WO2013015267A1

WO2013015267A1 - Radiographic equipment

Info

Publication number: WO2013015267A1
Application number: PCT/JP2012/068673
Authority: WO
Inventors: 大田　恭義; 中津川　晴康; 西納　直行
Original assignee: 富士フイルム株式会社
Priority date: 2011-07-27
Filing date: 2012-07-24
Publication date: 2013-01-31
Also published as: JP2014195481A

Abstract

In order to rapidly switch between video imaging and still image photography, an electronic cassette (15) is provided with a first radiation detector (40) and a second radiation detector (41). The first radiation detector (40) comprises a light emitting part (33) that absorbs radiation and produces visible light, and a first light detecting part (32) that detects the visible light. The second radiation detector (41) comprises the light emitting part (33), and a second light detecting part (34) that is positioned on the opposite side to the side where the radiation enters. A cassette control unit (70) drives the first radiation detector (40) during video imaging, and drives the second radiation detector (41) during still image photography.

Description

Radiography equipment

The present invention relates to a radiation imaging apparatus capable of moving image shooting and still image shooting.

2. Description of the Related Art In the medical field, a radiation imaging apparatus that captures a subject (a patient's imaging region) using radiation (for example, X-rays) is known for performing image diagnosis. The radiation imaging apparatus includes a radiation generator that emits radiation (for example, X-rays) toward a subject, and a radiation detector that is disposed opposite to the radiation generator and detects and images radiation that has passed through the subject. .

Some of these radiation imaging apparatuses are capable of both moving image shooting (also referred to as fluoroscopic imaging) and still image shooting (also simply referred to as shooting). The dose of radiation emitted from the radiation generator differs between when shooting moving images and when shooting still images. Movie shooting is performed at a low dose, and is used for positioning a patient for still image shooting, searching for a lesion, and the like. Still image capturing is performed at a high dose, and is used to obtain a clear radiographic image of a lesion. Generally, the dose during still image shooting is about 100 times the dose during moving image shooting.

In such a radiographic apparatus capable of capturing moving images and still images, it is necessary to switch the driving mode of the radiation detector depending on whether to perform moving image shooting or still image shooting. In order to easily switch the driving mode, in Patent Document 1, the radiation dose emitted from the radiation generator is monitored on the radiation detector side, and the radiation is synchronized with the timing of shifting from the low dose to the high dose. It has been proposed to switch the detector from the moving image shooting mode to the still image shooting mode. Thereby, the drive mode of a radiation detector can be switched, without performing transmission / reception of a synchronizing signal between a radiation generator and a radiation detector.

However, since the radiographic apparatus described in Patent Document 1 performs moving image shooting and still image shooting using a single radiation detector, the image quality and the still image shooting can be reduced between moving image shooting and still image shooting. The field of view size cannot be changed. In order to solve this problem such as image quality, Patent Documents 2 and 3 disclose a first radiation detector that performs still image shooting, and a first field detector that performs moving image shooting with a smaller field of view than the first radiation detector. A radiation imaging apparatus is described in which two radiation detectors are used and arranged in a radiation region from the radiation generator in a state where they are overlapped. When switching from moving image shooting to still image shooting, the second radiation detector is retracted from the radiation irradiation region from the radiation generator, and then the still image shooting is performed by the first radiation detector.

Japanese Patent Laid-Open No. 2003-307568 JP 2011-004966 A JP 2002-102213 A

However, the radiographic apparatuses described in Patent Documents 2 and 3 require a mechanical mechanism for moving the second radiation detector in and out of the radiation irradiation area in order to switch between moving image capturing and still image capturing. Therefore, switching between moving image shooting and still image shooting cannot be performed quickly, and there is a possibility of missing a shooting opportunity.

An object of the present invention is to provide a radiation imaging apparatus that can quickly and easily switch between moving image shooting and still image shooting.

In order to solve the above problems, a radiation imaging apparatus of the present invention transmits a first radiation detector that detects radiation emitted from a radiation generator and generates image data, and the first radiation detector. A second radiation detector that detects radiation and generates image data; and a control unit that causes the first radiation detector to perform moving image photographing and causes the second radiation detector to perform still image photographing. ing.

In addition, it is preferable to further include a radiation dose measurement unit that measures the dose of the radiation pulse emitted from the radiation generator, and a dose determination unit that compares the dose measured by the radiation dose measurement unit with a predetermined threshold. When the dose determination unit detects a low-dose pulse smaller than a predetermined threshold, the control unit causes the first radiation detector to perform moving image shooting, and the dose determination unit detects a high-dose pulse greater than the threshold. If this happens, the second radiation detector is caused to execute still image shooting.

The first radiation detector preferably has a smaller pixel arrangement density than the second radiation detector. The first radiation detector preferably has fewer pixels than the second radiation detector.

Further, it is preferable that the first radiation detector has a higher frame rate than the second radiation detector. Further, the first radiation detector preferably has a smaller area than the second radiation detector.

The first radiation detector absorbs radiation to generate visible light, and is disposed on the radiation incident side of the light emitting unit, and detects the visible light generated by the light emitting unit. The second radiation detector is arranged on the side opposite to the radiation incident side of the light emitting unit and detects the visible light generated by the light emitting unit to generate image data. It is preferable that it is comprised by the 2nd photon detection part to do.

In this case, it is preferable that the light emitting portion has a columnar crystal phosphor, and the tip portion of the columnar crystal phosphor faces the first photodetection portion. The first light detection unit preferably has a smaller area than the light emitting unit.

The second light detection unit has a smaller area than the light emission unit, and condenses visible light emitted from the light emission unit on the second light detection unit between the light emission unit and the second light detection unit. It is preferable to further include a Fresnel lens.

The first radiation detector is disposed on the radiation incident side of the first light emitting unit and absorbs radiation and generates visible light, and visible light generated by the first light emitting unit. A first light detection unit configured to detect light and generate image data, and the second radiation detector absorbs the radiation transmitted through the first light emission unit and the first light detection unit and is visible. A second light emitting unit for generating light is disposed on the side opposite to the radiation incident side of the second light emitting unit, and visible light generated by the second light emitting unit is detected to generate image data. You may be comprised by the 2nd photon detection part.

In this case, it is preferable that one of the first light emitting part and the second light emitting part has a columnar crystal phosphor, and the other has a GOS phosphor or a BaFX phosphor.

According to the radiation imaging apparatus of the present invention, the first radiation detector that detects the radiation emitted from the radiation generator and generates image data, and the radiation that has passed through the first radiation detector are detected and imaged. And a second radiation detector for generating data, and the first radiation detector is used for moving image shooting and the second radiation detector is used for still image shooting. The moving mechanism of the vessel is not necessary. Further, according to the present invention, since the switching between the moving image shooting and the still image shooting is performed quickly, the shooting opportunity of the still image shooting is not missed.

