JP2012045044A - Radiation image detection device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To accurately detect a start of irradiation even when there is simultaneously occurring system noise in the whole FPD (flat panel detector).SOLUTION: A detecting element part 41 comprises a first detecting element 44 and a second detecting element 46, and is provided in an imaging area. The first detecting element 44 accumulates signal charges corresponding to an incident amount of X-rays. The second detecting element 46 has a shielding member shielding visible light, has a configuration wherein visible light converted from X-rays by a scintillator does not enter a photodiode 48, and accumulates only dark charges. Voltages Vx, Vr of the first and second detecting elements 44, 46 are inputted to an operational amplifier 43. The operational amplifier 43 amplifies and outputs a difference between the voltages Vx, Vr. A control part detects that the irradiation is started based on an output voltage Vout from the operational amplifier 43.

Description

  The present invention relates to a radiological image detection apparatus that detects a radiographic image by receiving radiation that has passed through a subject.

  In the medical field, X-ray imaging systems using radiation, for example, X-rays, are known for performing image diagnosis. The X-ray imaging system includes an X-ray source that generates X-rays and an X-ray image detection device that receives an X-ray generated by the X-ray source and transmitted through a subject to detect an X-ray image. As an X-ray image detection apparatus, by using a TFT (Thin film transistor) active matrix substrate in which pixels that store signal charges according to the amount of incident X-rays are arranged, signal charges are stored for each pixel. An apparatus using an FPD (flat panel detector) that detects an X-ray image representing image information of a subject and outputs it as digital image data has been put into practical use.

  In the X-ray image detection apparatus, an FPD is provided on a standing imaging stand for photographing a subject (subject) in a standing posture and on a lying photographing stand for photographing a subject (subject) in a lying posture. In addition to the built-in stationary type, a portable X-ray image detection apparatus (hereinafter also referred to as “electronic cassette”) in which an FPD is built in a portable cassette type housing has been developed. The electronic cassette is used to image areas that are difficult to capture with the stationary type, and is carried along with a mobile X-ray source to the destination, and subjects who are difficult to move to the imaging room are laid on a bed in a hospital room Used to shoot in a state. Further, some electronic cassettes are known that can be used by being attached to an imaging stand used for a film cassette or an IP (imaging plate) cassette.

  In such an X-ray image detection apparatus, synchronous control for starting an accumulation operation of signal charges is performed in synchronization with a timing at which irradiation of the X-ray source is started by pressing an irradiation switch connected to the X-ray source. A control device such as a console for controlling the X-ray image detection apparatus is configured to synchronize the timing at which the X-ray source starts irradiation and the timing at which the X-ray image detection apparatus starts signal charge accumulation operation. When the irradiation is started, an irradiation start signal generated by the irradiation switch is received and output to the X-ray image detection apparatus as a synchronization signal. The X-ray image detection device periodically performs a reset operation for resetting the dark charges accumulated in the pixels regardless of whether or not irradiation is performed before starting the charge accumulation operation. The operation shifts from the operation to the signal charge accumulation operation to start photographing.

  However, when an X-ray image detection device and an X-ray source from different manufacturers constitute an imaging system, an interface for synchronization control (cable or connector interface) that is equipped as standard in the X-ray image detection device or its control device. Standards, synchronization signal formats, etc.) may not be compatible with the X-ray source interface. For this reason, a technique has been proposed in which the start of irradiation is detected by an X-ray image detection apparatus without using a synchronization signal to synchronize with an X-ray source (see Patent Documents 1 to 3).

  In the X-ray image detection apparatus described in Patent Document 1, an image reading operation is periodically performed at a predetermined frame rate before the start of irradiation is detected, and a difference between pixel values of two consecutive frames is examined. Thus, it is determined whether or not the difference is greater than or equal to a threshold value, and the start of irradiation is detected. When the irradiation is started, the pixel value rises due to the incidence of X-rays, so that the start of irradiation can be detected based on the difference between the pixel values between the two frames.

  The X-ray image detection apparatus described in Patent Document 2 detects the start of irradiation in a region where an X-ray irradiated from an X-ray source is incident on the FPD without passing through the subject (so-called blank region). The irradiation detection element is arranged, the rising of the output of the irradiation detection element is determined, and the start of irradiation is detected.

  The X-ray image detection apparatus described in Patent Document 3 sets a threshold value assuming a dark current output from an FPD pixel even when there is no X-ray irradiation, and sets a signal charge output from the pixel and the threshold value. In comparison, the start of irradiation is detected by determining whether or not the signal charge exceeds a threshold value.

JP 2003-126072 A Special Table 2002-543684 Publication JP 2008-507796 A

  However, in the detection methods of Patent Documents 1 to 3, if there is system noise that occurs simultaneously in the entire FPD, such as pixels and signal lines in the FPD, there is a risk of erroneous detection, and the start of irradiation cannot be detected accurately. There was a problem.

  The irradiation detection method of Patent Document 1 is a method of detecting the start of irradiation based on a difference in pixel values between two frames having a time difference in acquisition timing. For example, there is no system noise in the first frame and the second frame. In some cases, system noise occurs only in In such a case, when the pixel value increases due to system noise, it may be detected that irradiation is started.

  The irradiation detection method disclosed in Patent Document 2 is a method for determining the rising of the output of the irradiation detecting element arranged in the unexposed region. Therefore, when system noise occurs, the rising may be detected as the start of irradiation. is there.

  The irradiation detection method of Patent Document 3 detects the start of irradiation by comparing the threshold assuming a dark current with the signal charge output from the pixel, so that irradiation starts when system noise exceeding the assumed threshold occurs. May be detected.

  The present invention has been made in view of the above problems, and an object thereof is to occur simultaneously in the entire FPD in a radiological image detection apparatus having a function of detecting the start of irradiation regardless of a synchronization signal with a radiation source. Even in the presence of system noise, it is possible to accurately detect the start of irradiation.

  The radiological image detection apparatus of the present invention is an imaging unit for detecting a radiographic image of a subject by imaging radiation irradiated from a radiation source and transmitted through the subject, and accumulates signal charges corresponding to the incident amount of the radiation. An imaging unit having the radiation imaging region in which a plurality of pixels are arranged two-dimensionally, and a first driving unit that drives the pixel, and the darkness that occurs regardless of whether or not the radiation source is irradiated in the pixel A first drive means for performing three operations: a reset operation for resetting charge, an accumulation operation for accumulating the signal charge in the pixel, and a read operation for reading the signal charge from the pixel; and the radiation source Is an irradiation detection means for detecting the start of irradiation by the first detection element for accumulating the signal charge similarly to the pixel, and is arranged in the vicinity of the first detection element, A detection element unit having at least one set of detection elements composed of a second detection element that accumulates only the dark charge, and driving the first detection element and the second detection element to store each of them. Irradiation detecting means for detecting the start of the irradiation based on a difference in accumulated charge amount simultaneously read from the first detection element and the second detection element. And a control means for controlling the first drive means so that the pixel shifts from the reset operation to the accumulation operation when the irradiation detection means detects the start of the irradiation. Features.

  The irradiation detection means includes a differential amplifier that outputs a difference between analog signals output from the first detection element and the second detection element, and is based on the difference between the analog signals output from the differential amplifier. It is preferable to detect the start of irradiation.

