JP2016156719A - Radiation detector - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a radiation detector capable of high quality energy subtraction imaging.SOLUTION: The radiation detector includes: a first detection section which has a first scintillator covering plural photoelectric conversion parts formed on an array board; and a second detection section which has a second scintillator covering plural photoelectric conversion parts formed on the array board, and which is formed at the opposite side to the side where the first scintillator of the first detection section is provided. The thickness of the second scintillator is smaller than the thickness of the first scintillator.SELECTED DRAWING: Figure 1

Description

  Embodiments described herein relate generally to a radiation detector.

  One type of radiation detector is an X-ray detector. Many of the X-ray detectors in practical use currently employ an indirect conversion method. An indirect conversion type X-ray detector converts X-rays that have passed through and incident on a subject such as a human body into fluorescence (visible light) by a scintillator, and converts the fluorescence into signal charges by a photoelectric conversion element such as a photodiode. The signal charge is stored in the storage capacitor. The accumulated signal charge is output to the outside by a switching element (thin film transistor (TFT)). A plurality of sets of photoelectric conversion elements, storage capacitors, and switching elements are provided in a matrix. A set of photoelectric conversion elements, storage capacitors, and switching elements correspond to one pixel.

The X-ray detector images the inside of the subject using the difference in the X-ray transmittance in the tissue and structure inside the subject.
In imaging using an X-ray detector, an imaging technique called energy subtraction may be used.
In energy subtraction imaging, imaging of the same subject is performed a plurality of times.
For example, first, in a state where the subject is fixed, a low tube voltage is applied to the X-ray tube which is an X-ray source so that the X-ray energy spectrum is positioned on the low energy side, and X-rays are generated. X-rays transmitted through the subject are imaged using an X-ray detector.
Next, in a state where the subject is fixed, a high tube voltage is applied to the X-ray tube as an X-ray source to generate X-rays so that the X-ray energy spectrum is located on the higher energy side, and the subject The X-rays transmitted through are imaged using an X-ray detector.
In this case, the X-ray absorptance at different parts of the composition inside the subject varies depending on the quality of the irradiated X-rays. Therefore, a desired part can be selectively extracted from the difference between the X-ray image obtained by the first imaging and the X-ray image obtained by the next imaging.
When energy subtraction imaging is performed, for example, a part mainly containing light elements such as fat and blood can be selectively extracted, or a part including a relatively heavy element such as bone can be selectively extracted. Can do.

Here, it takes several seconds to change the tube voltage of the X-ray tube by the generator. Also, it takes several seconds to generate an X-ray image in the X-ray detector.
For this reason, in energy subtraction imaging, it takes several seconds or more to complete the first imaging and X-ray image generation, and it is necessary to keep the subject fixed until the next imaging.
However, when the subject is a human or the like, there is a movement such as breathing, so it is difficult to fix the subject for several seconds or more.
If the subject moves during energy subtraction imaging, the quality of the extracted X-ray image of the part may be lowered.

In this case, if the time required for changing the tube voltage of the X-ray tube is shortened, the influence of the movement of the subject can be reduced.
However, it is difficult for a general generator to change the tube voltage of the X-ray tube in a short time. Therefore, a special generator is required for energy subtraction imaging. As a result, the cost of the photographing system increases.
Thus, it has been desired to develop a radiation detector capable of performing high-quality energy subtraction imaging even when a general generator is used.

JP-A-8-313640

  The problem to be solved by the present invention is to provide a radiation detector capable of performing high-quality energy subtraction imaging.

The radiation detector according to the embodiment includes a first detection unit having a first scintillator that covers a plurality of photoelectric conversion units provided on the array substrate, and a second that covers the plurality of photoelectric conversion units provided on the array substrate. And a second detection unit provided opposite to the first detection unit on the side where the first scintillator is provided.
The thickness dimension of the second scintillator is shorter than the thickness dimension of the first scintillator.