It is a block diagram of a radiation information system. It is a side view which shows the arrangement | positioning state of each apparatus of a radiography system. It is a partially broken perspective view of an electronic cassette. It is sectional drawing which showed the structure of the electronic cassette typically. It is sectional drawing which shows the structure of a scintillator. It is explanatory drawing explaining the structure of a 1st radiation detector, a 2nd radiation detector, and a radiation dose measurement part. It is a block diagram which shows the electric constitution of an electronic cassette. It is a block diagram of a radiation generator and a console. It is explanatory drawing explaining the operation | movement timing of a radiography system. It is explanatory drawing explaining the 1st modification of an electronic cassette. It is explanatory drawing explaining the 2nd modification of an electronic cassette. It is explanatory drawing explaining the 3rd modification of an electronic cassette. It is explanatory drawing explaining the 4th modification of an electronic cassette. It is explanatory drawing explaining the 5th modification of an electronic cassette. It is explanatory drawing explaining the 6th modification of an electronic cassette. It is explanatory drawing explaining the 7th modification of an electronic cassette.

Referring to FIG. 1, a radiation information system (RIS) 10 is a system for managing information such as medical appointments and diagnosis records in a radiology department in a hospital. The RIS 10 includes a plurality of terminal devices 11, an RIS server 12, a radiography system 13 installed in each radiography room (or operating room) in the hospital, and an in-hospital network NW connected in a wired or wireless manner. It is comprised by.

As the terminal device 11, a personal computer (PC) or the like is used, which is operated by a photographer (doctor or radiographer). The photographer operates the terminal device 11 to input / view diagnostic information and facility reservation. A radiographic imaging request (imaging reservation) is also input via the terminal device 11.

The RIS server 12 is a computer including a storage unit 12A that stores an RIS database (DB). The storage unit 12A includes patient attribute information (patient name, sex, date of birth, age, blood type, patient ID, etc.), medical history, medical history, history of radiography, and data of radiographic images taken in the past. Other information related to the patient such as information related to the electronic cassette 15 included in each radiation imaging system 13 (identification number, model, size, sensitivity, usable imaging part, use start date, use frequency, etc.) Yes. The RIS server 12 is a process for managing the entire RIS 10 based on the information registered in the storage unit 12A (for example, a process for receiving an imaging request from each terminal device 11 and managing an imaging schedule of each radiation imaging system 13). )I do.

The radiation imaging system 13 captures a radiation image instructed from the RIS server 12 according to the operation of a doctor or a radiographer. The radiation imaging system 13 includes a radiation generator 14 that generates radiation, an electronic cassette 15 that detects radiation transmitted through an imaging region of a patient and generates a radiation image, a cradle 16 that charges the electronic cassette 15, and these And a console 17 for controlling the operation of each device. The electronic cassette 15 is a portable radiation imaging apparatus.

In FIG. 2, the radiation imaging room includes a radiation generator 14, a standing table 20 used for radiography of a standing patient 20 A (hereinafter referred to as standing imaging), and a supine patient 21 A. A supine table 21 used for photographing (hereinafter, referred to as supine photographing) is installed. The standing base 20 is provided with a cassette chamber 22 in which the electronic cassette 15 is mounted. When performing the standing position photographing, the electronic cassette 15 is loaded into the cassette chamber 22 of the standing base 20. When performing the supine position photographing, the electronic cassette 15 is loaded into the cassette chamber 23 of the supine position table 21.

In addition, in the radiography room, in order to enable standing radiography and supine radiography with one radiation generator 14, the radiation generator 14 is supported on the ceiling 26 while supporting the telescopic support 25. A moving mechanism 24 that moves two-dimensionally is provided. The support | pillar 25 is supporting the radiation generator 14 so that rotation of the surroundings of a horizontal axis (arrow A direction) and a vertical axis (arrow B direction) is possible.

The cradle 16 is formed with an accommodating portion 16A capable of accommodating the electronic cassette 15. The electronic cassette 15 is accommodated in the accommodating portion 16A when not in use, and the built-in battery is charged in this state. The electronic cassette 15 is taken out of the cradle 16 by the photographer at the time of radiographic image capturing, and is set on the holding unit 22 of the standing base 20 in the case of the standing position imaging, and the supine position base 21 in the case of the standing position imaging. Set in the cassette chamber 23 of the machine.

In FIG. 3, the electronic cassette 15 includes a housing 30, a radiation dose measurement sensor 31, a first light detection unit 32, a light emission unit 33, a second light detection unit 34, a base 35, and a storage case 36. Yes. The radiation dose measurement sensor 31, the first light detection unit 32, the light emission unit 33, and the second light detection unit 34 are stacked in the housing 30 in this order along the radiation incident direction, and each has an area. It has the same panel shape.

The housing 30 is made of a radiation transmissive material and has an overall shape of a rectangular parallelepiped. The housing 30 has a top plate 30A formed of a low radiation absorbing material such as carbon. The top plate 30A is irradiated with radiation that has passed through the imaging region of the patient. Portions other than the top plate 30A of the housing 30 are made of ABS resin or the like.

The top panel 30A is composed of a plurality of light emitting diodes (LEDs), and the operation state of the electronic cassette 15 such as the operation mode (eg, “ready state” or “data transmitting”) and the remaining battery capacity is indicated by the LED. A display unit 37 is provided for displaying by lighting. Note that the display unit 37 may be a display device configured by a light emitting element other than an LED, a liquid crystal display or an organic EL display that displays a state with characters or the like. Moreover, you may provide the display part 37 in parts other than the top plate 30A.

The storage case 35 is provided along one end side in the longitudinal direction of the top plate 30A. The storage case 35 stores a microcomputer (not shown) and a battery (not shown). The battery is chargeable and detachable. Various electronic circuits of the electronic cassette 15 including the radiation dose measuring sensor 31, the first light detection unit 32, and the second light detection unit 34 are operated by electric power supplied from the battery. In order to prevent these various electronic circuits from being damaged by radiation, a radiation shielding member (not shown) such as a lead plate is provided on the top plate 30 A side of the storage case 36.

In FIG. 4, the first light detection unit 32 is configured by forming a plurality of pixels 324 including a photoelectric conversion unit 321, a thin film transistor (TFT: Thin Film Transistor) 322, and a capacitor 323 over an insulating substrate 325. ing. The pixels 324 are arranged in a two-dimensional matrix.

The insulating substrate 325 and the layer on which the TFT 322 and the capacitor 323 are formed constitute a so-called TFT active matrix substrate (hereinafter referred to as a TFT substrate) 32A. The TFT 322 is made of amorphous silicon. The insulating substrate 325 is formed of a material having light transmissivity, such as a quartz substrate, a glass substrate, and a resin substrate, and having little radiation absorption.