  It is preferable that the first detection element and the second detection element are arranged in the imaging region.

  The first detection element and the pixel have the same structure, and the second detection element has a shielding member that shields radiation or visible light, except that the first detection element and the pixel have the shielding member. The first detection element and the pixel preferably have the same structure.

  The imaging means may be an indirect conversion type having a scintillator that converts the radiation into visible light. In this case, the pixels and the first and second detection elements are formed of photoelectric conversion elements that are formed on a substrate and convert visible light into electrical signals, and the scintillator It arrange | positions so that the whole surface of an arrangement | positioning surface may be opposed.

  When the imaging means is an indirect conversion type, the photoelectric conversion element includes, for example, a semiconductor layer, which is a photoelectric conversion layer, by a lower electrode disposed on the substrate side and an upper electrode disposed on the scintillator side. It is composed of a sandwiched layer structure. In this case, for example, the first detection element and the pixel are configured by the photoelectric conversion element in which the upper electrode is transparent, and the second detection element is configured by the photoelectric conversion element in which the upper electrode is opaque. The opaque upper electrode functions as the shielding member for shielding visible light from the scintillator toward the semiconductor layer.

  A plurality of the detection element units may be arranged in the imaging region. In that case, it is preferable that the detection element portions are arranged in different positions in the imaging region. The detection element unit may include a plurality of sets of the first detection element and the second detection element. The size of one detection element unit may be substantially the same as the size of one pixel. The detection element unit may be used for measuring an irradiation amount irradiated to a subject from the radiation source.

  The irradiation detection means may detect the stop of the irradiation based on the difference in the accumulated charge amount. In this case, when the stop is detected, the control unit shifts from the accumulation operation to the read operation. The radiographic image detection apparatus may be a portable electronic cassette.

  According to the present invention, similarly to the pixel, the first detection element that accumulates signal charges according to the amount of incident X-rays and the second detection that is arranged in the vicinity of the first detection element and accumulates only dark charges. Detecting the start of irradiation based on the difference in the amount of accumulated charge simultaneously read from the first detection element and the second detection element, using a detection element unit having at least one set of detection elements consisting of the detection elements Therefore, even when there is system noise that occurs simultaneously in the entire FPD, the start of irradiation can be accurately detected.

It is the schematic which shows the structure of a X-ray imaging system. It is explanatory drawing of FPD. It is a circuit diagram of an irradiation detection part. It is a timing chart which shows operation | movement of FPD at the time of imaging | photography. It is sectional drawing which shows the structure of the pixel part of an imaging panel. It is a top view which shows the structure of the pixel part of an imaging panel. It is sectional drawing which shows the structure of the 1st element part of an imaging panel. It is sectional drawing which shows the structure of the 2nd element part of an imaging panel. It is a top view which shows the structure of the 1st element part of an imaging panel. It is a top view which shows the structure of the 2nd element part of an imaging panel. It is a flowchart which shows the operation | movement at the time of imaging | photography. It is a circuit diagram of the irradiation detection part of a comparative example. It is a timing chart which shows the detail of detection operation. 14 is a timing chart for a case different from FIG. It is a timing chart of a case different from FIG. It is explanatory drawing of FPD with several element elements for a detection. It is a timing chart which shows operation in case there are a plurality of detection element parts. 18 is a timing chart of another example different from FIG. It is explanatory drawing of the element part for a detection which consists of a several element 1st element and 2nd element. It is explanatory drawing of FPD with the magnitude | size of one detection element part same as one pixel. It is a block diagram which shows the structure in the case of performing irradiation amount measurement. It is a block diagram which shows the structure in the case of performing dose management.

[First Embodiment]
In FIG. 1, an X-ray imaging system 10 includes an imaging table having a top plate 11 on which a subject H is placed, an X-ray source 12 that emits X-rays from an X-ray focal point 12 a toward the subject H, and a subject H. An electronic cassette 13 (radiation image detection device) that detects an X-ray image of the subject H upon receiving the transmitted X-rays is provided. The X-ray source 12 includes an X-ray tube that generates X-rays and an irradiation field limiter (collimator) that limits an X-ray irradiation field.

  The X-ray imaging system 10 includes a high voltage generation unit 14, an X-ray source control unit 15, a console 16, and a monitor 17. Imaging conditions such as tube voltage, tube current, and irradiation time are input to the X-ray source control unit 15 from the console 16 or an operation panel (not shown). The X-ray source control unit 15 sends the input imaging conditions to the high voltage generation unit 14. The X-ray source control unit 15 is connected to an irradiation switch 18 for inputting an irradiation start signal. The X-ray source control unit 15 outputs an irradiation start signal input from the irradiation switch 18 to a high voltage generation unit. 14 to the X-ray tube 12.

  The high voltage generator 14 generates a tube voltage and a tube current according to the imaging conditions input from the X-ray source controller 15, and gives the generated tube voltage and tube current to the X-ray source 12. When receiving the irradiation start signal, the X-ray source 12 starts X-ray irradiation according to a given tube voltage or tube current, and stops the irradiation when the irradiation time has elapsed.

  The console 16 is a control device that controls the electronic cassette 13. An irradiation start signal from the irradiation switch 18 is not input to the console 16. The console 16 transmits a control signal to the electronic cassette 13 via the communication unit 19 and receives X-ray image data detected by the electronic cassette 13. The monitor 17 displays an X-ray image received by the console 16 and displays an operation screen for operating the console 16.

  The electronic cassette 13 includes a FPD 21 that detects an X-ray image, a memory 22 that temporarily stores X-ray image data output from the FPD 21, and a console 16 in a flat, substantially rectangular parallelepiped housing. A communication unit 23 that performs communication of data and control signals in 22 is incorporated. For example, the communication unit 23 is a wireless communication unit that performs wireless communication using light such as radio waves and infrared rays. The electronic cassette 13 is a wireless type that incorporates a battery (not shown) that supplies power to each unit such as the FPD 21. . Of course, the communication unit 23 may be a wired communication unit that performs communication through a cable, or may receive power from a commercial power source through a power cable instead of a battery.

  2, the FPD 21 drives an imaging panel (imaging means) 61 (see FIG. 5) and a pixel 24 in which an imaging region 26 in which a plurality of pixels 24 that accumulate signal charges according to the amount of incident X-rays is arranged. A first gate driver 27 for controlling the reading of the signal charge, a signal processing circuit 28 for converting the signal charge read from the pixel 24 into digital data and outputting it, and a control unit 29 for controlling each part of the FPD 21. I have. The plurality of pixels 24 are two-dimensionally arranged in a matrix of n rows (x direction) × m columns (y direction) at a pixel pitch p. The size of one pixel 24 is, for example, about 150 μm.

  The FPD 21 has a scintillator 84 (see FIG. 5) that converts X-rays into visible light. The FPD 21 is an indirect conversion type that photoelectrically converts visible light converted by the scintillator 84 at the pixels 24. The scintillator 84 is disposed above the imaging region 26 so as to face the entire arrangement surface on which the pixels 24 are arranged.