It is a mimetic diagram for illustrating X-ray detector 1 concerning this embodiment. 1 is a schematic perspective view for illustrating an X-ray detector 1. FIG. 2 is a circuit diagram of an array substrate 2. FIG. 2 is a block diagram of the X-ray detector 1. FIG. It is a schematic diagram for demonstrating arrangement | positioning of the detection part 21 and the detection part 22 which concern on other embodiment. It is a schematic diagram for illustrating the detection part 21a which concerns on other embodiment. It is a model perspective view for illustrating the arrangement form of connection wiring parts 2e1 and 2e2. It is a model perspective view for illustrating the arrangement form of connection wiring parts 2e1 and 2e2 concerning other embodiments. FIG. 9 is a schematic cross-sectional view for illustrating a state in which the X-ray detector 1 illustrated in FIG. 8 is stored in a housing 6.

Hereinafter, embodiments will be illustrated with reference to the drawings. In addition, in each drawing, the same code | symbol is attached | subjected to the same component and detailed description is abbreviate | omitted suitably.
The radiation detector according to the present embodiment can be applied to various types of radiation such as γ rays in addition to X-rays. Here, as an example, a case of X-rays as a representative example of radiation will be described as an example. Therefore, by replacing “X-ray” in the following embodiments with “other radiation”, the present invention can be applied to other radiation.
Moreover, although the X-ray detector 1 which is a radiation detector can be used for a general medical use etc., for example, there is no limitation in a use.

FIG. 1 is a schematic diagram for illustrating an X-ray detector 1 according to the present embodiment.
In FIG. 1, the connection wiring portions 2e1, 2e2, etc. are omitted to avoid complication.
FIG. 2 is a schematic perspective view for illustrating the X-ray detector 1.
In FIG. 2, the detection unit 22 and the light shielding unit 23 are omitted in order to avoid complication.
FIG. 3 is a circuit diagram of the array substrate 2.
FIG. 4 is a block diagram of the X-ray detector 1.
In FIG. 4, in order to avoid complexity, the detection unit 22 is omitted.

As shown in FIGS. 1 and 2, the X-ray detector 1 includes a detection unit 21 (corresponding to an example of a first detection unit), a detection unit 22 (corresponding to an example of a second detection unit), A light shielding unit 23, a signal processing unit 3, and an image generation unit 4 are provided.
The detection unit 21 is provided with an array substrate 2 and a scintillator 5 (corresponding to an example of a first scintillator).
The array substrate 2 converts visible light converted from the X-rays 102 by the scintillator 5, that is, fluorescence, into signal charges.
The array substrate 2 includes a substrate 2a, a photoelectric conversion unit 2b, a control line (or gate line) 2c1, a data line (or signal line) 2c2, and a bias line 2c3.

The substrate 2a has a plate shape and is made of a translucent material such as glass.
A plurality of photoelectric conversion units 2b are provided on one surface of the substrate 2a.
The photoelectric conversion unit 2b has a rectangular shape and is provided in a region defined by the control line 2c1 and the data line 2c2. The plurality of photoelectric conversion units 2b are arranged in a matrix.
One photoelectric conversion unit 2b corresponds to one pixel.

Each of the plurality of photoelectric conversion units 2b is provided with a photoelectric conversion element 2b1 and a thin film transistor 2b2 which is a switching element.
Further, as shown in FIG. 3, a storage capacitor 2b3 for storing the signal charge converted in the photoelectric conversion element 2b1 can be provided. The storage capacitor 2b3 has, for example, a rectangular flat plate shape and can be provided under the thin film transistor 2b2. However, depending on the capacitance of the photoelectric conversion element 2b1, the photoelectric conversion element 2b1 can also serve as the storage capacitor 2b3.

The photoelectric conversion element 2b1 can be, for example, a photodiode.
The thin film transistor 2b2 performs switching between accumulation and emission of electric charges generated when fluorescence enters the photoelectric conversion element 2b1. The thin film transistor 2b2 can include a semiconductor material such as amorphous silicon (a-Si) or polysilicon (P-Si).
As shown in FIG. 3, the thin film transistor 2b2 includes a gate electrode 2b2a, a source electrode 2b2b, and a drain electrode 2b2c. Gate electrode 2b2a of thin film transistor 2b2 is electrically connected to corresponding control line 2c1. The source electrode 2b2b of the thin film transistor 2b2 is electrically connected to the corresponding data line 2c2. The drain electrode 2b2c of the thin film transistor 2b2 is electrically connected to the corresponding photoelectric conversion element 2b1 and the storage capacitor 2b3.