The photoelectric conversion unit 321 includes a first electrode 321A and a second electrode 321B, and a photoelectric conversion film 321C disposed therebetween. The photoelectric conversion film 321C is formed of amorphous silicon, and absorbs visible light emitted from the light emitting unit 33 described later to generate charges. The photoelectric conversion unit 321 constitutes a PIN-type or MIS-type photodiode and is provided on the TFT substrate 32A. A planarizing layer 326 that covers the photoelectric conversion unit 321 is provided on the TFT substrate 32A. The planarization layer 326 is formed of silicon nitride, silicon oxide, or the like, and the surface opposite to the radiation incident side is planarized.

The second light detection unit 34 has the same configuration as that of the first light detection unit 32, and the pixels 344 including the photoelectric conversion unit 341, the TFT 342, and the capacitor 343 are arranged in a two-dimensional matrix on the insulating substrate 345. A plurality are formed.

The photoelectric conversion unit 341 includes a first electrode 341A and a second electrode 341B, and a photoelectric conversion film 341C disposed therebetween. In addition, a planarization layer 346 that covers the photoelectric conversion portion 341 is provided, and the planarization layer 346 has a plane on the radiation incident side that is planarized. The insulating substrate 345 and the layer on which the TFT 342 and the capacitor 343 are formed constitute the TFT substrate 34A.

In the second light detection unit 34, the configuration order of each part with respect to the radiation incident direction is opposite to the configuration order of each part of the first light detection unit 32. That is, the planarization layer 326 of the first light detection unit 32 and the planarization layer 346 of the second light detection unit 34 face each other, and the light emitting unit 33 is disposed therebetween. The light emitting unit 33 generates and emits visible light in response to the incidence of radiation. The second light detection unit 34 has substantially the same planar shape and area as the first light detection unit 32.

The planarizing layer 326 of the first light detection unit 32 and the light emitting unit 33 are bonded to each other by a light-transmitting adhesive layer 327. Similarly, the planarization layer 346 of the second light detection unit 34 and the light emitting unit 33 are bonded to each other with a light-transmitting adhesive layer 347. Further, the insulating substrate 345 of the second light detection unit 34 is bonded to the base 35 with an adhesive layer 348.

A radiation dose measuring sensor 31 is formed on the radiation incident side of the first light detection unit 32. In the radiation dose measurement sensor 31, a wiring layer 311, an insulating layer 312, a photoelectric conversion unit 313, and a protective layer 314 are sequentially formed on an insulating substrate 325. The wiring layer 311 is a layer in which a wiring 73 (see FIG. 7) described later is patterned on the insulating substrate 315. The photoelectric conversion unit 313 is an element that detects light emitted from the light emitting unit 33 and transmitted through the first light detection unit 32, and a plurality of photoelectric conversion units 313 are formed on the insulating layer 312 in a matrix. The thickness of the radiation dose measuring sensor 31 is about 0.05 mm.

The photoelectric conversion unit 313 includes a first electrode 313A and a second electrode 313B, and a photoelectric conversion film 313C disposed therebetween. The photoelectric conversion film 313C is formed of an organic photoelectric conversion material. The photoelectric conversion film 313C is formed by applying an organic photoelectric conversion material onto the second electrode 313B using an inkjet head or the like.

In FIG. 5, the light emitting unit 33 includes a vapor deposition substrate 331, a scintillator 332, and a moisture-proof protective film 333. The vapor deposition substrate 331 is a light transmissive substrate such as a quartz substrate, a glass substrate, or a resin substrate. The scintillator 332 is formed by vapor depositing thallium activated cesium iodide (CsI: Tl) on the vapor deposition substrate 331. The scintillator 332 includes a non-columnar crystal 332A and a plurality of columnar crystals 332B provided on the non-columnar crystal 332A. The moisture-proof protective film 333 is formed of a light-proof moisture-proof material (for example, polyparaxylylene) and covers the periphery of the scintillator 332.

Note that the vapor deposition substrate 331 is not necessarily provided. For example, after forming the scintillator 332 on the vapor deposition substrate 331, the vapor deposition substrate 331 may be peeled off from the scintillator 332, and the scintillator 332 may be bonded to the second light detection unit 34. Alternatively, the scintillator 332 may be directly deposited on the second light detection unit 34. Further, instead of CsI: Tl, a phosphor material such as sodium activated cesium iodide (CsI: Na) may be used.

In the scintillator 332, the tip portion 332 C of the columnar crystal 332 B is arranged to face the first light detection unit 32. The vapor deposition substrate 331 is bonded to the second light detection unit 34 with an adhesive or the like. The plurality of columnar crystals 332B are separated from each other through the gap GP. The diameter of each columnar crystal 332B is about several μm to 10 μm.

The scintillator 332 absorbs radiation that is emitted from the radiation generator 14 and passes through the patient, the top plate 30A, the radiation dose measurement sensor 31, the first light detection unit 32, and the like and is incident on the light emitting unit 33 to generate visible light. appear. Since radiation enters the scintillator 332 from the first light detection unit 32 side, light emission in the scintillator 332 occurs mainly on the distal end portion 332C side. Visible light generated in the scintillator 332 travels toward the first light detection unit 32 and the second light detection unit 34 by the light guide effect of the columnar crystal 332B.

The visible light that has traveled toward the first light detection unit 32 is emitted from the pointed tip 332C, passes through the moisture-proof protective film 333, and enters the first light detection unit 32. The first light detection unit It is detected by 32 photoelectric conversion units 321. Further, part of the visible light incident on the first light detection unit 32 passes through the first light detection unit 32 and enters the radiation dose measurement sensor 31. Visible light incident on the radiation dose measurement sensor 31 is detected by the photoelectric conversion unit 313.

On the other hand, visible light traveling toward the second light detection unit 34 is incident on the non-columnar crystal 332A and is partially reflected by the non-columnar crystal 332A. The light is incident on the light detector 34. Visible light incident on the second light detection unit 34 is detected by the photoelectric conversion unit 341.

As shown in FIG. 6, the light emitting unit 33 and the first light detection unit 32 constitute a first radiation detector 40. The first radiation detector 40 is arranged in the order of the first light detection unit 32 and the light emitting unit 33 along the radiation traveling direction. Such an arrangement method is called an ISS (Irradiation Side Sampling) type. The light emitting unit 33 and the second light detection unit 34 constitute a second radiation detector 41. The second radiation detector 41 is arranged in the order of the light emitting unit 33 and the second light detection unit 34 along the radiation traveling direction. Such an arrangement method is called a PSS (Penetration Side Sampling) type.

Further, the light emitting unit 33 and the radiation dose measuring sensor 31 constitute an ISS type radiation dose measuring unit 42. As described above, since the light emission in the scintillator 332 of the light emitting unit 33 occurs in the vicinity of the first light detection unit 32, the first radiation detector 40 has high sensitivity to radiation.

In the first radiation detector 40, the arrangement pitch of the pixels 324 is larger than the arrangement pitch of the pixels 344 of the second radiation detector 42 (the arrangement density is small), and the number of pixels 324 (the number of effective pixels) is small. For this reason, the first radiation detector 40 is driven at a higher frame rate than the second radiation detector 42 and is used for moving image shooting. The other second radiation detector 41 is used for still image shooting.