  The pixel 24 includes a photodiode 31 that is a photoelectric conversion element that generates a charge upon incidence of visible light, a capacitor (not shown) that accumulates the charge generated by the photodiode 31, and a thin film transistor (TFT) 32 as a switching element. With. The TFT 32 has a gate electrode 63 (see FIG. 5) connected to the scanning line 33, a source electrode 66 (see FIG. 5) connected to the signal line 34, and a drain electrode 69 (see FIG. 5) connected to the photodiode 31. The The scanning lines 33 and the signal lines 34 are wired in a lattice pattern. The scanning lines 33 are the number of rows of the pixels 24 in the imaging region 26 (n rows), and the signal lines 34 are the number of columns of the pixels 24 (m Each column is wired. The scanning line 33 is connected to the first gate driver 27, and the signal line 34 is connected to the signal processing circuit 28.

  The first gate driver 27 resets dark charges generated in the pixels 24 regardless of whether X-rays are incident or not, and accumulates signal charges corresponding to the amount of incident X-rays in the pixels 24. , Driving means for performing three operations of reading out the signal charge from the pixel 24.

  While the first gate driver 27 turns off the TFT 32, the FPD 21 performs an accumulation operation for accumulating signal charges in the pixels 24. In the read operation, the first gate driver 27 sequentially generates gate pulses G1 to Gn, which are driving pulses for driving the TFT 32, sequentially activates the scanning lines 33 row by row, and connects to one scanning line 33. The TFTs 32 thus turned on are turned on line by line. The electric charge accumulated in the capacitor of the pixel 24 is read out to the signal line 34 and input to the signal processing circuit 28 when the TFT 32 is turned on.

  The signal processing circuit 28 includes an integration amplifier 36, a multiplexer (MUX) 37, and an A / D converter 38. The integrating amplifier 36 is individually connected to each signal line 34. The integrating amplifier 36 includes an operational amplifier and a capacitor connected between the input and output terminals of the operational amplifier, and the signal line 34 is connected to one input terminal of the operational amplifier. The other input terminal is connected to the ground (GND). The integrating amplifier 36 integrates the charges input from the signal line 34, converts them into voltage signals D1 to Dm, and outputs them. A common MUX 37 is connected to the output terminals of all the integrating amplifiers 36. An A / D converter 38 is connected to the output side of the MUX 37.

  The MUX 37 selects one integration amplifier 36 in order from a plurality of integration amplifiers 36 connected in parallel, and serially inputs voltage signals D1 to Dm output from the selected integration amplifier 36 to the A / D converter 38. . The A / D converter 38 converts the input voltage signals D <b> 1 to Dm into digital data and outputs the digital data to the memory 22. Thus, image data representing the X-ray image of the subject H is detected.

  The integrating amplifier 36 is provided with a reset switch 36a for resetting the integrated charge. The reset switch 36a is turned on by a reset pulse (reset signal) RST1 output from the control unit 29. The reset operation of dark charges generated in the pixel 24 is performed by the reset switch 36a as follows. The first gate driver 27 sequentially generates gate pulses G1 to Gn for the scanning line 33 to turn on the TFTs 32 of the pixels 24 row by row. While the TFT 32 is in the ON state, the dark charge accumulated in the pixel 24 flows to the integrating amplifier 36 through the signal line 34. The controller 29 outputs a reset pulse RST1 in synchronization with the gate pulses G1 to Gn, and resets the integrating amplifier 36 and the photodiode 31. Such a reset operation is periodically repeated until X-ray irradiation is started.

  The FPD 21 detects the irradiation state of the X-ray source 12, specifically, the start of X-ray irradiation by the X-ray source 12 (irradiation start) and the stop of irradiation (irradiation stop). An irradiation detection unit (irradiation detection means) is provided. The irradiation detection unit includes a detection element unit 41 provided in the imaging region 26, a second gate driver 42 that drives the detection element unit 41, and an operational amplifier 43 to which a signal from the detection element unit 41 is input. The control unit 29 determines the start and stop of X-ray irradiation based on the output signal from the operational amplifier 43.

  The detection element section 41 is composed of a set of detection elements including a first detection element (hereinafter referred to as a first element) 44 and a second detection element (hereinafter referred to as a second element) 46. The first element 44 is an element that accumulates signal charges corresponding to the amount of incident X-rays, like the pixel 24, and the second element 46 only accumulates dark charges that are generated regardless of whether or not X-rays are incident. It is an element to do. The first element 44 and the second element 46 are disposed adjacent to each other. The size (area occupied in the imaging region 26) and the arrangement pitch p of the first element 44 and the second element 46 are the same as those of the pixel 24, and the first element 44 and the second element 46 are within the imaging region 26. The pixel 24 is arranged in a partly replaced form.

  Since the detection element unit 41 does not function as the pixel 24, data corresponding to the position of the detection element unit 41 is missing in the X-ray image data output to the memory 22. The missing portion of data is interpolated based on the pixel value of the adjacent pixel, but the missing portion of data is preferably in an inconspicuous position. Since the detection element unit 41 only needs to be in a region where X-rays irradiated toward the subject H are incident, the detection element unit 41 is disposed near the outer periphery, avoiding the central region of the imaging region 26.

  As shown in FIG. 3, the first element 44 and the second element 46 are composed of photodiodes 47 and 48, a capacitor 50, and a TFT 49, as in the pixel 24. The photodiode 48 has a shielding film that shields visible light generated by the scintillator 84, and the second element 46 accumulates only dark charges by shielding the visible light incident by the shielding film.

  The second gate driver 42 is connected to the gate electrode 63 (see FIGS. 7 and 8) of the TFT 49 of each of the first element 44 and the second element 46 by the scanning line 54. The second gate driver 42 applies a driving gate pulse XG to the first element 44 and the second element 46 through the scanning line 54 to simultaneously turn on the TFTs 49 of the first element 44 and the second element 46. .

  The drain electrode 69 (see FIGS. 7 and 8) of each TFT 49 of the first element 44 and the second element 46 is connected to the photodiodes 47 and 48, respectively, and the source electrode 66 (see FIGS. 7 and 8) is respectively The output lines 51 and 52 are connected.

  Charges Cx and Cr generated by the photodiodes 47 and 48 are accumulated in the capacitors 50 of the first element 44 and the second element 46. When the TFT 49 is turned on, voltages Vx and Vr corresponding to the stored charges Cx and Cr of the first element 44 and the second element 46 are output to the two input terminals of the operational amplifier 43 through the output lines 51 and 52. . The voltage Vx is a voltage signal generated according to the accumulated charge Cx of the first element 44, and indicates a voltage value corresponding to the amount of incident X-rays. The voltage Vr is a voltage signal generated in accordance with the accumulated charge Cr of the second element 46, and indicates a voltage value corresponding to the dark charge accumulated in the second element 46. Since dark charges are also generated in the first element 44, the voltage Vx also includes a voltage value corresponding to the dark charges.

  The operational amplifier 43 is a differential amplifier that amplifies the difference between the two input voltages Vx and Vr and outputs the amplified difference value from the output terminal as an output voltage Vout that is an analog signal. The output voltage Vout represents the difference in the accumulated charge amount between the first element 44 and the second element 46.