A plurality of control lines 2c1 are provided in parallel with each other at a predetermined interval. For example, the control line 2c1 extends in the row direction.
The plurality of control lines 2c1 are electrically connected to each of the plurality of wiring pads 2d1 provided near the periphery of the substrate 2a. One end side of the plurality of wirings provided in the connection wiring part 2e1 is electrically connected to the plurality of wiring pads 2d1, respectively. The other end side of the plurality of wirings provided in the connection wiring part 2e1 is electrically connected to the control part 31 provided in the signal processing part 3, respectively.
The connection wiring part 2e1 can be a flexible printed circuit board, for example.

A plurality of data lines 2c2 are provided in parallel with each other at a predetermined interval. The data line 2c2 extends in the column direction, for example.
The plurality of data lines 2c2 are electrically connected to the plurality of wiring pads 2d2 provided near the periphery of the substrate 2a. One end side of the plurality of wirings provided in the connection wiring part 2e2 is electrically connected to the plurality of wiring pads 2d2. The other end side of the plurality of wirings provided in the connection wiring part 2e2 is electrically connected to the amplification / conversion circuit 32 provided in the signal processing part 3, respectively.
The connection wiring part 2e2 can be a flexible printed circuit board, for example.

A plurality of bias lines 2c3 are provided in parallel to each other with a predetermined interval. The bias line 2c3 extends in the same direction as the data line 2c2. The bias line 2c3 is electrically connected to the corresponding photoelectric conversion element 2b1 and the storage capacitor 2b3. The plurality of bias lines 2c3 are electrically connected to a bias power source (not shown).
The control line 2c1, the data line 2c2, and the bias line 2c3 can be formed using a low resistance metal such as aluminum or chromium.
Details regarding the arrangement of the connection wiring portions 2e1 and 2e2 will be described later.

In addition, a protective layer (not shown) that covers the photoelectric conversion unit 2b, the control line 2c1, the data line 2c2, and the bias line 2c3 can be provided.
The protective layer can be formed of an insulating material such as silicon nitride (SiN) or acrylic resin, for example.

The scintillator 5 is provided on the plurality of photoelectric conversion units 2b, and converts the X-rays 102 transmitted through the detection unit 22 and the light shielding unit 23 into visible light, that is, fluorescence. The scintillator 5 is provided so as to cover a region where the plurality of photoelectric conversion units 2b are provided on the substrate 2a.
The scintillator 5 can be formed using, for example, cesium iodide (CsI): thallium (Tl) or sodium iodide (NaI): thallium (Tl). In this case, an aggregate of columnar crystals can be formed using a vacuum deposition method or the like. If an aggregate of columnar crystals is formed, scattering of generated fluorescence can be suppressed. If the generated fluorescence can be suppressed from being scattered, the generated fluorescence can be suppressed from entering the adjacent photoelectric conversion unit 2b, so that the resolution can be improved.

The scintillator 5 can also be formed using, for example, gadolinium oxysulfide (Gd ₂ O ₂ S). The scintillator 5 made of gadolinium oxysulfide has a granular shape (columnar shape) and can be provided for each of the plurality of photoelectric conversion units 2b.
If the scintillator 5 is made of gadolinium oxysulfide and is an aggregate of granular bodies (columnar bodies), it is possible to improve sensitivity characteristics and productivity.

In addition, a reflection layer (not shown) can be provided so as to cover the surface side of the scintillator 5 (the incident surface side of the X-ray 102) in order to improve the use efficiency of fluorescence and improve the sensitivity characteristics.
Moreover, in order to suppress the deterioration of the characteristics of the scintillator 5 and the reflection layer (not shown) due to water vapor contained in the air, a moisture barrier (not shown) that covers the scintillator 5 and the reflection layer (not shown) can be provided.

