, for example, 600 μm. At the outermost surface of the scintillator layer 5, the thickness of the columnar crystals of the scintillator layer 5 is 8 to 12 μm.

[0035] The material forming the scintillator layer 5 is not limited to CsI:Tl, but may also be formed of thallium-activated sodium iodide (NaI:Tl), sodium-activated cesium iodide (CsI:Na), europium-activated cesium bromide (CsBr:Eu), sodium iodide (NaI), or the like.

[0036] When forming the scintillator layer 5 using the vacuum deposition method, a mask with an opening is used. In this case, the scintillator layer 5 is formed in the area facing the opening on the conversion substrate 2. The scintillator material deposited by deposition is also deposited on the surface of the mask. The scintillator material is also deposited near the opening of the mask, and crystals grow so as to gradually extend into the opening. When the crystals extend from the mask into the opening, deposition of the scintillator material onto the conversion substrate 2 near the opening is suppressed. Therefore, as shown in Figure 2, the thickness of the scintillator layer 5 near the periphery gradually decreases outward.

[0037] Alternatively, the scintillator layer 5 may have a plurality of scintillator sections arranged in a matrix, each having a rectangular prism shape, and provided in a one-to-one correspondence with the first conversion section 2b. When forming such a scintillator layer 5, a scintillator material obtained by mixing gadolinium oxysulfide (Gd2O2S) phosphor particles with a binder is applied to the conversion substrate 2, and the scintillator material is baked and hardened. Then, dicing is performed using a dicer or the like to form lattice-shaped grooves in the scintillator material. In the above case, air or an inert gas such as nitrogen (N2) for preventing oxidation is sealed between the plurality of scintillator sections. Alternatively, the space between the plurality of scintillator sections may be set to a space reduced in pressure below atmospheric pressure.

[0038] In this embodiment, the X-ray detection panel PNL further includes a light-reflecting layer 6. The light-reflecting layer 6 is provided on the scintillator layer 5. In other words, the light-reflecting layer 6 is provided on the X-ray incident side of the scintillator layer 5. The light-reflecting layer 6 is located at least in the effective area DA and covers the upper surface of the scintillator layer 5. The light-reflecting layer 6 is provided to improve the light (fluorescence) utilization efficiency and sensitivity characteristics. That is, the light-reflecting layer 6 reflects light generated in the scintillator layer 5 and traveling in the direction opposite to the side where the TFD 15a is provided, so that the light is directed toward the TFD 15a. However, the light-reflecting layer 6 is not necessarily required and may be provided depending on the sensitivity characteristics required of the X-ray detection panel PNL.

[0039] For example, a coating material made of a mixture of light-scattering particles such as titanium oxide (TiO2), a resin, and a solvent can be applied onto the scintillator layer 5, and then the coating material can be dried to form the light-reflecting layer 6.

[0040] The structure of the light-reflecting layer 6 and the manufacturing method of the light-reflecting layer 6 are not limited to the above examples and can be modified in various ways. For example, the light-reflecting layer 6 may be formed by depositing a layer made of a metal with high light reflectivity, such as a silver alloy or aluminum, on the scintillator layer 5. Alternatively, the light-reflecting layer 6 may be formed by providing, on the scintillator layer 5, a sheet whose surface includes a metal layer with high light reflectivity, such as a silver alloy or aluminum, or a resin sheet containing light-scattering particles.

[0041] A moisture-proof cover (moisture-proof sheet) 7 covers the scintillator layer 5 and the light-reflecting layer 6. The moisture-proof cover 7 is provided to prevent the properties of the light-reflecting layer 6 and the scintillator layer 5 from deteriorating due to moisture contained in the air. The moisture-proof cover 7 completely covers the exposed portion of the scintillator layer 5. A gap may be provided between the moisture-proof cover 7 and the light-reflecting layer 6, etc., or the moisture-proof cover 7 may be in contact with the light-reflecting layer 6, etc.

[0042] The moisture-proof cover 7 is formed of a sheet containing a metal. Examples of the metal include aluminum-containing metals, copper-containing metals, magnesium-containing metals, tungsten-containing metals, stainless steel, and Kovar. When the moisture-proof cover 7 contains a metal, the moisture-proof cover 7 can prevent or significantly reduce moisture penetration.

[0043] FIG. 4 is a circuit diagram showing the X-ray detection panel PNL, the circuit board 11, and the plurality of FPCs 2e1 and 2e2 according to this comparative example. 2 to 4, the circuit board 11 is provided with a drive circuit 11a and a detection circuit 11b. These circuits can be provided on a single board, or can be provided separately on multiple boards. The other ends of the multiple wires provided on the FPC 2e1 are electrically connected to the drive circuit 11a, respectively. The other ends of the multiple wires provided on the FPC 2e2 are electrically connected to the detection circuit 11b, respectively.

[0044] The driving circuit 11a is electrically connected to the plurality of scanning lines G via the FPC 2e1 and can sequentially scan the plurality of scanning lines G. The driving circuit 11a switches the TFT 13a between an on state and an off state. The driving circuit 11a has a plurality of gate drivers 11aa and a row selection circuit 11ab. A scanning signal S1 is input to the row selection circuit 11ab from an image processing unit (not shown) or the like provided outside the X-ray detector 1. The row selection circuit 11ab inputs the scanning signal S1 to the corresponding gate driver 11aa in accordance with the scanning direction of the X-ray image. The gate driver 11aa inputs the scanning signal S1 to the corresponding scanning line G.

[0045] For example, the driver circuit 11a sequentially inputs a scanning signal S1 to multiple scanning lines G via the FPC 2e1. The scanning signal S1 input to the scanning line G turns the TFT 13a on or off, and when the TFT 13a is turned on, the charge (image data signal S2) from the TFD 15a is output to the FPC 2e2.

[0046] The detection circuit 11b is electrically connected to the multiple signal lines U and can detect image data signals S2 generated by the multiple pixels PX. The detection circuit 11b includes multiple integrating amplifiers 11ba, multiple selection circuits 11bb, and multiple AD converters (Analog-to-Digital Converters) 11bc. Each integrating amplifier 11ba is electrically connected to one signal line U. The integrating amplifier 11ba sequentially receives image data signals S2 from the multiple first conversion units 2b. The integrating amplifier 11ba then integrates the current flowing within a certain time period and outputs a voltage corresponding to the integrated value to the selection circuit 11bb. In this manner, the value (amount of charge) of the current flowing through the signal line U within a predetermined time period can be converted into a voltage value. That is, the integrating amplifier 11ba converts image data information corresponding to the intensity distribution of fluorescence generated in the scintillator layer 5 into potential information.