  The output terminal of the operational amplifier 43 is connected to the control unit 29. The control unit 29 detects the start and stop of irradiation based on the output voltage Vout. The output lines 51 and 52 are connected to a reset switch 53 that resets the accumulated charges of the first element 44 and the second element 46 before the input terminal of the operational amplifier 43. The control unit 29 inputs a reset pulse (reset signal) RST <b> 2 to the reset switch 53 and turns on the reset switch 53. When the reset switch 53 is turned on, the accumulated charges Cx and Cr of the first element 44 and the second element 46 are discharged to the ground (GND).

  As shown in the timing chart of FIG. 4, the first gate driver 27 performs a reset operation until the irradiation detection unit detects the start of X-ray irradiation. During the reset operation, the first gate driver 27 sequentially outputs gate pulses G1 to Gn to the respective scanning lines 33 to turn on the TFTs 32 of the pixels 24. When the TFT 32 is turned on, the integrating amplifier 36 integrates the dark charge read from the pixel 24, and voltage signals D1 to Dm corresponding to the read dark charge are generated at the output terminal. During the reset operation, the MUX 37 does not select the integration amplifier 36, so that the voltage signals D 1 to Dm generated at the output terminal of the integration amplifier 36 are not output from the MUX 37.

  The control unit 29 generates the reset pulse RST1 so that the falling edges of the gate pulses G1 to Gn coincide with the falling edges of the reset pulse RST1, and the dark charges of the capacitor of the pixel 24 and the capacitor 36a of the integration amplifier 36 are generated. Reset. When a time T1 corresponding to the readout time for all the lines of the pixels 24 elapses, all the pixels 24 in the imaging region 26 are reset. Such a reset operation is repeated periodically.

  While the reset operation is being performed, the irradiation detection unit performs a detection operation for detecting the start of X-ray irradiation by the X-ray source 12. In the irradiation start detection operation, the second gate driver 42 generates a gate pulse XG at an interval of time T2 to turn on the TFT 49. When the TFT 49 is turned on, voltages Vx and Vr corresponding to the stored charges Cx and Cr are input to the operational amplifier 43. The control unit 29 reads the output voltage Vout of the operational amplifier 43 every time the voltages Vx and Vr are read to the operational amplifier 43. After reading the output voltage Vout, the control unit 29 generates the reset pulse RST2 so that the falling edge of the gate pulse XG coincides with the falling edge of the reset pulse RST2, and the accumulated charges of the first element 44 and the second element 46 are stored. Cx and Cr are reset.

  While the X-ray irradiation is not started, no signal charge corresponding to the amount of incident X-rays is generated in the first element 44. Therefore, the accumulated charges Cx and Cr are both at the output interval (T2) of the reset pulse RST2. It represents the dark charge accumulated between them. Therefore, during the reset operation, the voltages Vx and Vr have the same value, so that the output voltage Vout of the operational amplifier 43 is “0”.

  When irradiation of X-rays by the X-ray source 12 is started at time t01, signal charges corresponding to the amount of incident X-rays are generated in the first element 44. Since the signal charge is added to the dark charge to the accumulated charge Cx, the accumulated charge Cx increases when X-ray irradiation is started compared to before the start. When the TFT 49 is turned on, a voltage Vx corresponding to the accumulated charge Cx is input to the operational amplifier 43. On the other hand, since the photodiode 48 of the second element 46 is shielded, no signal charge is generated even if X-rays enter the FPD 21. Therefore, the accumulated charge Cr and the corresponding voltage Vr do not change before and after the start of X-ray irradiation.

  When the voltage Vx increases, a difference is generated between the voltage Vr and the output voltage Vout of the operational amplifier 43 increases. The control unit 29 reads the output voltage Vout of the operational amplifier 43 in synchronization with the generation timing of the gate pulse XG, and monitors whether the output voltage Vout has increased from “0”. When the voltage Vout increases, the start of X-ray irradiation is detected. When detecting the start of X-ray irradiation, the control unit 29 sends a control signal to the first gate driver 27 to stop the output of the gate pulses G1 to Gn. Thereby, the TFTs 32 of all the pixels 24 are turned off, and the pixels 24 shift from the reset operation to the accumulation operation.

  In the accumulation operation, signal charges are accumulated in the pixels 24 in accordance with the amount of incident X-rays. After detecting the irradiation start, the control unit 29 performs an irradiation stop detection operation. Even during the accumulation operation, the second gate driver 42 generates the gate pulse XG at intervals of time T2, periodically turns on the TFTs 49 of the first element 44 and the second element 46, and The voltages Vx and Vr of the first element 44 and the second element 46 are periodically input to the operational amplifier 43. The control unit 29 reads out the output voltage Vout of the operational amplifier 43 every time the voltages Vx and Vr are input, monitors whether the output voltage Vout has dropped to “0”, and detects the stop of irradiation. I do. After reading the output voltage Vout, the control unit 29 generates a reset pulse RST2 to reset the accumulated charges Cx and Cr.

  When the X-ray irradiation is stopped, the generation of the signal charge in the first element 44 is stopped, and the increase of the accumulated charge Cx is stopped. In this example, the accumulation period of the accumulated charge Cx corresponding to the voltage Vx read by the generation of the gate pulse XG immediately after the X-ray irradiation is stopped partially overlaps with the X-ray irradiation period. The charge Cx includes a signal charge generated before the X-ray irradiation is stopped. For this reason, there is a difference between the voltages Vx and Vr output to the operational amplifier 43 for the first time from the first element 44 and the second element 46 after the irradiation is stopped, so that the output voltage Vout of the operational amplifier 43 does not become “0”. .

  On the other hand, since the accumulation period of the accumulated charge Cx read by the next gate pulse XG does not overlap with the X-ray irradiation period, the second time from the first element 44 and the second element 46 after the irradiation is stopped. The difference between the voltages Vx and Vr read out is “0”, and the output voltage Vout of the operational amplifier 43 is also “0”. The controller 29 detects the irradiation stop at time t02 when the output voltage Vout becomes “0”.

  When the stop of irradiation is detected by the control unit 29, the control unit 29 shifts the pixel 24 from the accumulation operation to the read operation through the first gate driver 27. In the read operation, the first gate driver 27 sequentially generates gate pulses G1 to Gn, and the signal charges accumulated in the pixels 24 are read to the signal processing unit 28 through the signal line 34 row by row.

  When the read operation is started, the output of the gate pulse XG from the second gate driver 42 and the output of the reset pulse RST2 from the control unit 29 are stopped. Thereby, the irradiation detection operation for detecting the start and stop of irradiation is stopped until the reset operation is started again.

  As shown in FIGS. 5 and 6, in the imaging panel 61 of the FPD 21, the photodiode 31 and the TFT 32 are formed on the glass substrate 62 to form the pixel 24. The cross section shown in FIG. 5 is a cross section shown by the AA line of the plan view shown in FIG. On the glass substrate 62, first, a gate electrode 63 and a bridge wiring 64 are formed. The bridge wiring 64 is arranged to be orthogonal to the signal line 34 and the output lines 51 and 52, and is a wiring for obtaining a contact between one of the signal line 34 and the output lines 51 and 52 and the source electrode 66 of the TFT 32. . In FIG. 5 showing the configuration of the pixel 24, the bridge wiring 64 is connected to the signal line 34 through the contact hole 65.