The detection unit 22 is provided opposite to the side of the detection unit 21 where the scintillator 5 is provided. The detection unit 22 is provided on the side of the detection unit 21 where the X-ray 102 is incident.
In this case, the detection unit 22 can be provided so that the scintillator 15 (corresponding to an example of a second scintillator) is on the side opposite to the detection unit 21 side (X-ray tube 100 side).
The detection unit 22 is provided with the array substrate 2 and the scintillator 15 described above. The scintillator 15 is provided on the plurality of photoelectric conversion units 2b and converts incident X-rays 102 into fluorescence. The scintillator 15 is provided so as to cover an area on the substrate 2a where the plurality of photoelectric conversion units 2b are provided.
The thickness dimension H2 of the scintillator 15 is shorter than the thickness dimension H1 of the scintillator 5.
Details regarding the thickness dimension H1 of the scintillator 5 and the thickness dimension H2 of the scintillator 15 will be described later.

The material and configuration of the scintillator 15 can be the same as the material and configuration of the scintillator 5 described above, but the following is preferable.
In the thin scintillator 15, scattering of generated fluorescence hardly occurs. Therefore, the scintillator 15 is made of the above-described gadolinium oxysulfide or the like and can include a granular material. In this way, it is possible to improve sensitivity characteristics and productivity in the detection unit 22.
In the thick scintillator 5, scattering of generated fluorescence is likely to occur. Therefore, the scintillator 5 can include the aggregate of columnar crystals described above. In this way, the resolution in the detection unit 21 can be improved.

The light shielding unit 23 suppresses the fluorescence generated in the scintillator 15 from being transmitted through the substrate 2 a and entering the detection unit 21.
The light shielding unit 23 suppresses the transmission of fluorescence from the scintillator 15 and transmits the X-rays 102 transmitted through the detection unit 22.
Moreover, the light-shielding part 23 can also suppress the fluorescence from the scintillator 5 from being transmitted to the detection part 22 side.
The light shielding unit 23 is provided between the detection unit 21 and the detection unit 22.
The light shielding part 23 may be a plate.
The material of the light shielding part 23 is not particularly limited as long as it can transmit the X-ray 102 and suppress the transmission of fluorescence. The material of the light shielding portion 23 can be, for example, an opaque resin or a metal such as aluminum.

  The thickness dimension and material of the light shielding part 23 can be appropriately set in consideration of the transparency of the X-ray 102 and the fluorescence light shielding property. For example, by performing experiments and simulations, the thickness dimension and material of the light shielding portion 23 can be determined so that desired X-ray transparency and fluorescence light shielding properties can be obtained.

Further, as shown in FIG. 1, there may be a gap between the light shielding part 23 and the detection part 22, or the light shielding part 23 and the detection part 22 may be in close contact as shown in FIG.
As shown in FIG. 1, there may be a gap between the light shielding unit 23 and the detection unit 21, or the light shielding unit 23 and the detection unit 21 may be in close contact as shown in FIG. 9. However, if a gap is provided between the light shielding unit 23 and the detection unit 21, external vibrations are transmitted to the side of the scintillator 5 or the array substrate 2 where the photoelectric conversion unit 2 b is provided via the light shielding unit 23. Can be suppressed.

The signal processing unit 3 is provided on the opposite side of the detection unit 21 from the side on which the X-ray 102 is incident.
As shown in FIG. 4, the signal processing unit 3 includes a control unit 31 and an amplification / conversion circuit 32.
The control unit 31 is electrically connected to the detection unit 21 and the detection unit 22 via the connection wiring unit 2e1 and the connection wiring unit 2e2.

The control unit 31 selectively applies the control signal S1 to the plurality of photoelectric conversion units 2b via the plurality of control lines 2c1.
The control unit 31 includes a plurality of gate drivers 31a and a row selection unit 31b.
The gate driver 31a applies the control signal S1 to the corresponding control line 2c1.
The row selection unit 31b sends a control signal S1 to the corresponding gate driver 31a according to the scanning direction of the X-ray image.
For example, the control unit 31 sequentially applies the control signal S1 to each control line 2c1 via the connection wiring unit 2e1 and the control line 2c1. The thin film transistor 2b2 is turned on by the control signal S1 applied to the control line 2c1, and the signal charge (image data signal S2) from the photoelectric conversion element 2b1 can be received. (See Figure 2)
The amplification / conversion circuit 32 includes a plurality of integrating amplifiers 32a and a plurality of A / D converters 32b.
The integrating amplifier 32a receives the image data signal S2 from each of the plurality of photoelectric conversion elements 2b1 via the data line 2c2 and the connection wiring portion 2e2, and amplifies and outputs the received image data signal S2. The image data signal S2 output from the integrating amplifier 32a is parallel / serial converted and input to the A / D converter 32b.
The A / D converter 32b converts the input image data signal S2 (analog signal) into a digital signal.