[0047] The selection circuit 11bb selects the integrating amplifier 11ba to perform reading and sequentially reads out the image data signals S2 converted into potential information. The AD converter 11bc sequentially converts the read-out image data signals S2 into digital signals. The digital image data signals S2 are input to the image processing unit via wiring. The digital image data signals S2 may be transmitted to the image processing unit wirelessly. The image processing unit creates an X-ray image based on the digital image data signals S2. The image processing unit may also be integrated with the circuit board 11.

[0048] Fig. 5 is a plan view showing the X-ray detection panel PNL according to this comparative example, and is a diagram for explaining the positional relationship between the effective area DA and the scintillator layer 5. In Fig. 5, the scintillator layer 5 is marked with diagonal lines slanting upward to the right, and the bonded body 8 is marked with diagonal lines slanting downward to the right. Fig. 6 is a cross-sectional view showing the X-ray detection panel PNL of Fig. 5 taken along line VI-VI, and also showing the FPC 2e1.

[0049] 5 and 6, the conversion substrate 2 has an effective area DA, a frame-shaped first non-detection area NDA1 located around the effective area DA, and a second non-detection area NDA2 outside the first non-detection area NDA1. In this comparative example, the second non-detection area NDA2 has a frame-like shape. The base material 2a is located in the effective area DA and the frame-shaped non-detection areas (the first non-detection area NDA1 and the second non-detection area NDA2) surrounding the effective area DA.

[0050] The scintillator layer 5 is located at least in the effective area DA. The conversion substrate 2 further has a plurality of pads 2d1 and a plurality of pads 2d2. The pads 2d1 and 2d2 are located in a second non-detection area NDA2. In this embodiment, the plurality of pads 2d1 are arranged along the left side of the base material 2a, and the plurality of pads 2d2 are arranged along the bottom side of the base material 2a. Note that FIG. 5 shows the plurality of pads schematically, and the number, shape, size, position, and pitch of the plurality of pads are not limited to the example shown in FIG. 5.

[0051] One scanning line G extends through the effective area DA, the first non-detection area NDA1, and the second non-detection area NDA2, and is electrically connected to one of the pads 2d1. One signal line U extends through the effective area DA, the first non-detection area NDA1, and the second non-detection area NDA2, and is electrically connected to one of the pads 2d2. One pad 2d1 is electrically connected to one of the multiple wirings provided on the FPC 2e1, and one pad 2d2 is electrically connected to one of the multiple wirings provided on the FPC 2e2 (Figure 2).

[0052] The X-ray detection panel PNL further includes an assembly 8. The assembly 8 is provided around the scintillator layer 5. The assembly 8 has a frame-like shape and extends continuously around the scintillator layer 5. The assembly 8 is joined to the conversion substrate 2 (for example, the insulating layer 25).

[0053] The moisture-proof cover 7 is provided on the conversion substrate 2, the scintillator layer 5, and the light-reflecting layer 6. The moisture-proof cover 7 is located in the effective area DA and the non-detection area (first non-detection area NDA1). In the plan view shown in FIG. 5, the moisture-proof cover 7 completely covers the scintillator layer 5. As shown in FIG. 6, the portion of the scintillator layer 5 that is not covered by the conversion substrate 2 and the assembly 8 is completely covered by the moisture-proof cover 7. The moisture-proof cover 7 is bonded to the assembly 8. The moisture-proof cover 7, together with the conversion substrate 2 and the assembly 8, seals the scintillator layer 5 and the light-reflecting layer 6.

[0054] For example, if the moisture-proof cover 7 and the assembly 8 are joined in an environment where the pressure is reduced below atmospheric pressure, the moisture-proof cover 7 can be brought into contact with the light-reflecting layer 6 and the like. Generally, the scintillator layer 5 has voids that account for approximately 10 to 40% of its volume. If the moisture-proof cover 7 and the assembly 8 are joined in an environment where the pressure is reduced below atmospheric pressure, damage to the moisture-proof cover 7 can be suppressed even when the X-ray detector 1 is transported by aircraft or the like. For the above reasons, it is preferable that the pressure in the space defined by the assembly 8 and the moisture-proof cover 7 be lower than atmospheric pressure.

[0055] The bonding body 8 is located between the conversion substrate 2 and the moisture-proof cover 7 and is made of a thermoplastic resin. By heating the vicinity of the periphery of the moisture-proof cover 7, the bonding body 8 bonds the conversion substrate 2 and the moisture-proof cover 7 together. The FPC 2e1 is fixed to the conversion substrate 2 (X-ray detection panel PNL) by a connecting material AD, and is electrically connected to the pad 2d1. The connecting material AD is made of ACF.

[0056] FIG. 7 is a circuit diagram showing the conversion board 2, the circuit board 11, the image transmission unit 4, and the power supply PS1 of the X-ray detector 1 according to this comparative example. As shown in Fig. 7, pixels PX are arranged in a matrix with a row direction X and a column direction Y. For example, pixels PX are effective for detecting radiation (e.g., X-rays). Each pixel PX includes a TFT 13a, a TFD 15a, etc. In the figure, the capacitor represents the self-capacitance of the TFD 15a.

[0057] Each TFD 15a is electrically connected to one corresponding TFT 13a among the plurality of TFTs 13a. Each TFD 15a is also electrically connected to a bias line BL1. More specifically, the anode of each TFD 15a is electrically connected to the bias line BL1, and the cathode is electrically connected to one corresponding TFT 13a. The bias line BL1 is electrically connected to a power supply PS1. The X-ray detector 1 according to this comparative example is configured as described above.

[0058] When X-ray imaging is performed using the X-ray detector 1, the X-ray detector 1 can wirelessly communicate with an X-ray tube device (not shown), and can, for example, synchronize with the X-ray tube device. The X-ray tube device irradiates a subject with X-rays, and the X-ray detector 1 can detect the X-rays that have passed through the subject. This allows the X-ray detector 1 to perform X-ray imaging of the subject together with the X-ray tube device.