In FIG. 7, the first photodetecting portion 32 extends along the row direction, and includes a plurality of gate wirings 50 for turning on / off each TFT 322, and a column direction intersecting the row direction. A plurality of data wirings 51 are provided for reading out the charges accumulated in the capacitor 323 through the TFT 322 in the on state. The first radiation detector 40 is provided with a gate line driver 52, a signal processing unit 53, and an image memory 54 in addition to the first light detection unit 32.

The gate wiring 50 is connected to the gate line driver 52. The data wiring 51 is connected to the signal processing unit 53. When the electronic cassette 15 is irradiated with radiation that has passed through the imaging region of the patient (radiation carrying image information of the imaging region), visible light having a light amount corresponding to the radiation dose is emitted from the light emitting unit 33. . In the photoelectric conversion unit 321 of each pixel 324, a charge having a magnitude corresponding to the incident light amount of visible light is generated. This charge is accumulated in the capacitor 323.

When charges are accumulated in the capacitor 323, the TFTs 322 are sequentially turned on in units of rows by a signal supplied from the gate line driver 52 via the gate wiring 50. The electric charge accumulated in the capacitor 323 of the pixel 324 in which the TFT 322 is turned on is transmitted through the data wiring 51 as an analog electric signal and input to the signal processing unit 53. In this way, the charges accumulated in the capacitor 323 of each pixel 324 are sequentially read out in units of rows.

As is well known, the signal processing unit 53 includes an amplifier (not shown) and a sample hold circuit (not shown) for each data wiring 51. The electric signal transmitted through each data line 51 is amplified by an amplifier and then held in a sample and hold circuit. A multiplexer (not shown) and an A / D converter (not shown) are sequentially connected to the output side of the sample hold circuit. The electric signal held in each sample and hold circuit is selected by a multiplexer and converted into digital image data by an A / D converter. An image memory 54 is connected to the signal processing unit 53, and image data output from the A / D converter of the signal processing unit 53 is stored in the image memory 54.

Similarly, the second light detection unit 34 is provided with a plurality of gate wirings 60 and a plurality of data wirings 61. The second radiation detector 41 is provided with a gate line driver 62, a signal processing unit 63, and an image memory 64 in addition to the second light detection unit 34. The gate line 60 is connected to the gate line driver 62, and the data line 61 is connected to the signal processing unit 63. An image memory 64 is connected to the signal processing unit 63.

As described above, since the first light detection unit 32 has a low arrangement density of the pixels 324, the number of the gate wirings 50 and the data wirings 51 is equal to the number of the gate wirings 60 and the data wirings 61 of the second light detection unit 34. Less than the number. Further, since the first radiation detector 40 is for moving image shooting, the gain of the amplifier of the signal processing unit 53 is set to a value larger than the gain of the amplifier of the signal processing unit 63 in the second radiation detector 41. Yes. Since the configuration of the second radiation detector 41 other than this is the same as the configuration of the first radiation detector 40, detailed description thereof is omitted.

The image memories 54 and 64 are connected to a cassette control unit 70 that controls the overall operation of the electronic cassette 15. The cassette control unit 70 is a microcomputer, and includes a CPU 70A, a RAM 70B, and a nonvolatile ROM 70C such as a flash memory.

The cassette control unit 70 is connected to a wireless communication unit 71 that wirelessly transmits and receives various types of information to and from external devices. The wireless communication unit 71 corresponds to a wireless LAN (Local Area Network) standard represented by IEEE (Institute of Electrical and Electronics Electronics) (802.11a / b / g / n). The cassette control unit 70 performs wireless communication with the console 17 via the wireless communication unit 71.

The radiation dose measuring unit 42 is used for measuring the dose of radiation (radiation dose per unit time) irradiated to the electronic cassette 15 from the radiation generator 14. The radiation generator 14 emits, as radiation, a low-dose pulse for moving image shooting and a high-dose pulse for still image shooting according to the operation of the photographer.

The radiation dose measurement sensor 31 of the radiation dose measurement unit 42 is provided with the same number of wirings 73 as the photoelectric conversion unit 313. In addition to the radiation dose measurement sensor 31, the radiation dose measurement unit 42 is provided with a signal detection unit 74. Each photoelectric conversion unit 313 is connected to the signal detection unit 74 via a dedicated wiring 73. The signal detection unit 74 includes an amplifier, a sample hold circuit, and an A / D converter (all not shown) for each wiring 73, and is connected to the cassette control unit 70 and the dose determination unit 75.

The signal detection unit 74 performs sampling of a signal transmitted from the photoelectric conversion unit 313 via the wiring 73 in a predetermined cycle under the control of the cassette control unit 70, converts the sampled signal into digital data, and determines the dose. The data are sequentially output to the unit 75. The dose determination unit 75 determines the dose of radiation emitted from the radiation generator 14 based on the data input from the signal detection unit 74 (that is, a low-dose pulse for moving image shooting and a high-dose pulse for still image shooting). To determine which of the two). This determination result is output to the cassette control unit 70.

The electronic cassette 15 is provided with a power supply unit 77 and is connected to the various electronic circuits described above by wiring (not shown). The power supply unit 77 incorporates the above-described battery so as not to impair the portability of the electronic cassette 15, and supplies power from the battery to various electronic circuits. The power supply unit 77 is connected to the cassette control unit 70. The cassette controller 70 can selectively turn on / off the power supply to the first radiation detector 40 and the second radiation detector 41.

In FIG. 8, the console 17 is configured by a computer and includes a CPU 170 that controls the operation of the entire apparatus, a ROM 171 that stores various programs including a control program in advance, a RAM 172 that temporarily stores various data, And an HDD 173 for storing data, which are connected to each other via a bus line BL. In addition, a communication I / F 174 and a wireless communication unit 175 are connected to the bus line BL, and a display 176 is connected via a display driver 177. Furthermore, an operation unit 178 is connected to the bus line BL via an operation input detection unit 179.

The communication I / F 174 is connected to the connection terminal 14A of the radiation generator 14 via the connection terminal 17A and the communication cable 78. The CPU 170 transmits and receives information such as the exposure conditions to and from the radiation generator 14 by a wired method using the communication I / F 174 and the like. The wireless communication unit 175 communicates with the wireless communication unit 71 of the electronic cassette 15 and transmits and receives various types of information such as image data between the CPU 170 and the electronic cassette 15.

The display driver 177 generates and outputs a signal for displaying various information on the display 176. The CPU 170 displays an operation menu, a radiation image, and the like on the display 176 via the display driver 177. The operation unit 178 includes a keyboard and the like, and various information and operation instructions are input thereto. The operation input detection unit 179 detects an operation on the operation unit 178 and transmits a detection result to the CPU 170. The operation unit 178 is connected to a foot switch (not shown) that is arranged on the floor of the radiation imaging room and performs switching between moving image shooting and still image shooting. The foot switch is turned on / off when the photographer steps on the foot.