The image generation unit 4 is electrically connected to the signal processing unit 3 via the wiring 4a.
The image generation unit 4 generates an X-ray image based on the image data signal S2 converted into digital signals by the plurality of A / D converters 32b. For example, the image generation unit 41 generates an X-ray image by sequentially processing the image data signal S2 in accordance with the rows and columns of the plurality of photoelectric conversion units 2b arranged in a matrix.
The image generation unit 4 can also perform predetermined correction (for example, offset correction, gain correction, defect correction, etc.) when generating an X-ray image.
The generated X-ray image is output from the image generation unit 4 to the outside.

Next, the thickness dimension H1 of the scintillator 5 and the thickness dimension H2 of the scintillator 15 will be further described.
As shown in FIG. 1, the generator 101 applies a predetermined tube voltage to the X-ray tube 100. The X-ray tube 100 emits X-rays 102 having a predetermined intensity and energy (a predetermined energy spectrum) in accordance with the applied tube voltage.
A part of the X-ray 102 emitted from the X-ray tube 100 is absorbed when passing through the subject 300, and the X-ray 102 that has not been absorbed passes through the subject 300 and enters the X-ray detector 1.

Here, the X-ray 102 emitted from the X-ray tube 100 includes a low energy component 102a and a high energy component 102b.
Therefore, the X-ray 102 that has passed through the subject 300 also includes the low energy component 102a and the high energy component 102b.
The X-ray absorptance at different parts of the subject 300 differs in X-ray quality, that is, the low energy component 102a and the high energy component 102b.
Therefore, the detection unit 22 mainly generates an X-ray image using the low energy component 102a of the X-ray 102, and the detection unit 21 generates an X-ray image mainly using the high energy component 102b of the X-ray 102. If so, energy subtraction imaging can be performed with a single X-ray irradiation.

  The low-energy component 102a of the X-ray 102 is more easily absorbed (has less transmission power) than the high-energy component 102b. Therefore, the scintillator 15 of the detection unit 22 near the X-ray tube 100 mainly converts the low energy component 102a of the X-ray 102 into fluorescence. The low energy component 102a that is easily absorbed does not reach the detection unit 21 because it is converted into fluorescence by the scintillator 15. On the other hand, a part of the high energy component 102 b that is difficult to be absorbed is absorbed by the scintillator 15, but most of the high energy component 102 b passes through the detection unit 22 and reaches the detection unit 21. Therefore, the scintillator 5 of the detection unit 21 mainly converts the high energy component 102b of the X-ray 102 into fluorescence.

In this case, if the amount of absorption (conversion amount) of the high energy component 102 b increases in the scintillator 15, it may be difficult to generate an X-ray image by the detection unit 21.
In order to reduce the amount of absorption (conversion amount) of the high energy component 102b in the scintillator 15, the thickness dimension H2 of the scintillator 15 may be shortened.
Even if the thickness dimension H2 of the scintillator 15 is shortened, the low energy component 102a is more easily absorbed than the high energy component 102b, so that sufficient fluorescence can be obtained.

On the other hand, since the high energy component 102b of the X-ray 102 is less likely to be absorbed than the low energy component 102a, it is preferable to make the thickness dimension H1 of the scintillator 5 longer in order to obtain sufficient fluorescence.
Therefore, the thickness dimension H2 of the scintillator 15 is shorter than the thickness dimension H1 of the scintillator 5.

  In this way, the detection unit 22 generates an X-ray image mainly using the low energy component 102a of the X-ray 102, and the detection unit 21 mainly uses the high energy component 102b of the X-ray 102. Can be generated.

The thickness dimension H1 of the scintillator 5 and the thickness dimension H2 of the scintillator 15 include the material of the scintillator 5 and the scintillator 15, the ratio and strength of the low energy component 102a and the high energy component 102b, the transmittance of the high energy component 102b in the light shielding portion 23, and the like. It can be appropriately determined in consideration.
For example, it is possible to appropriately determine the thickness dimension H1 of the scintillator 5 and the thickness dimension H2 of the scintillator 15 by performing experiments and simulations in consideration of these factors.