[0059] In the X-ray detector 1 and X-ray imaging method according to the comparative example configured as described above, when continuous pulsed X-rays are irradiated onto the X-ray detector 1 while the TFT 13a of the pixel PX is on, noise occurs, reducing the S / N ratio of the image data signal S2. Therefore, it is important to synchronize the irradiation of continuous pulsed X-rays by the X-ray tube device with the drive of the X-ray detector 1 during X-ray imaging. To synchronize the drive of the X-ray detector 1 with the drive of the X-ray tube device, the X-ray detector 1 can communicate wirelessly with the X-ray tube device. The portable X-ray detector 1 is equipped with wireless communication means.

[0060] However, when the X-ray detector 1 synchronizes with the X-ray tube device using wireless communication means, wireless communication may be temporarily interrupted depending on the position of the subject (e.g., patient) and the position of the wireless communication means of the X-ray detector 1. If a problem specific to wireless types occurs and the X-ray detector 1 cannot synchronize with the X-ray tube device, X-ray imaging cannot be performed at the desired timing, and an image data signal S2 with a low S / N ratio will be derived. In particular, if the X-ray detector 1 cannot synchronize with the X-ray tube device at the start of X-ray irradiation, there is a risk that unnecessary X-rays will be irradiated onto the subject.

[0061] Therefore, with the X-ray detector 1 according to this comparative example, it is difficult to obtain an X-ray detector 1 that can continuously detect X-rays in synchronization with the X-rays based on the incident X-rays. Therefore, it is desirable that the X-ray detector 1 can reliably synchronize with the X-ray tube assembly without being provided with a communication means (wired communication means or wireless communication means).

[0062] (First embodiment) Next, the configuration of the X-ray detector 1 and the X-ray imaging method according to the first embodiment will be described. The X-ray detector 1 has the same configuration as the above-mentioned comparative example except for the configuration described in this first embodiment. Fig. 8 is a circuit diagram showing the conversion board 2, circuit board 11, image transmission unit 4, power supply PS1, and monitor unit M of the X-ray detector 1 according to this first embodiment.

[0063] As shown in Figure 8, the driver circuit 11a sequentially scans the scan lines G1, G2, and G3, turning on the TFT 13a and accumulating a certain amount of charge from the bias line BL1 in the TFD 15a. When the TFT 13a is subsequently turned off and X-rays are irradiated, the X-rays are converted into visible light in the scintillator layer 5, and the visible light irradiates the TFD 15a. Charges (electrons and holes) generated by the photoelectric effect combine with the different charges stored in the TFD 15a, reducing the accumulated charge in the TFD 15a.

[0064] When the driver circuit 11a scans the multiple scan lines G again, the TFT 13a switches on and the TFD 15a is recharged through the bias line BL1. The reduced image data signal S2 passes through the signal line U as a recharge current from the TFD 15a and is detected as an electrical signal by the detection circuit 11b.

[0065] By scanning the scanning line G1, an image data signal S2 corresponding to the scanning line G1 is transferred from the detection circuit 11b to the image transmission unit 4. Thereafter, by scanning the scanning line G2, an image data signal S2 corresponding to the scanning line G2 is transferred from the detection circuit 11b to the image transmission unit 4. The above-mentioned integrating amplifier 11ba can integrate the current that flows while the TFT 13a is in the ON state.

[0066] The monitor unit M is electrically connected between a power supply PS1 serving as a first power supply and a bias line BL1. The monitor unit M can monitor the current flowing from the power supply PS1 to the bias line BL1. Data monitored by the monitor unit M is sent to the image transmission unit 4. In the first embodiment, the image transmission unit 4 functions as a control circuit. However, a circuit unit other than the image transmission unit 4, such as the circuit board 11, may function as the control circuit.

[0067] The power supply PS1, monitor unit M, bias line BL1, and multiple first conversion units 2b function as a detection unit DT. The detection unit DT can generate an X-ray detection signal (radiation detection signal) by detecting X-rays (radiation). The X-ray detection signal is generated by a current flowing from the power supply PS1 to the bias line BL1. The image transmission unit (control circuit) 4 can control the output of the image data signal S2 based on the X-ray detection signal (radiation detection signal). For example, the image transmission unit 4 can permit or prohibit the output of the image data signal S2 to an external image display device.

[0068] The image transmission unit 4, to which the image data signal S2 is input from the detection circuit 11b, is equipped with a memory for performing image processing. Every time the drive circuit 11a scans the scanning line G, digital data based on the image data signal S2 is sequentially transferred to the memory in the image transmission unit 4 and recorded.

[0069] As shown in Fig. 9, the X-ray detector 1 may further include a power supply PS2 as a second power supply and a second conversion unit 2c. The second conversion unit 2c is electrically connected to the power supply PS2, is independent of the pixels PX, and converts incident X-rays (radiation) into electric charges. In the example shown in Fig. 9, the second conversion unit 2c includes a TFD 15b and a portion of the scintillator layer 5 that overlaps with the TFD 15b. The TFD 15b is located in the non-detection area NDA, for example, but may also be located in the active area DA.

[0070] The power supply PS2 is electrically connected to the TFD 15b so as to apply a reverse bias to the TFD 15b. The monitor M is electrically connected between the power supply PS2 and the second conversion unit 2c (TFD 15b). The monitor M can monitor the current flowing from the power supply PS2 to the second conversion unit 2c (TFD 15b). In the example shown in FIG. 9, the power supply PS2, the monitor M, and the second conversion unit 2c function as a detection unit DT. An X-ray detection signal is generated by the current flowing from the power supply PS2 to the second conversion unit 2c. If the monitor M detects a current value equal to or greater than a threshold, the image transmission unit (control circuit) 4 can determine that X-rays have been irradiated.