The radiation generator 14 performs radiation based on the radiation I / F 141 that transmits and receives various information such as the exposure conditions between the radiation source 140 that generates radiation and the console 17, and the exposure conditions received from the console 17. A radiation source controller 142 for controlling the source 140.

Next, the operation of the RIS 10 will be described. When taking a radiographic image, an imaging request is input from the terminal device 11. In this imaging request, a patient to be imaged, an imaging region to be imaged are designated, and tube voltage, tube current, etc. are designated as necessary.

1 notifies the RIS server 12 of the content of the input photographing request. The RIS server 12 stores the content of the imaging request notified from the terminal device 11 in the storage unit 12A. The console 17 accesses the RIS server 12 to acquire the content of the imaging request and the attribute information of the patient to be imaged, and displays the content of the imaging request and the attribute information of the patient on the display 176 (see FIG. 8). .

The radiographer (radiologist, etc.) performs preparatory work for radiographic imaging based on the content of the radiography request displayed on the display 176. For example, when photographing the affected part of the patient 21 A lying on the prone table 21, the electronic cassette 15 is loaded in the cassette chamber 23 of the prone table 21.

When the preparatory work is completed, the photographer performs an operation for notifying the completion of the preparatory work through the operation unit 178 of the console 17. Using this operation as a trigger, the console 17 sets the operation mode of the electronic cassette 15 to the ready state. When the electronic cassette 15 is in a ready state, the radiation control unit 70 drives the radiation dose measurement unit 42 and the dose determination unit 75 to irradiate the radiation pulse (a low-dose pulse for moving image shooting or a still image). A standby operation for detecting a high-dose pulse for imaging) is started. The console 17 notifies the photographer that the camera is ready to shoot by switching the display on the display 176.

The photographer who has confirmed this notification issues a shooting instruction via the operation unit 178. For example, in the case of still image shooting, the console 17 transmits an instruction signal instructing the start of exposure to the radiation generator 14. The radiation generator 14 emits a high-dose pulse for taking a still image from the radiation generator 14 with a tube voltage and a tube current corresponding to the exposure conditions received from the console 17.

The cassette control unit 70 of the electronic cassette 15, when detecting the high-dose pulse by the radiation dose measurement unit 42 and the dose determination unit 75, drives the second radiation detector 41 to perform an imaging operation, and performs the second radiation detector. The image data obtained by 41 is transmitted to the console 17 via the wireless communication unit 71. In the console 17, the input image data is displayed on the display 176 as a still image.

When performing cardiovascular system diagnosis and treatment, several doctors as a photographer form a team. This team has an auxiliary role in charge of operations such as adjusting the position of the prone position 21 on which the patient is placed and rotating the radiation generator 14 according to the imaging region of the patient. And a doctor who operates a catheter or guide wire inserted into a patient while observing a moving image (perspective image). Since both hands are closed to operate the catheter and guide wire, the doctor switches between moving image shooting and still image shooting using the above-described foot switch. Movie shooting is used for patient positioning and lesion search. Still image capturing is used to obtain a clearer radiation image of a lesion.

Next, with reference to a timing chart shown in FIG. 9, a case will be described in which video shooting is temporarily interrupted and still image shooting is performed. At the time of moving image shooting, the radiation generator 14 irradiates the patient with a low-dose pulse for moving image shooting at a predetermined interval. At this time, the radiation dose measurement unit 42 performs radiation sampling at an interval shorter than the irradiation interval of the low dose pulse. The dose determination unit 75 compares the radiation dose at the rise of the radiation detected by the radiation dose measurement unit 42 with a predetermined threshold value, and determines that the radiation dose (intensity) is lower than this threshold value as a low dose pulse. To do.

When the low-dose pulse is detected by the dose determination unit 75, the cassette control unit 70 drives the first radiation detector 40 in synchronization with the low-dose pulse to execute the moving image capturing operation MP. In this moving image shooting operation MP, all the gate wirings 50 are selected at once by the gate line driver 52, all the TFTs 322 are turned on, and the charges accumulated in the capacitor 323 are discarded (reset).

Next, all the gate wirings 50 are not selected, all the TFTs 322 are turned off, and the capacitor 323 is in a charge accumulation state. The photoelectric conversion unit 321 generates charges corresponding to the radiation transmitted through the imaging region of the patient and accumulates them in the capacitor 323. Then, after the low-dose pulse irradiation is completed, the gate wiring 50 is sequentially driven by the gate line driver 52, whereby the charges accumulated in the capacitor 323 are read out, and image data is generated by the signal processing unit 53. .

In this moving image shooting operation MP, the cassette control unit 70 stops the supply of the power supply voltage from the power supply unit 77 to each part of the second radiation detector 41 and turns it off (OFF). Thereby, the influence of the power supply noise on the reading operation of the first radiation detector 40 is reduced.

Each time a low-dose pulse is detected, a moving image capturing operation MP is performed, and image data is sequentially transmitted from the image memory 64 to the console 17 via the wireless communication unit 71. In the console 17, the input image data is displayed on the display 176 as a moving image.

When a still image shooting instruction is given by operating a foot switch or the like during this moving image shooting, a high-dose pulse for shooting a still image is emitted from the radiation generator 14 toward the patient. The dose of this high dose pulse is about 100 times that of the low dose pulse. The dose determination unit 75 compares the radiation dose at the time of rising of the radiation detected by the radiation dose measurement unit 42 with a predetermined threshold value, and determines that the dose is a high dose pulse when the radiation dose is larger than this threshold value.

When the high-dose pulse is detected by the dose determination unit 75, the cassette control unit 70 drives the second radiation detector 41 in synchronization with the high-dose pulse to execute the still image shooting operation SP. This still image shooting operation SP is the same as the moving image shooting operation MP, and image data is generated by the second radiation detector 41. This image data is transmitted to the console 17 via the wireless communication unit 71, and is displayed on the display 176 as a still image on the console 17. Note that this still image may be displayed on another display other than the display 176.

Also, in this still image shooting operation SP, the cassette control unit 70 stops the supply of the power supply voltage from the power supply unit 77 to each part of the first radiation detector 40 to be turned off (OFF). Thereby, since the moving image capturing operation MP is interrupted, the influence of power supply noise on the reading operation of the second radiation detector 41 is reduced.

As described above, since the second radiation detector 41 has a high arrangement density of the pixels 344, a high-definition still image can be obtained. In contrast, the first radiation detector 40 is driven at high speed and generates a moving image at a high frame rate because the arrangement density of the pixels 324 is small and the number of the pixels 324 is small.

The first radiation detector 40 and the second radiation detector 41 are stacked in the radiation traveling direction, and the second radiation detector 41 detects the radiation transmitted through the first radiation detector 40. Therefore, when switching from moving image shooting to still image shooting, there is no need to move the second radiation detector 41, and still image shooting is quickly switched.