Next, the operation of the X-ray detector 1 will be described.
X-rays 102 are emitted from an X-ray tube 100 provided outside the X-ray detector 1.
For example, the generator 101 applies a predetermined tube voltage to the X-ray tube 100. The X-ray tube 100 emits X-rays 102 having a predetermined intensity and energy according to the applied tube voltage.
A part of the X-ray 102 emitted from the X-ray tube 100 is absorbed when passing through the subject 300, and the X-ray 102 that has not been absorbed passes through the subject 300 and enters the X-ray detector 1.

The low energy component 102 a of the X-ray 102 incident on the X-ray detector 1 is converted into fluorescence by the scintillator 15 of the detection unit 22. The intensity of the fluorescence is proportional to the intensity of the low energy component 102 a of the incident X-ray 102. A part of the fluorescence generated inside the scintillator 15 enters the corresponding photoelectric conversion element 2b1.
The fluorescence that has entered the photoelectric conversion element 2b1 is converted into electric charges composed of electrons and holes inside the photoelectric conversion element 2b1. The converted charge is stored in the storage capacitor 2b3.

The control unit 31 selectively applies the control signal S1 to the plurality of photoelectric conversion units 2b. The thin film transistor 2b2 of the selected photoelectric conversion unit 2b is turned on, and the accumulated charge is output from the storage capacitor 2b3 as the image data signal S2.
The integrating amplifier 32a receives the image data signal S2 from each of the plurality of photoelectric conversion elements 2b1, and amplifies and outputs the received image data signal S2. The image data signal S2 output from the integrating amplifier 32a is parallel / serial converted and input to the A / D converter 32b. The A / D converter 32b converts the input image data signal S2 (analog signal) into a digital signal.

  The image generation unit 4 generates a first X-ray image based on the low energy component 102a based on the image data signal S2 converted into digital signals by the plurality of A / D converters 32b. For example, the image generation unit 41 sequentially processes the image data signal S2 according to the rows and columns of the plurality of photoelectric conversion units 2b arranged in a matrix to generate a first X-ray image.

The high energy component 102 b of the X-ray 102 that has entered the X-ray detector 1 passes through the detection unit 22 and the light shielding unit 23 and enters the scintillator 5 of the detection unit 21.
A part of the fluorescence generated inside the scintillator 15 passes through the array substrate 2 and travels toward the detection unit 21, but incidence to the scintillator 5 is suppressed by the light shielding unit 23.
The high energy component 102b of the X-ray 102 incident on the scintillator 5 is converted into fluorescence by the scintillator 5. Thereafter, the second X-ray image by the high energy component 102b is generated in the same manner as described above.
In this way, the X-ray detector 1 can perform energy subtraction imaging with a single X-ray 102 irradiation.
The generated first X-ray image and second X-ray image are output from the image generation unit 4 to the outside.

The X-ray absorptance at different parts of the subject 300 differs between the low energy component 102a and the high energy component 102b.
Therefore, a desired part can be selectively extracted from the difference between the first X-ray image and the second X-ray image.
For example, it is possible to selectively extract a site mainly composed of light elements such as fat and blood, or to selectively extract a site containing a relatively large amount of heavy elements such as bone.

Since the X-ray detector 1 according to the present embodiment can perform energy subtraction imaging with a single X-ray 102 irradiation, imaging time in energy subtraction imaging can be shortened. Therefore, a high-quality X-ray image can be obtained even with a moving subject 300.
That is, the X-ray detector 1 according to the present embodiment can perform high-quality energy subtraction imaging.
Further, since it is only necessary to perform X-ray irradiation once, that is, a predetermined tube voltage may be applied to the X-ray tube 100, it is not necessary to change the tube voltage applied to the X-ray tube 100 during imaging.
Therefore, there is no need to use a special generator that can change the tube voltage of the X-ray tube 100 in a short time, so that the cost of the imaging system does not increase.

FIG. 5 is a schematic diagram for illustrating the arrangement of the detection unit 21 and the detection unit 22 according to another embodiment.
As illustrated in FIG. 5, the detection unit 22 is provided on the side of the detection unit 21 on which the X-ray 102 is incident.