[0071] Alternatively, one or more of the first conversion units 2b of the pixels PX may also function as the detection unit DT. In this case, the X-ray detection signal (radiation detection signal) is generated by a current flowing from the TFD 15a to the conductive layer (pixel electrode) CL connected to the TFD 15a. The pixel PX (first conversion unit 2b) that extracts the X-ray detection signal may be fixed at all times. In this case, it is possible to suppress deviation of the time axis when extracting the X-ray detection signal. However, since one horizontal scanning period (one frame period) is a short time period, the pixel PX (first conversion unit 2b) that extracts the X-ray detection signal may be changed as needed. This is because there is no significant adverse effect on the derivation of the period (detection period) for detecting the X-ray detection signal. Alternatively, the X-ray detection signal (radiation detection signal) is generated by a current flowing through one or more of the multiple signal lines U.

[0072] Next, an X-ray imaging method applied to the X-ray detector 1 of the first embodiment will be described. Fig. 10 is a flowchart for explaining the X-ray imaging method applied to the X-ray detector 1 of the first embodiment. Fig. 11 is a flowchart following Fig. 10 for explaining the X-ray imaging method.

[0073] 10 and 11, when X-ray imaging is started by scanning multiple scanning lines G in a first scanning cycle T1 including a first accumulation period Ta and a first readout period Tg, first, in step ST1, the image transmission unit (control circuit) 4 determines whether an X-ray detection signal has been acquired from the detection unit DT. During the first readout period Tg, a first scanning signal S1a is sequentially applied to the multiple scanning lines G, and the TFT 13a is switched on. During the first readout period Tg, the image data signal S2 is sequentially transferred to the detection circuit 11b. The first accumulation period Ta is a period during which the drive circuit 11a does not output the first scanning signal S1a to the multiple scanning lines G.

[0074] If the image transmission unit 4 does not acquire an X-ray detection signal (step ST1), the process proceeds to step ST8, where the image transmission unit 4 prohibits the output of the image data signal S2 to the outside of the X-ray detector 1, and then in step ST9, the image transmission unit 4 maintains the first accumulation period Ta, the first readout period Tg, and the first scanning signal S1a.

[0075] Thereafter, in step ST7, it is determined whether or not to end the X-ray imaging. If the X-ray imaging is to be continued without ending it, the process proceeds to step ST1, and the operations from step ST1 onwards are repeatedly executed.

[0076] On the other hand, if the image transmission unit 4 acquires an X-ray detection signal (step ST1), it proceeds to step ST2, where the image transmission unit 4 determines whether the detection period T2, which is the period for detecting the X-ray detection signal, is within the first scanning period T1.

[0077] If the detection period T2 is not within the first scanning period T1 (step ST2), the process proceeds to step ST10, where the image transmission unit 4 allows the output of the image data signal S2 to the outside of the X-ray detector 1, and then in step ST11, the image transmission unit 4 determines that a single X-ray has been detected, and then the process proceeds to step ST9.

[0078] On the other hand, if the detection period T2 is within the first scanning period T1 (step ST2), the process proceeds to step ST3, where the image transmission unit 4 determines whether the detection period T2 is the same as the first scanning period T1.

[0079] If the detection period T2 is the same as the first scanning period T1 (step ST3), the process proceeds to step ST12, where the image transmission unit 4 permits the output of an image data signal S2 to the outside of the X-ray detector 1. Subsequently, in step ST13, the image transmission unit 4 determines that continuous X-rays have been detected, and then the process proceeds to step ST9. The signal output to the outside of the X-ray detector 1 is, for example, an Ethernet signal (command notification signal) or a hardware signal.

[0080] On the other hand, if the detection period T2 is not the same as the first scanning period T1 (step ST3), the process proceeds to step ST4, where the image transmission unit 4 permits the output of the image data signal S2 to the outside of the X-ray detector 1, and then in step ST5, the image transmission unit 4 can determine that continuous pulsed X-rays have been detected. In other words, when the X-ray detection signal is detected twice within the time period of the detection period T2, the image transmission unit 4 can determine that continuous pulsed X-rays have been detected.

[0081] Subsequently, in step ST6, the image transmission unit 4 adjusts the first accumulation period Ta to the second accumulation period Tb, adjusts the first readout period Tg to the second readout period Tg', adjusts the first scanning signal S1a to the second scanning signal S1b, and adjusts the timing of the second accumulation period Tb. If X-ray imaging is to be continued without terminating it (step ST7), the image transmission unit 4 proceeds to step ST1 and repeatedly executes the operations from step ST1 onward. In this way, the first scanning period T1 can be adjusted to the second scanning period (T2). However, even if adjustment is performed in a single step ST6, the image transmission unit 4 may not be able to accurately adjust the second scanning period (T2). Therefore, it is preferable to execute step ST6, then proceed to step ST1, and repeatedly execute the operations from step ST1 onward. This allows the second scanning period (T2) to be accurately adjusted (corrected), contributing to the derivation of an image data signal S2 with a high S / N ratio.

[0082] As a result, the image transmission unit (control circuit) 4 controls the driving of the drive circuit 11a based on the X-ray detection signal, and can automatically adjust the first scanning period T1. The X-ray detector 1 can be used not only for capturing still images but also for capturing moving images.

[0083] FIG. 12 is a timing chart showing the scanning signal S1 (first scanning signal S1a and second scanning signal S1b), the X-ray detection signal DS, the image data signal S2 to be output to the outside, the continuous pulse X-rays XR1, and the current I1 generated based on the continuous pulse X-rays before and after detecting the continuous pulse X-rays in the first embodiment.

[0084] As shown in FIG. 12, in a first scanning period T1, a first scanning signal S1a is sequentially applied to multiple scanning lines G. The time period of the first scanning period T1 includes a first accumulation period Ta and a first readout period Tg. The first accumulation period Ta is a period during which the driving circuit 11a does not scan the multiple scanning lines G, and charges can be accumulated in the multiple pixels PX. The first readout period Tg is a period different from the first accumulation period Ta, and is a period during which the driving circuit 11a sequentially scans the multiple scanning lines G, and the detection circuit 11b detects the image data signal S2.

[0085] Furthermore, in a second scanning period (T2), the second scanning signal S1b is sequentially applied to the plurality of scanning lines G. The time period of the second scanning period (T2) includes a second accumulation period Tb and a second readout period Tg'. The second accumulation period Tb is a period during which the driving circuit 11a does not scan the plurality of scanning lines G, and charges can be accumulated in the plurality of pixels PX. The second readout period Tg' is a period different from the second accumulation period Tb, and is a period during which the driving circuit 11a sequentially scans the plurality of scanning lines G, and the detection circuit 11b detects the image data signal S2.