The following is a modification of the electronic cassette. In the modification shown in FIG. 10, the radiation dose measurement sensor 31, the light emitting unit 33, and the second light detection unit 34 have the same configuration as in the above embodiment, but the first light detection unit 32 a includes the light emission unit 33 and the first light detection unit 32 a. The second embodiment is different from the above embodiment in that the area is smaller than that of the second light detection unit 34 (the visual field range is small). Even in this case, it is preferable that the arrangement pitch of the pixels 324 of the first light detection unit 32a is larger than the arrangement pitch of the pixels 344 of the second light detection unit 34 (the arrangement density is small).

When the first light detection unit 32a having such a small area is used, the shape (particularly the edge portion) may be reflected in the image data obtained by the second light detection unit 34. Since this reflection becomes a fixed pattern, the reflection may be removed by performing correction processing of the fixed pattern in the signal processing unit 63 or an external image processing apparatus (not shown).

In the modification shown in FIG. 11, a plurality of first light detection units 32a to 32c having a small area are laid down to have an area equivalent to that of the light emitting unit 33 and the second light detection unit 34. In this case, in the radiographic images obtained by the first light detection units 32a to 32c, a gap corresponding to the joint between the first light detection units 32a to 32c is generated. Processing is performed. Since this radiation image is used as a moving image, it has little influence on diagnosis.

In the modification shown in FIG. 12, the first light detection unit 32 a and the second light detection unit 34 a are smaller than the light emission unit 33 and have a smaller area, and between the second light detection unit 34 a and the light emission unit 33. A Fresnel lens 80 is disposed on the surface. The visible light emitted from the light emitting unit 33 in the direction of the second light detection unit 34a is collected by the Fresnel lens 80 and is incident on the second light detection unit 34a. A visual field range equivalent to that of the light emitting unit 33 can be detected.

As the second light detection unit 34a having a small area, it is possible to use a CMOS image sensor or a CCD image sensor configured based on a silicon substrate or a wide gap semiconductor substrate such as silicon carbide (SiC). . Since a SiC substrate is about 500 times more resistant to radiation than a silicon substrate, it is preferable to use a SiC substrate.

On the other hand, silicon substrates have low radiation resistance, and there is a risk of deterioration due to cumulative exposure to radiation. For this reason, when a CMOS image sensor or a CCD image sensor configured based on a silicon substrate is used as the second light detection unit 34a, the Fresnel lens 80 is made of radiation (X It is preferable to form with a glass material containing an element having a high attenuation effect. The Fresnel lens 80 is also suitable when the substrate of the second light detection unit 34a is a SiC substrate. It is also preferable to provide a light-transmitting member such as a glass plate containing an element such as lead, strontium, or barium between the light emitting unit 33 and the Fresnel lens 80 separately.

In the example illustrated in FIG. 12, the first light detection unit 32 a has a smaller area than the light emitting unit 33, but as illustrated in FIG. 6, the first light detection unit 32 having the same size as the light emitting unit 33. May be used.

In the modification shown in FIG. 13, the radiation dose measurement sensor 31, the first light detection unit 32, the first light emission unit 33A, the second light detection unit 34, and the second light emission unit are arranged along the radiation traveling direction. 33B are arranged in order. The first light emitting unit 33A and the second light emitting unit 33B have the same configuration as the light emitting unit 33 described above. In this modification, the first radiation detector 40 is an ISS type radiation detector composed of a first light detection unit 32 and a first light emitting unit 33A, and the second radiation detector 41 is: It is an ISS type radiation detector comprised by the 2nd light detection part 34 and the 2nd light emission part 33B. In this case, the light reflecting layer 81A is formed on the surface opposite to the radiation incident side of the first light emitting portion 33A, and the light reflecting layer 81B is formed on the surface opposite to the radiation incident side of the second light emitting portion 33B. It is preferable to form. The

light reflecting layers

81A and 81B are formed of a metal film such as aluminum.

In the modification shown in FIG. 14, the radiation dose measuring sensor 31, the first light emitting unit 33A, the first light detecting unit 32, the second light emitting unit 33B, and the second light detecting unit are arranged along the radiation traveling direction. 34 are arranged in order. In this modification, the first radiation detector 40 is a PSS type radiation detector composed of a first light detection unit 32 and a first light emitting unit 33A, and the second radiation detector 41 is This is a PSS type radiation detector composed of a second light detection unit 34 and a second light emission unit 33B. In this case, it is preferable that the light reflecting layer 81A is formed on the radiation incident side surface of the first light emitting unit 33A, and the light reflecting layer 81B is formed on the radiation incident side surface of the second light emitting unit 33B.

In the modification shown in FIG. 15, the radiation dose measurement sensor 31, the first light detection unit 32, the first light emission unit 33 A, the second light emission unit 33 B, and the second light detection unit are arranged along the radiation traveling direction. 34 are arranged in order. In this modification, in the configuration shown in FIG. 6, the light emitting unit 33 is configured by a first light emitting unit 33A and a second light emitting unit 33B. The first radiation detector 40 is an ISS type radiation detector composed of a first light emitting unit 33A and a first light detecting unit 32, and the second radiation detector 41 is a second light emitting unit. This is a PSS type radiation detector composed of 33B and the second light detection unit.

13 to 15, the first light emitting unit 33A and the second light emitting unit 33B may be formed of phosphors having different characteristics. For example, a columnar crystal phosphor having a columnar crystal structure such as CsI: Tl or CsI: Na is used for the second light emitting portion 33B, and a barium halide (BaFX (X = Br, X) is used for the first light emitting portion 33A. Cl, I)) phosphors are used. In this case, it is preferable that the tip of the columnar crystal phosphor is opposed to the second light detection unit 34.

Although the columnar crystal phosphor has high resolution and high performance, it is more expensive than the BaFX phosphor, so that the second radiation detector 41 for still image photography that requires high-quality photography is required. A columnar crystal phosphor is used for the light emitting unit 33B, and a BaFX phosphor is used for the first light emitting unit 33A of the first radiation detector 40 for moving image shooting that does not require high image quality. This can reduce costs without sacrificing the desired performance. In addition, the columnar crystal phosphor has a shock resistance that deteriorates as the thickness increases. However, in this configuration, the columnar crystal phosphor can be thinned, and thus the impact resistance is improved.

The first light emitting unit 33A absorbs radiation (X-rays) with lower energy than the second light emitting unit 33B having the columnar crystal phosphor by the BaFX phosphor. This configuration is effective when the tube voltage of the radiation source 140 is changed between moving image shooting and still image shooting to increase the tube voltage during still image shooting in order to obtain a high-contrast still image.