However, the detection unit 22 is provided so that the scintillator 15 is on the detection unit 21 side (the side opposite to the X-ray tube 100 side).
That is, the detection unit 22 is provided with the scintillator 15 facing the scintillator 5.
Therefore, the surface of the array substrate 2 provided in the detection unit 22 on the side where the photoelectric conversion unit 2b is provided and the surface of the array substrate 2 provided in the detection unit 21 on the side where the photoelectric conversion unit 2b is provided face each other.

  Here, when the distance between the photoelectric conversion unit 2b included in the detection unit 22 and the photoelectric conversion unit 2b included in the detection unit 21 is increased, pixel deviation is likely to occur. Further, the pixel shift becomes significant in the peripheral region of the array substrate 2 where the X-rays 102 are incident from an oblique direction. Pixel shift becomes a factor of resolution degradation.

In the present embodiment, since the surfaces on the side of the array substrate 2 where the photoelectric conversion units 2b are provided are opposed to each other, the photoelectric conversion unit 2b included in the detection unit 22 and the photoelectric conversion unit 2b included in the detection unit 21 The distance between can be shortened.
Therefore, pixel shift can be reduced, and resolution can be improved.

FIG. 6 is a schematic diagram for illustrating a detection unit 21a according to another embodiment.
As shown in FIG. 6, the detection unit 21a is provided with an array substrate 12a and a scintillator 5.
The detector 22a is provided with an array substrate 12b and a scintillator 15.
The pitch dimension Pb of the photoelectric conversion unit 2b provided on the array substrate 12b is shorter than the pitch dimension Pa of the photoelectric conversion unit 2b provided on the array substrate 12a.
The pitch dimension Pa and the pitch dimension Pb consider the radiation angle θ of the X-ray 102.

Here, as described above, the pixel shift becomes significant in the peripheral regions of the array substrates 12a and 12b where the X-rays 102 are incident from an oblique direction.
In the present embodiment, the pitch dimension Pb is shorter than the pitch dimension Pa, and the pitch dimension Pa and the pitch dimension Pb are considered in consideration of the radiation angle θ of the X-ray 102. Therefore, the peripheral regions of the array substrates 12a and 12b The pixel shift in can be reduced. Therefore, the resolution can be improved.

Next, the arrangement of the connection wiring portions 2e1 and 2e2 will be further described.
FIG. 7 is a schematic perspective view for illustrating the arrangement of the connection wiring portions 2e1 and 2e2.
In FIG. 7, in order to avoid complication, the elements other than the connection wiring parts 2 e 1 and 2 e 2, the array substrate 2, the detection part 21, the detection part 22, and the signal processing part 3 are omitted.
As shown in FIG. 7, the connection wiring part 2 e 1 provided in the detection unit 21 can be provided so as to face the connection wiring part 2 e 1 provided in the detection unit 22.
The connection wiring part 2e2 provided in the detection part 21 can be provided so as to face the connection wiring part 2e2 provided in the detection part 22.
In this way, each of the detection unit 21 and the detection unit 22 can be connected to the signal processing unit 3. Therefore, the detection unit 21 and the detection unit 22 can be controlled separately.

FIG. 8 is a schematic perspective view for illustrating an arrangement form of the connection wiring portions 2e1 and 2e2 according to another embodiment.
FIG. 9 is a schematic cross-sectional view for illustrating a state in which the X-ray detector 1 illustrated in FIG.

As described above, the connection wiring portion 2e1 electrically connects the control unit 31 and the control line 2c1.
Here, the control unit 31 selectively applies the control signal S1 to the plurality of photoelectric conversion units 2b. The control signal S1 is a signal for outputting the image data signal S2 from the desired photoelectric conversion unit 2b. Therefore, the control signal S1 can be applied to the photoelectric conversion unit 2b provided in the detection unit 21 and the photoelectric conversion unit 2b provided in the detection unit 22 at the same timing.
That is, the control unit 31 applies the control signal S1 to the control line 2c1 provided on the array substrate 2 of the detection unit 21 and the control line 2c1 provided on the array substrate 2 of the detection unit 22 at the same timing. May be.