[0086] Before the above detection unit DT detects X-rays, the X-ray detector 1 alternately repeats the operations of the first accumulation period Ta and the first readout period Tg. The time period of the first scanning cycle T1 is the sum of the first accumulation period Ta and the first readout period Tg. The time period of the first scanning cycle T1 is preset. The drive circuit 11a scans a plurality of scanning lines G at the first scanning cycle T1 under the control of the image transmission unit (control circuit) 4.

[0087] Here, as an example, the waveform of the current I1 detected by the flow monitor unit M in the TFD15b is also shown in FIG. 12. It is assumed that the current I1 is intermittently detected at a detection period that is synchronized with the continuous pulsed X-rays XR1 and is the period for detecting the X-ray detection signal DS.

[0088] FIG. 12 shows an example in which X-rays are detected four times, and the continuous pulsed X-rays XR1 are irradiated in a pulsed manner at the above detection period. When the image transmission unit 4 determines that the detection period is shorter than the first scanning cycle T1, the drive circuit 11a scans a plurality of scanning lines G at a second scanning cycle (T2) shorter than the first scanning cycle T1 (T2 < T1) under the control of the image transmission unit 4. The time period of the second scanning cycle is the same as the time period of the above detection cycle T2. The timing of the second scanning cycle also coincides with the timing of the detection cycle T2.

[0089] Until X-rays are detected twice, the drive circuit 11a scans a plurality of scanning lines G at the first scanning cycle T1, and the image transmission unit 4 prohibits the output of the image data signal S2 from the X-ray detector 1 to the outside. After X-rays are detected twice within the time period of the first scanning period T1, the scanning period can be changed to a second scanning period (T2). The driving circuit 11a scans multiple scanning lines G in the second scanning period (T2), and the image transmission unit 4 permits the output of an image data signal S2 to the outside of the X-ray detector 1. When the image transmission unit 4 determines that the detection period is shorter than the first scanning period T1, it can permit the output of the image data signal S2. The time period of the second scanning period (T2) is the sum of the second accumulation period Tb and the second readout period Tg'. Note that the time period of the second readout period Tg' may be the same as the time period of the first readout period Tg, but is not particularly limited to this.

[0090] The X-ray detector 1 can adjust the timing of the second accumulation period Tb. The drive circuit 11a, under the control of the image transmission unit 4, can adjust the timing of the second accumulation period Tb by adjusting the first accumulation period Ta to the second accumulation period Tb and adjusting the first readout period Tg to the second readout period Tg'. The drive circuit 11a overlaps the second accumulation period Tb with the period for detecting the X-ray detection signal DS, but does not overlap the second readout period Tg'.

[0091] In other words, the timing of the second accumulation period Tb is adjusted so that the last pulse of the second scanning signal S1b in the second readout period Tg' falls a period T3 earlier than the rise of the X-ray detection signal DS. That is, the timing of the second accumulation period Tb is adjusted so that the first pulse of the second scanning signal S1b in the second readout period Tg' rises a period T4 later than the rise of the X-ray detection signal DS. The time duration of the period T4 is the time duration of the second scanning cycle (T2) minus the time duration of the second readout period Tg' and the time duration of the period T3 (T4 = T2 - Tg' - T3).

[0092] By adjusting the timing of the second accumulation period Tb in this way, it is possible to avoid a situation in which the continuous pulsed X-rays XR1 are irradiated during the second readout period Tg', thereby suppressing the generation of noise and contributing to the derivation of an image data signal S2 with a high S / N ratio. It is desirable to determine the period T3 for each X-ray detector 1 in consideration of the time period required to detect the continuous pulsed X-rays XR1.

[0093] Since the first accumulation period Ta of the first signal 1 of the X-ray detection signal DS is indefinite, the output to the outside of the image data signal S2 generated at the same time as signal 1 remains prohibited. The second signal 2 of the image data signal S2 was generated at the same time as the second signal 2 of the X-ray detection signal DS. The second signal 2 of the image data signal S2 is permitted to be output to the outside. The image data signal S2 can be recorded in a storage unit SM, which will be described later, or the like.

[0094] FIG. 13 is a timing chart showing the scanning signal S1 (first scanning signal S1a), the X-ray detection signal DS, the image data signal S2 to be output to the outside, the single X-ray XR2, and the current I2 generated based on the single X-ray before and after detecting the single X-ray in the first embodiment.

[0095] 13, the driving circuit 11a scans a plurality of scanning lines G in a first scanning period T1 under the control of the image transmission unit 4. When the image transmission unit 4 determines that the detection period T2 for detecting the X-ray detection signal DS is longer than the first scanning period T1 (T2>T1), the driving circuit 11a continues scanning the plurality of scanning lines G in the first scanning period T1 under the control of the image transmission unit 4.

[0096] The X-ray detector 1 further includes a memory unit SM that stores the image data signal S2. For example, the memory unit SM is provided in the image transmission unit 4 (FIGS. 8 and 9). Here, a period longer than the first scanning period T1 is defined as a reference period (maximum period) T1+Tout. The time period of the reference period is the sum of the time period of the first scanning period T1 and the time period Tout. When the image transmission unit 4 determines that the detection period for detecting the X-ray detection signal DS is longer than the reference period, it can permit the image data signal S2 recorded in the memory unit SM to be output to the outside. As described above, it is possible to determine the time to start outputting the image data signal S2 to the outside.

[0097] In other words, in X-ray imaging using a single X-ray XR2, the period of the single X-ray XR2 cannot be detected, and therefore the reference period is preset in the X-ray detector 1. If the second signal 2 of the X-ray detection signal DS cannot be detected even after the reference period has elapsed since the first signal 1 of the X-ray detection signal DS was detected, the digital data of the first signal 1 of the image data signal S2 is recorded in the memory unit SM so that the first signal 1 of the image data signal S2 can be read out later.

[0098] FIG. 14 is a timing chart showing the scanning signal S1 (first scanning signal S1a), the X-ray detection signal DS, the image data signal S2 to be output to the outside, the continuous X-rays XR3, and the current I3 generated based on the continuous X-rays XR3 before and after detecting the continuous X-rays XR3 in the first embodiment.