In each of the modifications shown in FIGS. 13 to 15, a gadolinium oxide (GOS) phosphor may be used for the first light emitting portion 33A and a columnar crystal phosphor may be used for the second light emitting portion 33B. In this case, the first light emitting unit 33A absorbs radiation (X-rays) having a relatively higher energy than the second light emitting unit 33B having the columnar crystal phosphor by the GOS phosphor. Since the columnar crystal phosphor such as CsI has a characteristic that the sensitivity gradually decreases due to the cumulative irradiation of radiation, the first light emitting unit 33A absorbs the radiation on the higher pressure side than the absorption energy of the columnar crystal phosphor. A decrease in sensitivity of the second light emitting unit 33B is suppressed.

In this case, in the modification shown in FIG. 15, the GOS phosphor is formed by applying or bonding to the first light detection unit 32. The columnar crystal phosphor is formed by vapor deposition or bonding to the second light detection unit 34. The vapor deposition of the columnar crystal phosphor includes direct vapor deposition and indirect vapor deposition. Indirect vapor deposition is a method in which a columnar crystal phosphor is vapor-deposited on a vapor deposition substrate, the columnar crystal phosphor is bonded to the second light detection unit 34, and then the vapor deposition substrate is peeled off. The columnar crystal phosphor and the GOS phosphor are bonded together by bonding or by pouching in a state where both are pressed. Alternatively, the columnar crystal phosphor may be directly or indirectly deposited on the GOS phosphor, and then the columnar crystal phosphor and the second light detection unit 34 may be bonded together.

In the modification shown in FIG. 16, the radiation dose measuring sensor 31, the first light detection unit 32, the second light detection unit 34, and the light emitting unit 33 are sequentially arranged along the radiation traveling direction. In this configuration, the first radiation detector 40 is an ISS type radiation detector composed of a first light detector 32 and a light emitter 33, and the second radiation detector 41 is a second light. This is an ISS type radiation detector composed of a detector 34 and a light emitter 33B. In this case, it is preferable to form the light reflecting layer 82 on the surface of the light emitting unit 33 opposite to the radiation incident side.

Moreover, in the said embodiment and each modification, in the 1st radiation detector 40, although the photoelectric converting film 321C of the 1st photon detection part 32 is comprised with the amorphous silicon, the photoelectric converting film 321C is made into organic photoelectric. You may comprise with the material containing conversion material. In this case, an absorption spectrum showing high absorption mainly in the visible light region is obtained, and the photoelectric conversion film 321C hardly absorbs electromagnetic waves other than visible light emitted from the scintillator 332. Thereby, the noise which generate | occur | produces because a radiation is absorbed by the photoelectric converting film 321C is suppressed.

The photoelectric conversion film 321C made of an organic photoelectric conversion material can be formed by attaching an organic photoelectric conversion material onto the TFT substrate 32A using a droplet discharge head such as an ink jet head, and is included in the TFT substrate 32A. The insulating substrate 325 is not required to have heat resistance. For this reason, the insulating substrate 325 can be made of a material other than glass.

When the photoelectric conversion film 321 C is made of an organic photoelectric conversion material, radiation is hardly absorbed by the photoelectric conversion film 321 C, and thus attenuation of radiation due to transmission through the first light detection unit 32 is suppressed. Therefore, it is preferable that the photoelectric conversion film 321C is made of an organic photoelectric conversion material when the first radiation detector 40 is an ISS type.

The organic photoelectric conversion material that constitutes the photoelectric conversion film 321C is preferably such that its absorption peak wavelength is closer to the emission peak wavelength of the scintillator 332 in order to absorb the visible light emitted from the scintillator 332 most efficiently. Ideally, the absorption peak wavelength of the organic photoelectric conversion material matches the emission peak wavelength of the scintillator 332, but if the difference between the two is small, the visible light emitted from the scintillator 332 can be sufficiently absorbed. is there. Specifically, the difference between the absorption peak wavelength of the organic photoelectric conversion material and the emission peak wavelength of the scintillator 332 is preferably within 10 nm, and more preferably within 5 nm.

Examples of organic photoelectric conversion materials that can satisfy these conditions include quinacridone organic compounds and phthalocyanine organic compounds. Since the absorption peak wavelength in the visible region of quinacridone is 560 nm, if quinacridone is used as the organic photoelectric conversion material and CsI: Tl is used as the material of the scintillator 332, the difference in the peak wavelengths can be made within 5 nm. The amount of charge generated in the scintillator 332 can be substantially maximized.

The photoelectric conversion film 321C preferably contains an organic p-type compound or an organic n-type compound. An organic p-type compound is a donor organic semiconductor typified by a hole-transporting organic compound and has a property of easily donating electrons. More specifically, the organic p-type compound is an organic compound having a smaller ionization potential when two organic materials are used in contact with each other. Any organic compound can be used as the donor organic semiconductor as long as it has an electron donating property. The organic n-type compound is an acceptor organic semiconductor mainly represented by an electron transporting organic compound, and has a property of easily accepting electrons. More specifically, the organic n-type compound is an organic compound having a higher electron affinity when two organic compounds are used in contact with each other. As the acceptor organic semiconductor, any organic compound can be used as long as it has an electron accepting property.

In addition, the photoelectric conversion unit 321 only needs to include at least the

electrodes

321A and 321B and the photoelectric conversion film 321C, but in order to suppress an increase in dark current, at least one of an electron blocking film and a hole blocking film is provided. It is preferable to provide both.

The active layer of the TFT 322 is preferably an amorphous oxide containing at least one of In, Ga, and Zn (for example, an In—O system), and at least two of In, Ga, and Zn are used. Amorphous oxides containing (for example, In—Zn—O, In—Ga—O, and Ga—Zn—O) are more preferable, and amorphous oxides including In, Ga, and Zn are particularly preferable. As the In—Ga—Zn—O-based amorphous oxide, an amorphous oxide whose composition in a crystalline state is represented by InGaO ₃ (ZnO) _m (m is a natural number of less than 6) is preferable. = 4 is more preferable.

Further, the active layer of the TFT 322 may be formed of an organic semiconductor material. In this case, examples of the organic semiconductor material include phthalocyanine compounds described in JP2009-212389A, pentacene, vanadyl phthalocyanine, and the like.

If the active layer of the TFT 322 is formed of an amorphous oxide or an organic semiconductor material, it does not absorb radiation such as X-rays, or even if it is absorbed, the amount of noise remains very small. The

Further, the active layer of the TFT 322 may be formed of carbon nanotubes. In this case, the switching speed of the TFT 322 is increased. Further, the degree of light absorption in the visible light region in the TFT 322 can be reduced. When the active layer is formed of carbon nanotubes, the performance of the TFT 322 is remarkably deteriorated just by mixing a very small amount of metallic impurities into the active layer. Therefore, the highly pure carbon nanotubes are separated and extracted by centrifugation or the like. Therefore, it must be used for forming the active layer.