  A gate insulating layer 67 is formed on the gate electrode 63 and the bridge wiring 64. A semiconductor active layer 68 is formed on the gate insulating layer 67 above the gate electrode 63. A source electrode 66 and a drain electrode 69 of the TFT 32 are formed on the gate insulating layer 67 so as to overlap a part of the semiconductor active layer 68. A protective film 71 is formed on the source electrode 66 and the drain electrode 69.

  A signal line 34 and output lines 51 and 52 are formed in the upper layer of the protective film 71. In the same layer as these, a connection line 73 for connecting the drain electrode 69 and the lower electrode 72 of the photodiode 31, and a bridge wiring 74 (see FIG. 6) and a connection line 75 are formed. The bridge wiring 74 is a wiring for obtaining a contact between the gate electrode 63 and one of the scanning lines 33 and 47. In FIG. 6 showing the planar configuration of the pixel 24, the gate electrode 63 is connected via the contact hole 74a. And the scanning line 33 are connected. The connection line 75 connects the source electrode 66 and the bridge wiring 64.

  An interlayer insulating film 76 is formed on the protective film 71, and the photodiode 31 is formed on the interlayer insulating film 76. The photodiode 31 has a semiconductor layer sandwiched between a lower electrode 72 and an upper electrode 77, and the semiconductor layer 78 includes a P-type semiconductor layer 78a, an I layer (intrinsic semiconductor layer) 78b, and an N-type semiconductor in order from the lower layer. This is a PIN photodiode in which the layer 78c is formed. The upper electrode 77 of the photodiode 31 is formed of a transparent electrode that transmits visible light.

  The lower electrode 72 of the photodiode 31, the drain electrode 69 of the TFT 32, and the interlayer insulating film 76 constitute a capacitor. An insulating film 81 is formed above the upper electrode 77, and a common wiring 82 for applying a bias voltage to the upper electrode 77 of each photodiode 31 is formed via the insulating film 81.

  After the photodiode 31 is formed, a scintillator 84 is formed above the protective layer 83 with the protective layer 83 interposed therebetween. The scintillator 84 is a phosphor made of cesium iodide (CSI), gadolinium oxysulfide (GOS), or the like, and converts X-rays into visible light. The scintillator 84 has a size that covers the entire surface of the imaging region 26 in which the pixels 24 are arranged. In the photodiode 31 of the pixel 24, since the upper electrode 77 is transparent, visible light emitted from the scintillator 84 enters the semiconductor layer 78.

  As shown in FIGS. 7 to 10, the first element 44 and the second element 46 constituting the detection element unit 41 have substantially the same configuration as the pixel 24. The same parts are denoted by the same reference numerals, description thereof is omitted, and differences will be mainly described.

  The first element 44 shown in FIGS. 7 and 9 has the same layer structure as that of the pixel 24. The difference is that the gate electrode 63 of the TFT 49 is replaced with the scanning line 54 via the contact hole 74b instead of the scanning line 33. And the source electrode 66 of the TFT 49 is connected to the output line 51 via the contact hole 65 instead of the signal line 34. The upper electrode 77 of the photodiode 47 of the first element 44 is also transparent like the pixel 24.

  The second element 46 shown in FIGS. 8 and 10 also has the same layer structure as the pixel 24 and the first element 44 except for the upper electrode 86 of the photodiode 48. The difference from the first element 44 is that the source electrode 66 of the TFT 49 is connected to the output line 52 via the contact hole 65 instead of the output line 51. The upper electrode 86 is made of an opaque material that does not transmit visible light. Visible light from the scintillator 84 does not enter the semiconductor layer 78 of the photodiode 48. Therefore, the second element 46 outputs only dark charges.

  Thus, since the pixel 24 and the first element 44 and the second element 46 have substantially the same structure, the first element 44 and the second element 46 can be formed simultaneously in the manufacturing process of the pixel 24. Manufacturing efficiency is good. Further, in the second element 46, since the upper electrode 86 and the shielding member that shields visible light are used together, the number of parts is not increased and the manufacturing efficiency is good as compared with the case where the shielding member is added separately.

  Hereinafter, the operation of the above configuration will be described with reference to the flowchart shown in FIG. When photographing the subject H with the electronic cassette 13, the position of the electronic cassette 13 and the photographing part of the subject H are positioned. Imaging conditions such as tube current, tube voltage, and irradiation time are set for the X-ray source control unit 15. Then, the electronic cassette 13 is activated to start a reset operation (step (S) 101). As shown in FIG. 4, in the reset operation, the first gate driver 27 sequentially outputs gate pulses G1 to Gn through the scanning line 33 to drive the TFTs 32 of the pixels 24 for each line, and according to the drive timing. The control unit 29 outputs a reset pulse RST1. As a result, the dark charge accumulated in the pixel 24 is reset.

  When the reset operation is started, the irradiation detection unit is also activated to start the irradiation start detection operation (S102). In the irradiation start detection operation, the second gate driver 42 outputs the gate pulse XG to the TFT 49 of the first element 44 and the second element 46 through the scanning line 54, and the voltage output by the first element 44 and the second element 46. Read Vx and Vr. The control unit 29 monitors the increase in the output voltage Vout of the operational amplifier 43. Until the start of irradiation is detected, the detection operation for starting irradiation is continued (N in S102).

  When the irradiation switch 18 is pressed, the X-ray source 12 starts X-ray irradiation with a tube current and a tube voltage corresponding to the imaging conditions. When the irradiation is started, the voltage Vx rises and a difference from the voltage Vr occurs, so that the output voltage Vout of the operational amplifier 43 rises. The control unit 29 detects the rise of the output voltage Vout and detects the start of irradiation (Y in S102).

  When detecting the start of irradiation, the control unit 29 stops the reset operation of the pixel 24 (S103) and starts the accumulation operation (S104). In the accumulation operation, X-rays are incident on the pixels 24 and signal charges corresponding thereto are accumulated. After the accumulation operation is started, the control unit 29 starts an irradiation stop detection operation (S105). In the X-ray irradiation period, the voltage Vx output from the first element 44 increases during the reset pulse RST2 generated at the interval of time T2. In the irradiation stop detection operation, the control unit 29 monitors that the output voltage Vout of the operational amplifier 43 drops to “0”.

  When the irradiation time set in the imaging conditions has elapsed and the irradiation of the X-ray source 12 is stopped, the incidence of X-rays on the first element 44 is also stopped. When the output voltage Vout of the operational amplifier 43 becomes “0”, the control unit 29 detects the irradiation stop (Y in S105). When the irradiation stop is detected, the control unit 29 stops the accumulation operation (S106) and starts the reading operation (S107). In the read operation, the accumulated charge in the pixel 24 is read to the integrating amplifier 36 through the signal line 34. The voltage signals D1 to Dm read through the MUX 37 are converted into digital image data by the A / D converter 38, and the image data is written into the memory 22. When reading of all the lines of the pixels 24 is completed (S108), the image data written in the memory 22 is transmitted to the console 16 by the communication unit 23.

  When the reading operation is started, the irradiation stop detection operation is also stopped. If there is a next shooting (Y in S109), the reset operation is started again. If there is no next shooting (N in S109), the power of the electronic cassette 13 is turned off to end the shooting.