[0099] 14, the driving circuit 11a scans a plurality of scanning lines G in a first scanning cycle T1 under the control of the image transmission unit 4. When the image transmission unit 4 determines that the time period of the detection cycle T2 for detecting the X-ray detection signal DS is the same as the time period of the first scanning cycle T1 (T2=T1), the driving circuit 11a continues scanning the plurality of scanning lines G in the first scanning cycle T1 under the control of the image transmission unit 4.

[0100] The detection unit DT can reset the X-ray detection signal DS every time the time period of the first scanning period T1 elapses under the control of the image transmission unit 4. This allows the detection period T2 to coincide with the first scanning period T1. When the image transmission unit 4 determines that the time period of the detection period T2 is the same as the time period of the first scanning period T1, it can permit the image data signal S2 stored in the memory unit SM to be output to the outside. As described above, it is possible to determine the time to start outputting the image data signal S2 to the outside.

[0101] In other words, before the detection unit DT detects X-rays, the X-ray detector 1 alternately and repeatedly performs an operation during a first accumulation period Ta and an operation during a first readout period Tg. The time period of the first scanning period T1 is the sum of the first accumulation period Ta and the first readout period Tg. The time period of the first scanning period T1 is set in advance.

[0102] When the image transmission unit 4 determines that the time period of the detection cycle T2 is the same as the time period of the first scanning cycle T1, it can permit the output of the image data signal S2 to the outside, but the scanning cycle may remain the first scanning cycle T1 and may not be changed. By configuring the X-ray detector 1 so that the X-ray detection signal DS is repeatedly reset at the first scanning cycle T1, in the case of continuous X-rays XR3, the detection cycle T2 of the X-ray detection signal DS coincides with the first scanning cycle T1.

[0103] In the case of continuous X-rays XR3, there is no fluctuation in X-ray irradiation (rising and falling edges) after the timing when X-ray irradiation starts during the first readout period Tg, so noise is not a problem. An image data signal S2 with a large S / N ratio can be derived.

[0104] Furthermore, in the case of continuous X-rays XR3, the accumulation period (first accumulation period Ta) is never indefinite, so external output is permitted from the first signal 1 of the image data signal S2, which is generated at the same time as the first signal 1 of the X-ray detection signal DS. However, continuous X-rays XR3 generally require time to rise, and the voltage level of the first signal 1 of the image data signal S2 tends to be low. Therefore, if the rise of the continuous X-rays XR3 is slow, it is preferable for the image transmission unit 4 to externally output the second signal 2 of the image data signal S2. Note that the rise time of the continuous X-rays XR3 varies depending on the configuration of the X-ray imaging system including the X-ray detector 1. Therefore, the image transmission unit 4 may start external output from the first signal 1 of the image data signal S2 or from the second signal 2 of the image data signal S2.

[0105] According to the X-ray detector 1 and X-ray imaging method of the first embodiment configured as described above, the X-ray detector 1 can detect X-rays emitted from the X-ray tube assembly, thereby determining whether continuous pulsed X-rays XR1, single X-rays XR2, or continuous X-rays XR3 are being irradiated, and can synchronize with the X-ray tube assembly. Even if the portable X-ray detector 1 cannot receive a synchronization signal from the X-ray tube assembly and fails to communicate wirelessly with the X-ray tube assembly, it can still synchronize with the X-ray tube assembly.

[0106] For example, when the X-ray detector 1 is irradiated with continuous pulsed X-rays XR1, the timing of the second accumulation period Tb can be adjusted so that the continuous pulsed X-rays XR1 is irradiated during the second accumulation period Tb, rather than during the second readout period Tg'. Since the generation of noise can be suppressed, the S / N ratio of the image data signal S2 can be increased.

[0107] Furthermore, even if wireless communication with the X-ray tube device fails, the X-ray detector 1 can quickly establish synchronization with the X-ray tube device, thereby preventing unnecessary X-ray irradiation of the subject. From the above, the first embodiment can provide an X-ray detector 1 that can continuously detect X-rays in synchronization with the X-rays based on the incident X-rays (radiation). Furthermore, the X-ray detector 1 can reliably synchronize with the X-ray tube assembly without being provided with a communication means.

[0108] (Second embodiment) Next, the configuration of an X-ray detector 1 according to a second embodiment will be described. The X-ray detector 1 has the same configuration as the X-ray detector 1 of the first embodiment, except for the configuration described in this second embodiment. The X-ray detector 1 of this second embodiment is a direct conversion type X-ray detector. FIG. 15 is a circuit diagram showing a part of the conversion board 2 of the X-ray detector 1 according to this second embodiment. FIG. 16 is an enlarged cross-sectional view showing a part of the conversion board 2 shown in FIG. 15.

[0109] 15 and 16, the X-ray detector 1 does not include a TFD 15a, a scintillator layer 5, a light-reflecting layer 6, a moisture-proof cover 7, or a power supply PS1. The X-ray detector 1 further includes an X-ray conversion layer 30, which is a radiation conversion layer, a plurality of capacitors 17, a bias electrode layer 31, a power supply PS3 as a first power supply, an insulating layer 32, and an X-ray grid 33 as an X-ray shielding portion (radiation shielding portion).

[0110] Each capacitor 17 is electrically connected to one of a plurality of conductive layers (pixel electrodes) CL. The capacitor 17 has a lower electrode 17a formed on the substrate 2a and an upper electrode 17b formed on the insulating layer 14 and facing the lower electrode 17a. The upper electrode 17b is electrically connected to the drain electrode DE of the TFT 13a. An insulating layer 18 is formed on the insulating layer 14, the semiconductor layer SC, the upper electrode 17b, the source electrode SE, and the drain electrode DE. The conductive layer CL is formed on the insulating layer 18 and is electrically connected to the drain electrode DE.

[0111] The X-ray conversion layer 30 is in contact with the plurality of conductive layers CL of the plurality of pixels PX, and can convert incident X-rays into charges of a magnitude corresponding to the intensity of the X-rays. The bias electrode layer 31 is formed on the X-ray conversion layer 30. The bias electrode layer 31 is formed so as to be able to apply a predetermined bias voltage to the X-ray conversion layer 30. The insulating layer 32 is formed on the bias electrode layer 31. The X-ray grid 33 is formed on the insulating layer 32. The X-ray grid 33 overlaps between the conductive layers CL. Therefore, the X-ray grid 33 has the function of blocking X-rays incident from pixels PX adjacent to the pixel PX, and can suppress a decrease in resolution characteristics due to scattered X-rays.