In the present embodiment, the connection wiring part 2 e 1 is shared by the detection part 21 and the detection part 22.
That is, one end side of the connection wiring part 2 e 1 is electrically connected to the control part 31. The other end of the connection wiring portion 2e1 is electrically connected to a control line 2c1 provided on the array substrate 2 of the detection unit 21 and a control line 2c1 provided on the array substrate 2 of the detection unit 22.
Therefore, since wiring can be simplified, wiring work can be facilitated, and hence productivity of the X-ray detector 1 can be improved.
Further, since the control unit 31 can share a circuit that outputs the control signal S1, the control unit 31 can be simplified, and the manufacturing cost of the control unit 31 can be reduced.

As shown in FIG. 9, the housing 6 has a box shape, and the X-ray detector 1 is stored therein.
A support plate 6b is provided inside the housing 6, and a detection unit 21, a light shielding unit 23, and a detection unit 22 are provided on one surface of the support plate 6b. The signal processing unit 3 is provided on the other surface of the support plate 6b. A connector 6 a is provided on a side surface of the housing 6 so that the signal processing unit 3 and the image generation unit 4 can be electrically connected.
As shown in FIG. 8, if the connection wiring part 2e1 is made common to the detection part 21 and the detection part 22, one side of the detection part 21 and the detection part 22 is set as a space where the connection wiring part 2e1 is not provided. be able to. Therefore, the wiring 6c that electrically connects the signal processing unit 3 and the connector 6a can be provided in this space. Therefore, the wiring work inside the housing 6 can be facilitated, and the productivity of the X-ray detector 1 can be improved.

  As mentioned above, although several embodiment of this invention was illustrated, these embodiment is shown as an example and is not intending limiting the range of invention. These novel embodiments can be implemented in various other forms, and various omissions, replacements, changes, and the like can be made without departing from the spirit of the invention. These embodiments and modifications thereof are included in the scope and gist of the invention, and are included in the invention described in the claims and equivalents thereof. Further, the above-described embodiments can be implemented in combination with each other.

  1 X-ray detector, 2 array substrate, 2b photoelectric conversion unit, 2b1 photoelectric conversion element, 2b2 thin film transistor, 2c1 control line, 2c2 data line, 2e1 connection wiring unit, 2e2 connection wiring unit, 3 signal processing unit, 4 image generation unit 5, scintillator, 15 scintillator, 21 detection unit, 22 detection unit, 23 light shielding unit, 31 control unit, 31a gate driver, 31b row selection unit, 32 amplification / conversion circuit, 100 X-ray tube, 101 generator, 102 X-ray, 102a Low energy component, 102b High energy component, 300 Subject

Claims (7)

A first detection unit having a first scintillator covering a plurality of photoelectric conversion units provided on the array substrate;
A second detection unit having a second scintillator that covers a plurality of photoelectric conversion units provided on the array substrate, and is provided facing the side of the first detection unit on which the first scintillator is provided When,
With
A radiation detector in which a thickness dimension of the second scintillator is shorter than a thickness dimension of the first scintillator.   Provided between the first detection unit and the second detection unit, suppresses transmission of fluorescence from the second scintillator, and transmits radiation transmitted through the second detection unit. The radiation detector according to claim 1, further comprising a light shielding portion.   3. The radiation detector according to claim 1, wherein the second detection unit is provided with the second scintillator facing the first scintillator. 4. The first scintillator includes an aggregate of columnar crystals,
The radiation detector according to claim 1, wherein the second scintillator includes a granular material. A control unit that applies control signals to a control line provided on the array substrate of the first detection unit and a control line provided on the array substrate of the second detection unit;
The control unit applies the control signal to the control line provided on the array substrate of the first detection unit and the control line provided on the array substrate of the second detection unit at the same timing. The radiation detector as described in any one of Claims 1-4.   One end side is electrically connected to the control unit, and the other end side is electrically connected to a control line provided on the array substrate of the first detection unit and a control line provided on the array substrate of the second detection unit. The radiation detector according to claim 5, further comprising a connection wiring portion connected to the.   The pitch dimension of the plurality of photoelectric conversion units provided on the array substrate of the second detection unit is shorter than the pitch dimension of the plurality of photoelectric conversion units provided on the array substrate of the first detection unit. The radiation detector as described in any one of -6.

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