  As in the present invention, the detection element unit 41 including a pair of the first element 44 and the second element 46 is used, and the voltage Vx output from the first element 44 and the dark charge output from the second element 46 are used. When the irradiation detection operation is performed based on the difference from the voltage Vr corresponding to, the irradiation detection is accurately performed even when there is a system noise generated in the FPD 21 as compared with the irradiation detection unit of the comparative example shown in FIG. be able to. Hereinafter, the reason will be described in more detail with reference to FIGS.

  The irradiation detection unit of the comparative example shown in FIG. 12 does not include the second element 46, inputs the reference voltage Vrf having a fixed voltage value to the operational amplifier 43, and outputs the voltage Vx output from the first element 44 and the reference voltage. 3 is different from the irradiation detection unit of the present invention shown in FIG. 3 in that irradiation detection is performed based on the difference from Vrf. The reference voltage Vrf is a fixed value reference voltage set assuming dark charges generated by the first element 44. Other points are the same as in the present invention.

  As shown in FIG. 13, when system noise occurs in the FPD 21, the system noise is superimposed on the accumulated charges Cx, Cr, voltages Vx, Vr of the first element 44 and the second element 46 of the FPD 21. System noise includes drift noise due to the dark current of TFT, large current fluctuation when a large current flows through a drive circuit such as a gate driver, and the capacitance of a capacitor when vibration or shock is applied to the case. It consists of noise caused by fluctuations, etc., and there are many noises generated in the same phase. System noise appears as high-frequency noise that rises steeply, low-frequency noise that fluctuates relatively slowly, or a mixture of both.

  In the case shown in FIG. 13, the X-ray irradiation start timing substantially coincides with the generation timing of the reset pulse RST2, and the accumulation period until the next reset pulse RST2 in which the first element 44 accumulates signal charges. Is the largest case. When the accumulation period in which the first element 44 accumulates signal charges is long, the amount of signal charges accumulated in the first element 44 increases, so that the voltage Vx output to the operational amplifier 43 also has a relatively large value.

  Therefore, the difference W between the voltage Vx and the voltage Vr corresponding to the output voltage Vout in the present invention shown in FIG. 13A is also equivalent to the voltage Vx and the reference corresponding to the output voltage Vout in the comparative example shown in FIG. The difference W from the voltage Vrf is also a relatively large value. If the output voltage Vout is large, the voltage Vr exceeding the reference voltage Vrf can be obtained even in the comparative example. Even when system noise occurs, the S / N ratio becomes large, so that relatively stable detection is possible. In the case shown in FIG. 13, in both cases of the present invention and the comparative example, the irradiation start is detected at the time t1 when the first output voltage Vout is read after the X-ray irradiation starts.

  In the case shown in FIG. 13, the timing at which X-ray irradiation stops and the timing at which the reset pulse RST2 is generated substantially coincide with each other, and signal charges are generated in the first element 44 almost simultaneously with the X-ray irradiation stop. Is stopped. For this reason, in both the present invention and the comparative example, the first output voltage Vout becomes “0”, and irradiation stop is detected at the time point t2 when the output voltage Vout is read.

  On the other hand, the cases shown in FIGS. 14 and 15 are cases in which the timing of starting and stopping the irradiation of X-rays does not coincide with the timing of generating the reset pulse RST2. If the X-ray irradiation start timing does not coincide with the timing at which the reset pulse RST2 is generated, the accumulation period until the next reset pulse RST2 in which the first element 44 accumulates signal charges is shortened. When the accumulation period is short, the difference W between the voltage Vx and the reference voltage Vrf, which corresponds to the output voltage Vout, decreases, and the S / N ratio decreases. When the S / N ratio decreases, detection becomes unstable and the possibility of erroneous detection due to system noise superposition increases. In addition, if the set value of the reference voltage Vrf is too high, there is a situation in which only a small voltage Vx below the reference voltage Vrf is generated despite the occurrence of signal charges, and detection becomes impossible.

  In the case of FIG. 14, the voltage Vx of the first element 44 at the time when the X-ray irradiation is started is smaller than that in the case of FIG. In addition, a system noise in which a low frequency and a high frequency are mixed immediately after the first gate pulse XG is output after the start of the X-ray irradiation. The increase in the voltage Vx of the element 44 is delayed. For this reason, in the comparative example shown in FIG. 14B, it takes time until the voltage Vx exceeds the reference voltage Vrf, and the start of irradiation is detected at time t4.

  In the comparative example, the stop of X-ray irradiation is detected at the same time t6 as in the present invention shown in FIG. 14A, but after the X-ray irradiation is stopped, the first gate pulse XG Since the voltage Vx is small at the time point t5 when is output, the output voltage Vout rises in small increments, and the output voltage Vout becomes unstable. If the output voltage Vout is unstable, erroneous detection may occur.

  On the other hand, in the present invention shown in FIG. 14A, even if the voltage Vx is small, the difference from the voltage Vr is used as the output voltage Vout, so that the voltage Vout rises immediately after the gate pulse XG is output. The irradiation start is detected at time t3 earlier than time t4. In the case of the present invention, since the difference between the first element 44 and the second element 46 is output, the in-phase system noise generated in the both elements is canceled out, so that there is no influence by the system noise. Also, immediately after the X-ray irradiation is stopped, the output voltage Vout does not become unstable as in the comparative example.

  In the case shown in FIG. 15, since the voltage Vx is small, in the comparative example shown in FIG. 15B, the voltage Vx rises at the time t7 immediately after the start of X-ray irradiation, but does not exceed the reference voltage Vrf. The irradiation start cannot be detected, and the irradiation start is detected at time t8 when the next gate pulse XG is output. On the other hand, in the present invention shown in FIG. 15A, the irradiation start is detected at time t7 even if the voltage Vx is small. In the case shown in FIG. 15, the detection of the irradiation stop is detected at the same time t8 in the present invention and the comparative example.

  As described above, the present invention performs accurate and stable detection even when the voltage Vr is small or there is system noise, as compared with the comparative example in which irradiation is detected based on the reference voltage Vrf that is a fixed value. be able to.

  Also, the dark charge changes depending on the TFT characteristics, and the TFT characteristics are easily affected by temperature fluctuations due to the temperature around the FPD 21 and the length of the driving time. For this reason, in the case of detecting irradiation with a fixed reference voltage Vrf, it is not possible to follow the fluctuation of dark charge, so there is a high risk of erroneous detection. On the other hand, the present invention takes the difference between the voltage Vr corresponding to the dark charge and the voltage Vx, so that the fluctuation of the dark charge can be absorbed, and accurate irradiation detection can be performed.

  In addition, since irradiation detection is performed based on the output voltage Vout of the operational amplifier 43, which is an analog signal, irradiation detection can be performed in a shorter time than when irradiation detection is performed based on data after A / D conversion. it can. Specifically, it takes a short time from the start of irradiation until the start of irradiation is detected. For this reason, since the time from the start of irradiation until the transition to the accumulation operation is shortened, useless exposure that does not contribute to accumulation of signal charges can be reduced. In addition, since the time from when the irradiation is stopped to when the irradiation is stopped is short, an extra accumulation operation time after the irradiation is stopped can be shortened. Since dark current is generated even during the accumulation operation, if the accumulation operation time is long, noise (so-called offset noise) due to the dark current increases and image quality deteriorates. By shortening the extra storage operation time, the influence of dark current on image quality can be minimized.