[0112] The power supply PS3 is electrically connected to the bias electrode layer 31. In each pixel PX, the first conversion unit 2b is a portion of the X-ray conversion layer 30 that overlaps the conductive layer CL. In the second embodiment, one or more of the multiple first conversion units 2b of the multiple pixels PX, the bias electrode layer 31, and the power supply PS3 also serve as the detection unit DT. The X-ray detection signal (radiation detection signal) DS is generated by a current flowing from the above-mentioned portion of the X-ray conversion layer 30 to the conductive layer CL in contact with the above-mentioned portion. The direct conversion type X-ray detector 1 is configured as described above.

[0113] Next, an X-ray imaging method using the X-ray detector 1 will be described. X-rays are irradiated onto the X-ray conversion layer 30 of the conversion substrate 2 by passing through the subject, etc. As a result, charges 30a generated in the X-ray conversion layer 30 move to the conductive layer CL of any pixel PX due to an electric field oriented by the bias voltage applied to the bias electrode layer 31. Note that in FIG. 16, electrons e are shown moving toward the bias electrode layer 31, and holes h are shown moving toward the conductive layer CL. The charges 30a that have moved to the conductive layer CL are stored in the capacitor 17 via the drain electrode DE of the TFT 13a.

[0114] Next, a scanning signal S1 for switching the TFT 13a between an ON state and an OFF state is input to the scanning line G. Here, the scanning signal S1 is output from the above-mentioned drive circuit 11a. When the TFT 13a is turned ON by the scanning signal S1, the charge accumulated in the capacitor 17 of each pixel PX is output to the signal line U as an image data signal S2, which is a charge signal, and the image data signal S2 is transmitted to the detection circuit 11b and the image transmission unit 4.

[0115] In the X-ray detector 1 and X-ray imaging method according to the second embodiment configured as described above, the X-ray detector 1 can obtain the same effects as those of the first embodiment. That is, even if the X-ray detector 1 is a direct conversion type, it can determine which of the continuous pulse X-rays XR1, the single X-rays XR2, and the continuous X-rays XR3 is being irradiated, and can reliably synchronize with the X-ray tube device.

[0116] In the direct conversion type X-ray detector 1, the detection unit DT can be modified in various ways. 17, the X-ray detector 1 of the modified example further includes a monitor unit M. The monitor unit M is electrically connected between the power supply PS3 and the bias electrode layer 31. The monitor unit M can monitor the current flowing from the power supply PS3 to the bias electrode layer 31. Data monitored by the monitor unit M is sent to an image transmission unit (control circuit) 4.

[0117] In this modification, the power supply PS3, the bias electrode layer 31, and the monitor unit M function as a detection unit DT. An X-ray detection signal (radiation detection signal) DS is generated by a current flowing from the power supply PS3 to the bias electrode layer 31. The X-ray detector 1 shown in FIG. 17 can also achieve the same effects as those of the second embodiment.

[0118] To monitor the bias current, the current on the primary circuit side of the power supply PS3 can be monitored, or a shunt circuit can be attached to the high voltage circuit to monitor the current. The principle of detecting the X-ray detection signal (radiation detection signal) DS is the same for both direct conversion type X-ray detectors 1 and indirect conversion type X-ray detectors 1.

[0119] Although several embodiments of the present invention have been described, these embodiments are presented as examples and are not intended to limit the scope of the invention. These novel embodiments can be embodied in various other forms, and various omissions, substitutions, and modifications can be made without departing from the spirit of the invention. These embodiments and their modifications are included within the scope and spirit of the invention, and are also included in the scope of the invention and its equivalents as defined in the claims.

[0120] For example, the above-described technology is not limited to application to the X-ray detection panel PNL, the X-ray detection module 10, and the X-ray detector 1, but can be applied to various radiation detection panels such as other X-ray detection panels, various radiation detection modules such as other X-ray detection modules, and various radiation detectors such as other X-ray detectors. The radiation detector may include a radiation detection panel that detects radiation instead of the X-ray detection panel PNL. [Explanation of symbols]

[0121] 1...X-ray detector, 10...X-ray detection module, PNL...X-ray detection panel, 2...conversion board, 2a...substrate, 2b...first conversion section, 2c...second conversion section, 4...image transmission section, 5...scintillator layer, 6...light reflection layer, 7...moisture-proof cover, 11...circuit board, 11a...drive circuit, 11b...detection circuit, 13a...TFT, 15a, 15b...TFD, 17...capacitor, 30...X-ray conversion layer, 31...bias electrode layer, BL1...bias line, CL...conductive layer, DT...detection section, G...scanning line, U...signal line, PX...image element, M...monitor unit, SM...storage unit, PS1, PS2, PS3...power supply, DS...X-ray detection signal, I1, I2, I3...current, S1...scanning signal, S1a...first scanning signal, S1b...second scanning signal, S2...image data signal, T1...first scanning cycle, T2...detection cycle, Ta...first accumulation period, Tb...second accumulation period, Tg...first readout period, Tg'...second readout period, XR1...continuous pulsed X-ray, XR2...single X-ray, XR3...continuous X-ray, DA...effective area (detection area), X...row direction, Y...column direction.

Claims

1. A substrate; a plurality of scanning lines disposed above the substrate; a plurality of signal lines provided above the substrate; a plurality of pixels provided above the substrate, each of the pixels including: a switching element electrically connected to a corresponding one of the plurality of scanning lines and a corresponding one of the plurality of signal lines; a pixel electrode electrically connected to the switching element; and a first conversion unit electrically connected to the pixel electrode, provided in a region overlapping the pixel electrode, and configured to convert incident radiation into an image data signal; a driving circuit electrically connected to the plurality of scanning lines and configured to sequentially scan the plurality of scanning lines; a detection circuit electrically connected to the plurality of signal lines and detecting the image data signals generated by the plurality of pixels; a detection unit that detects radiation and generates a radiation detection signal; a control circuit for controlling output of the image data signal based on the radiation detection signal, Radiation detector.