The amorphous oxide and organic semiconductor material constituting the active layer of the TFT 322 and the organic photoelectric conversion material constituting the photoelectric conversion film 321C can all be formed at a low temperature. Therefore, the insulating substrate 325 is not limited to a substrate having high heat resistance such as a quartz substrate or a glass substrate, and a flexible substrate made of synthetic resin, aramid, or bionanofiber can be used. Specifically, flexible materials such as polyesters such as polyethylene terephthalate, polybutylene phthalate, and polyethylene naphthalate, polystyrene, polycarbonate, polyethersulfone, polyarylate, polyimide, polycycloolefin, norbornene resin, and poly (chlorotrifluoroethylene). A conductive substrate can be used. If such a flexible substrate made of a synthetic resin is used, the weight can be reduced. Note that the insulating substrate 325 includes an insulating layer for ensuring insulation, a gas barrier layer for preventing permeation of moisture and oxygen, an undercoat layer for improving flatness or adhesion to electrodes, and the like. May be provided.

The bio-nanofiber is a composite of a cellulose microfibril bundle (bacterial cellulose) produced by bacteria (Acetobacter Xylinum) and a transparent resin. The cellulose microfibril bundle has a width of 50 nm and a size of 1/10 of the visible light wavelength, and has high strength, high elasticity, and low thermal expansion. By impregnating and curing a transparent resin such as acrylic resin or epoxy resin in bacterial cellulose, a bio-nanofiber having a light transmittance of about 90% at a wavelength of 500 nm can be obtained while containing 60 to 70% of the fiber. Bionanofiber has a low coefficient of thermal expansion (3-7ppm) comparable to silicon crystals, and is as strong as steel (460MPa), highly elastic (30GPa), and flexible. Thinner.

In addition, you may comprise the 2nd light detection part 34 similarly to the 1st light detection part 32 comprised as mentioned above.

Further, in the above-described embodiment and each modification, the configurations of the first and second light detection units may be changed as appropriate. The first and second light detection units may each use a TFT substrate, a silicon substrate, or a SiC substrate. As described above, since the silicon substrate has low radiation resistance, at least one of the first and second light detection units is formed of the silicon substrate, and the light emitting unit is disposed on the radiation incident side, the It is preferable to provide an optical member having a radiation attenuation effect between the photodetector formed of a silicon substrate and the light emitting portion. Examples of the optical member include a radiation-absorbing glass plate containing an element such as lead, strontium, and barium.

Moreover, in the said embodiment and each modification, all of the 1st radiation detector 40, the 2nd radiation detector 41, and the radiation dose measurement part 42 convert a radiation into light with a scintillator, and make this light into an electric charge. Although it is an indirect conversion type radiation detector for conversion, it may be a direct conversion type radiation detector that converts radiation directly into electric charges by a photoconductive layer such as amorphous selenium.

Moreover, in the said embodiment and each modification, although the radiation dose measurement sensor 31 is arrange | positioned in the upstream of a radiation from the 1st photon detection part 32 and the 2nd photon detection part 34, it replaces with this, The radiation dose measuring sensor 31 may be disposed on the downstream side of the radiation from the first light detection unit 32 and the second light detection unit 34. Further, the radiation dose measuring sensor 31 may be incorporated in the first light detection unit 32 or the second light detection unit 34.

In the above embodiment and each modification, an electronic cassette is exemplified as the radiation imaging apparatus, but the present invention can be applied to a radiation detection apparatus such as a mammography apparatus instead of the electronic cassette.

DESCRIPTION OF SYMBOLS 15 Electronic cassette 31 Radiation

dose measurement sensor

32, 32a-32c 1st light detection part 33

Light emission part

34, 34a 2nd light detection part 40 1st radiation detector 41 2nd radiation detector 42 Radiation

dose measurement part

321, 341

Photoelectric conversion unit

324, 344 Pixel 332 Scintillator 332 A Non-columnar crystal 332 B Columnar crystal 332 C Tip portion 333 Moisture-proof protective film

Claims

A first radiation detector for detecting radiation emitted from the radiation generator and generating image data;
A second radiation detector that detects radiation transmitted through the first radiation detector and generates image data;
A control unit that causes the first radiation detector to perform moving image shooting and causes the second radiation detector to perform still image shooting;
A radiation imaging apparatus comprising:
A radiation dose measuring unit for measuring a dose of radiation pulses emitted from the radiation generator;
A dose determination unit that compares the dose measured by the radiation dose measurement unit with a predetermined threshold;
The control unit causes the first radiation detector to perform moving image capturing when the low-dose pulse smaller than a predetermined threshold is detected by the dose determination unit, and the dose determination unit causes a high value higher than the threshold. If a dose pulse is detected, cause the second radiation detector to perform still image capture;
The radiation imaging apparatus according to claim 1, wherein:
The radiation imaging apparatus according to claim 2, wherein the first radiation detector has a smaller pixel arrangement density than the second radiation detector.
The radiation imaging apparatus according to claim 3, wherein the first radiation detector has a smaller number of pixels than the second radiation detector.
The radiation imaging apparatus according to claim 4, wherein the first radiation detector has a higher frame rate than the second radiation detector.
The radiation imaging apparatus according to claim 2, wherein the first radiation detector has a smaller area than the second radiation detector.
The first radiation detector is disposed on the radiation incident side of the light emitting unit to absorb visible radiation and generates visible light, and detects the visible light generated by the light emitting unit. And a light detection unit of
The second radiation detector is disposed on the light emitting unit and on the side opposite to the radiation incident side of the light emitting unit, and detects the visible light generated by the light emitting unit to generate image data. 2 light detection units,
The radiation imaging apparatus according to any one of claims 1 to 6, wherein:
The radiographic imaging according to claim 7, wherein the light emitting unit has a columnar crystal phosphor, and a tip of the columnar crystal phosphor is opposed to the first light detection unit. apparatus.
The radiation imaging apparatus according to claim 7, wherein the first light detection unit has a smaller area than the light emitting unit.
The second light detection unit is smaller in area than the light emission unit, and visible light emitted from the light emission unit is transmitted between the light emission unit and the second light detection unit to the second light detection unit. The radiation imaging apparatus according to claim 9, further comprising a Fresnel lens for condensing the light.
The first radiation detector is disposed on a radiation incident side of the first light emitting unit and absorbs radiation and generates visible light, and is generated by the first light emitting unit. A first light detection unit that detects the visible light generated to generate image data,
The second radiation detector includes: a second light emitting unit that generates visible light by absorbing radiation transmitted through the first light emitting unit and the first light detecting unit; and the second light emitting unit. The second light detection unit is arranged on the side opposite to the radiation incident side and detects the visible light generated by the second light emitting unit to generate image data.
The radiation imaging apparatus according to any one of claims 1 to 6, wherein:
The radiographic imaging according to claim 11, wherein one of the first light emitting unit and the second light emitting unit has a columnar crystal phosphor, and the other has a GOS phosphor or a BaFX phosphor. apparatus.