  In this example, the first element 44 and the second element 46 are described as being adjacent to each other. However, the first element 44 and the second element 46 may be disposed in the vicinity. The second element 46 removes dark charges caused by system noise included in the accumulated charge Cx of the first element 44, and performs irradiation detection based on the output Vout corresponding to the signal charge accumulated by the first element 44. belongs to. System noise is noise that is generated simultaneously throughout the entire imaging area 26, and it is considered that there is no significant change in waveform or size throughout the entire imaging area 26, but the interval between the second element 46 and the first element 44. If is too large, there may be a difference in waveform and size between the system noises generated in both. Therefore, it is preferable that the system noises are arranged in the vicinity where the same level of system noise is expected to occur.

  As a specific range of the vicinity position, when the size of one pixel is about 50 μm to about 200 μm, it is a range of an interval of several pixels to several hundred pixels, more preferably several pixels to It is preferably about several tens of pixels. Of course, it is most preferable that they are adjacent as in this example. This is because the dark charges accumulated in both the first element 44 and the second element 46 are considered to be approximately equal if they are adjacent to each other.

[Second Embodiment]
The FPD 91 shown in FIG. 16 is provided with a plurality of detection element units 41 in the imaging region 26. The FPD 21 is the same as the FPD 21 of the first embodiment except that there are a plurality of detection element portions 41. In the FPD 91, the plurality of detection element units 41 are arranged not only near the outer periphery of the imaging region 26 but also distributed throughout the imaging region 26, such as the central region of the imaging region 26. In the FPD 91, scanning lines 54 for driving the first element 44 and the second element 46 and output lines 51 and 52 for reading out the voltages Vx and Vr are added according to the number of the detection element units 41. Yes. The second gate driver 42 outputs gate pulses XG <b> 1 to XG <b> 3 to each scanning line 54 to selectively drive the plurality of detection element units 41.

  As shown in FIG. 17, the control unit 29 detects the start of irradiation when the output voltage Vout of the operational amplifier 43 rises when the voltages Vx and Vr of each detection element unit 41 are read. The irradiation stop detection operation is performed by sequentially reading out the voltages Vx and Vr of the plurality of detection element units 41 by the gate pulses XG1 to XG3 that are sequentially output to the voltages Vx and Vr of all the detection element units 41. When the output voltage Vout of the corresponding operational amplifier 43 becomes “0”, the irradiation stop is detected.

  As for the irradiation stop detection operation, as shown in FIG. 18, after the start of irradiation is detected, the outputs of the gate pulses XG1 to XG3 are switched from sequential output to simultaneous output. The voltages Vx and Vr may be read simultaneously. When the simultaneous reading is performed, as in the FPD 91, when the output lines 51 and 52 and the operational amplifier 43 are shared by the plurality of detection element units 41, the voltages of the plurality of detection element units 41 are used. A voltage obtained by adding Vx and Vr is input to the operational amplifier 43. Even in this case, when the X-ray irradiation is stopped, the signal charges are not accumulated in the first elements 44 of all the detection element portions 41, and the voltage Vx corresponding to the dark charges is output. It may be monitored that the voltage Vout becomes “0”. As shown in FIG. 18, by simultaneously reading a plurality of detection element units 41 and detecting the irradiation stop, the time from the irradiation stop to the irradiation stop detection timing is shortened compared to the example of FIG. 17. can do.

  Depending on the imaging region, a part of the imaging region 26 may be used without using the entire imaging region 26. For example, when imaging an elbow or knee, the size of the imaging region is smaller than that of the imaging region 26, so it is sufficient to set the X-ray irradiation field within that range. However, when the detection element unit 41 is provided only in the vicinity of the outer periphery of the imaging region 26 as in the FPD 21, in order to irradiate the detection element unit 41 with X-rays, a wide region with respect to the imaging region is X-rayed. It must be set as the irradiation field. Thus, it is not preferable to expand the X-ray irradiation field beyond the range necessary for imaging because it may cause unnecessary exposure.

  As in the FPD 91, if the detection element units 41 are distributed over the entire imaging area 26, even if the size or position of the irradiation field changes in the imaging area 26, any one of the detection elements 41 is used. Since there is a high possibility that the element unit 41 enters the irradiation field, it is not necessary to expand the irradiation field beyond the range necessary for imaging, and unnecessary exposure is reduced. The electronic cassette 13 is often used for imaging using a part of the imaging region 26, such as imaging of an elbow or knee, as compared to a stationary X-ray image detection apparatus. Therefore, the configuration in which the plurality of detection element portions 41 are distributed in the imaging region 26 is particularly effective for the electronic cassette 13.

  Further, the configuration in which the plurality of detection element portions 41 are arranged in a distributed manner is also effective when X-ray imaging is performed using a scattered X-ray removal grid plate. As is well known, the grid plate for removing scattered X-rays is formed by alternately arranging two kinds of materials such as a shielding material for shielding X-rays such as lead and a transmitting material for transmitting X-rays such as aluminum in a stripe pattern. The scattered X-rays formed by the transmission of the X-rays through the subject H are removed. At the time of shooting, the grid plate is disposed between the subject H and the FPD.

  When the grid plate is used, only the X-rays that have passed through the transmissive material are incident on the FPD imaging region 26. Therefore, in the imaging region 26, the X-rays are shielded by the shielding material, so that the X-rays are A shielded area is formed in which no incident or X-ray energy is reduced. In the case where there is only one detection element unit 41, if the detection element unit 41 is in the shielding region, the detection element unit 41 does not function, so that it may be impossible to perform irradiation detection. If a plurality of detection element units 41 are distributed in the imaging region 26, it is unlikely that all the detection element units 41 are located in the shielding region, and such a situation can be avoided.

  The X-ray absorption rate in the body varies depending on the site, such as a soft tissue having a low X-ray absorption rate and a bone portion having a high X-ray absorption rate. Therefore, for example, when imaging a portion having many bone parts, a shielding area may be generated in the imaging area 26 by the bone parts as in the case of the grid plate. A configuration in which a plurality of detection element units 41 are distributed in the imaging region 26 is also effective in such a case.

  In the FPD 21 according to the first embodiment, the detection element unit 41 is arranged near the outer periphery of the imaging region 26 because it is preferable that the number of missing pixels is small, but the irradiation field is not unnecessarily widened. Considering the viewpoint of reducing exposure and the viewpoint of avoiding inconvenience due to the generation of a shielding area due to photographing using a grid plate, a plurality of detection element units 41 are arranged in an imaging area like the FPD 91 of this embodiment. 26 is preferably distributed over the entire area. Even when a plurality of detection element units 41 are arranged in a distributed manner, the pixels lost by the detection element units 41 are subjected to interpolation processing using the pixel values of the surrounding pixels 24 as described in the first embodiment. Therefore, it is possible to reduce the influence of pixel loss on image quality by interpolation processing.

  As described above, the configuration in which the plurality of detection element units 41 are arranged in a distributed manner enables stable irradiation detection even when there are variations in use conditions such as the imaging region and the presence or absence of a grid plate. This configuration is the configuration of the detection element unit 41 (the first element 44 and the first element 44) that enable accurate and stable irradiation detection even when there is system noise that is caused by the temperature of the usage environment where the FPD is placed. Together with the combination of the two elements 46, it contributes to the realization of an irradiation detection function having a high resistance to fluctuations in the environment and conditions in which the FPD is actually used.