2. the control circuit controls driving of the drive circuit based on the radiation detection signal, and automatically adjusts a first scanning period, which is a period for scanning the plurality of scanning lines; The radiation detector according to claim 1 .

3. a radiation conversion layer that is in contact with the pixel electrodes of the pixels and converts incident radiation into charges having a magnitude corresponding to the intensity of the radiation; a plurality of capacitors, each of the capacitors electrically connected to one of the plurality of pixel electrodes; a bias electrode layer capable of applying a bias voltage to the radiation conversion layer; a first power supply electrically connected to the bias electrode layer; In each of the pixels, the first conversion unit is a portion of the radiation conversion layer that overlaps the pixel electrode, one or more of the plurality of first conversion units of the plurality of pixels, the bias electrode layer, and the first power supply also serve as the detection unit; the radiation detection signal is generated by a current flowing from the portion of the radiation conversion layer to the pixel electrode in contact with the portion. The radiation detector according to claim 1 .

4. a radiation conversion layer that is in contact with the pixel electrodes of the pixels and converts incident radiation into charges having a magnitude corresponding to the intensity of the radiation; a plurality of capacitors, each of the capacitors electrically connected to one of the plurality of pixel electrodes; a bias electrode layer capable of applying a bias voltage to the radiation conversion layer; a first power supply electrically connected to the bias electrode layer; a monitor unit electrically connected between the first power supply and the bias electrode layer, and configured to monitor a current flowing from the first power supply to the bias electrode layer; In each of the pixels, the first conversion unit is a portion of the radiation conversion layer that overlaps the pixel electrode, the first power supply, the bias electrode layer, and the monitor unit function as the detection unit, the radiation detection signal is generated by a current flowing from the first power supply to the bias electrode layer. The radiation detector according to claim 1 .

5. a scintillator layer that converts incident radiation into visible light; a plurality of conversion elements each converting visible light incident from the scintillator layer side into an electric charge having a magnitude corresponding to the intensity of the visible light; In each of the pixels, the first conversion unit includes one of the plurality of conversion elements electrically connected to the pixel electrode and a portion of the scintillator layer overlapping the conversion element, one or more of the plurality of first conversion units of the plurality of pixels also serve as the detection unit; the radiation detection signal is generated by a current flowing from the conversion element to the pixel electrode connected to the conversion element. The radiation detector according to claim 1 .

6. a scintillator layer that converts incident radiation into visible light; a plurality of conversion elements each converting visible light incident from the scintillator layer side into charges of a magnitude corresponding to the intensity of the visible light; a bias line electrically connected to the plurality of conversion elements; a first power supply electrically connected to the bias line; a monitor unit electrically connected between the first power supply and the bias line, and configured to monitor a current flowing from the first power supply to the bias line; In each of the pixels, the first conversion unit includes one of the plurality of conversion elements electrically connected to the pixel electrode and a portion of the scintillator layer overlapping the conversion element, the first power supply, the monitor unit, the bias line, and the plurality of first conversion units function as the detection unit; the radiation detection signal is generated by a current flowing from the first power supply to the bias line; The radiation detector according to claim 1 .

7. the radiation detection signal is generated by a current flowing through one or more of the plurality of signal lines; The radiation detector according to claim 1 .

8. A second power source; a second conversion unit electrically connected to the second power supply and configured to convert incident radiation into electric charges independently of the plurality of pixels; a monitor unit electrically connected between the second power supply and the second conversion unit and configured to monitor a current flowing from the second power supply to the second conversion unit; the second power supply, the monitor unit, and the second conversion unit function as the detection unit, the radiation detection signal is generated by a current flowing from the second power supply to the second conversion unit. The radiation detector according to claim 1 .

9. the drive circuit scans the plurality of scanning lines in a first scanning period under the control of the control circuit; When the control circuit determines that the detection period for detecting the radiation detection signal is shorter than the first scanning period, the drive circuit scans the plurality of scanning lines at a second scanning period that is shorter than the first scanning period under the control of the control circuit; The time duration of the second scanning period is the same as the time duration of the detection period. The radiation detector according to claim 1 .

10. The time duration of the first scanning period is: a first accumulation period during which the drive circuit does not scan the plurality of scanning lines and charges can be accumulated in the plurality of pixels; a first readout period, which is a period different from the first accumulation period, during which the drive circuit sequentially scans the plurality of scanning lines and the detection circuit detects the image data signal; The time period of the second scan period is: a second accumulation period during which the drive circuit does not scan the plurality of scanning lines and charges can be accumulated in the plurality of pixels; a second readout period, which is a period different from the second accumulation period, during which the drive circuit sequentially scans the plurality of scanning lines and the detection circuit detects the image data signal; the drive circuit, under the control of the control circuit, adjusts the first accumulation period to the second accumulation period, adjusts the first readout period to the second readout period, and adjusts the timing of the second accumulation period so that the second accumulation period overlaps with a period for detecting the radiation detection signal; The radiation detector according to claim 9.

11. When the control circuit determines that the detection period is shorter than the first scanning period, it permits output of the image data signal. The radiation detector according to claim 9.

12. the drive circuit scans the plurality of scanning lines in a first scanning period under the control of the control circuit; When the control circuit determines that the time period of the detection cycle for detecting the radiation detection signal is the same as the time period of the first scanning cycle, the drive circuit continues scanning the plurality of scanning lines in the first scanning period under the control of the control circuit; The radiation detector according to claim 1 .

13. the detection unit resets the radiation detection signal every time the time period of the first scanning cycle elapses under the control of the control circuit; The radiation detector of claim 12.

14. the control circuit permits output of the image data signal when determining that the time period of the detection cycle is the same as the time period of the first scanning cycle; The radiation detector of claim 12.

15. the drive circuit scans the plurality of scanning lines in a first scanning period under the control of the control circuit; When the control circuit determines that the detection period for detecting the radiation detection signal is longer than the first scanning period, the drive circuit continues scanning the plurality of scanning lines in the first scanning period under the control of the control circuit; The radiation detector according to claim 1 .

16. further comprising a storage unit for storing the image data signal; If a period longer than the first scanning period is taken as a reference period, then: when the control circuit determines that the detection period for detecting the radiation detection signal is longer than the reference period, it permits output of the image data signal stored in the storage unit. The radiation detector of claim 15.

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