  Also, as described with reference to FIGS. 5 to 10, the configuration in which the pixel 24 and the first element 44 and the second element 46 have substantially the same structure increases the number of detection element portions 41 as in the FPD 91. The effect is greater. In addition, if the structure is the same, even in the case of a design change such as changing the position of the detection element unit 41 or increasing the number, it is possible to cope with relatively small changes such as changes in the number and position of scanning lines and output lines. Is possible.

[Third Embodiment]
As in the FPD 96 shown in FIG. 19, one detection element unit 97 may be composed of a plurality of sets (two sets in this example) of the first element 44 and the second element 46. If the first element 44 and the second element 46 are composed of a plurality of sets, even if one set does not function due to a failure or being located in the shielding region described in the second embodiment, the remaining Since the set can function to detect irradiation, the safety is high.

[Fourth Embodiment]
As in the FPD 101 shown in FIG. 20, the size of one detection element unit 97 is substantially the same as the size of one pixel 24, specifically, one detection element unit 97 occupies the imaging region 26. The area occupied by one pixel 24 may be substantially the same. The detection element unit 97 is composed of a set of detection elements of the first element 103 and the second element 104 that function in the same manner as the first element 44 and the second element 46. Since the detection element unit 97 is a missing part in the image data, the number of the detection element units 97 is preferably small. By making the size of one detection element unit 97 the same as that of the pixel 24, the missing portion of the image data can be reduced as compared with the FPD of the above embodiment having the size of two pixels. Can do.

  In each of the first to fourth embodiments, one detection element unit 41 is composed of the same number of first elements 44 and second elements 46 (one or two). However, for example, one second element 46 and a plurality of first elements 44 are configured such that one detection element unit 41 is constituted by one second element 46 and two first elements 44. One detection element unit 41 may be configured. In this case, one second element 46 is shared by a plurality of first elements 44 in one detection element unit 41. Specifically, when the voltage Vx of each of the plurality of first elements 44 is read to the operational amplifier 43, the voltage Vr of one shared second element 46 is read to the operational amplifier 43. Then, the difference between the voltage Vx of each first element 44 and the voltage Vr of one second element 46 is read as the output voltage Vout. According to such a configuration, since one second element 46 may be disposed for the plurality of first elements 44 that are dispersedly arranged, the second element 46 is compared with the case where the second element 46 is provided for each. The number can be reduced, and the missing portion of the image data can be reduced.

[Fifth Embodiment]
In the above-described embodiment, the example of performing the irradiation detection using the detection element unit has been described. However, the irradiation amount irradiated to the subject H may be measured using the detection element unit. For example, automatic exposure control can be performed by measuring the irradiation amount. In this case, as shown in FIG. 21, a dose calculation unit 106 and a determination unit 107 are provided in the control unit 29. The dose calculator 106 integrates the output voltage Vout from the operational amplifier 43 that is read each time the gate pulse XG is output, and measures the dose. The determination unit 107 stores an exposure value set before photographing, and compares the measured irradiation amount with the set exposure value to determine whether or not the irradiation amount has reached the exposure value. When the determination unit 107 determines that the irradiation amount has reached the exposure value, the control unit 29 causes the first gate driver 27 to output the gate pulses G1 to Gn and shifts from the accumulation operation to the read operation. Thereby, an X-ray image having an appropriate exposure value can be obtained regardless of the irradiation time of the X-ray source 12.

  When performing such automatic exposure control, a plurality of detection element portions 41 are arranged in the imaging region 26 as in the FPD 91 shown in FIG. It is preferable that the detection element unit 41 to be used can be selected depending on the imaging region. This is because, in the case of automatic exposure control, the area for measuring the dose varies depending on the imaging region.

  Moreover, you may use the element part 41 for a detection other than irradiation detection or automatic exposure control, for example for X-ray dose management. In recent years, in medical facilities such as hospitals, it has been required to manage an exposure history such as when and how much dose has been irradiated to which patient. The dose is determined by the product of the tube current and the irradiation time (so-called mAs value). In shooting, the shooting conditions including the irradiation time are set. In this case, the approximate dose can be grasped, but in actual shooting, the exposure set in the shooting conditions such as when automatic exposure control is performed. In many cases, the actual irradiation time is different from the time, and it is difficult to accurately grasp the dose actually irradiated.

  If the detection element unit 41 is used, the irradiation start and the irradiation stop can be accurately detected, so that the accurate dose used for imaging can be grasped based on the detection signal. For example, as shown in FIG. 22, when the electronic cassette 13 detects the start and stop of irradiation, a detection signal is sent to the console 16 through the communication unit 23 in real time. When the CPU 16a receives the irradiation start detection signal through the communication unit 19, the CPU 16a activates the timer 16b. Then, the timer 16b is stopped when the irradiation stop detection signal is received, and the irradiation time is measured.

  As described above, since the detection element unit 41 includes the first element 44 and the second element 46, accurate irradiation detection is possible, and irradiation detection is performed based on the output voltage Vout of the analog signal of the operational amplifier 43. Therefore, the time difference between the start and stop of irradiation and the detection timing thereof is small. For this reason, the irradiation time measured by the timer 16b is very close to the actual irradiation time. The CPU 16a obtains a dose by multiplying the measured irradiation time by a tube current and stores the dose data in a data storage 16c such as an HDD or a memory in association with the name or date of the patient who is the subject H. By using the detection element unit 41 in this way, the irradiation time can be accurately measured, so that accurate dose management is also possible.

  In the above embodiment, the example in which the detection element unit 41 is disposed in the imaging region 26 has been described. However, the detection element unit 41 may be disposed in a region where X-rays outside the imaging region 26 enter, such as the periphery of the imaging region 26. However, as described above, when arranged outside the imaging region 26, the X-ray irradiation field must be larger than the range necessary for imaging so that the X-rays are incident on the detection element unit 41. From the viewpoint of reducing exposure, it is preferable that the detection element unit 41 is disposed in the imaging region 26 as described in the above embodiment.

  In the above-described embodiment, the example in which both the irradiation start and the irradiation stop are detected has been described. However, the irradiation stop may not be detected, and only the irradiation start may be detected. This is because, as in the case of performing automatic exposure control, the storage operation may be shifted to the reading operation regardless of whether or not the irradiation is stopped.

  In the above embodiment, the electronic cassette which is a portable X-ray image detection apparatus has been described as an example. However, the present invention may be applied to a stationary X-ray image detection apparatus.

  In the above embodiment, the indirect conversion type has been described as an example of the FPD. However, the present invention may be applied to a direct conversion type FPD using a conversion layer (such as amorphous selenium) that directly converts X-rays into electric charges. In this case, a material such as lead that shields X-rays is used as the shielding member of the second element 46. The shielding lead is disposed at a position corresponding to the second element 46 on the upper surface of the conversion layer.

  Moreover, you may combine said each embodiment. In each of the above embodiments, X-rays have been described as an example of radiation. However, the present invention may use radiation other than X-rays, such as γ-rays. The present invention is not limited to the above-described embodiment, and various modifications can be made without departing from the spirit of the invention.

10 X-ray imaging system 13 Electronic cassette (radiation image detector)
21, 91, 96, 101 FPD
24 pixels 26 imaging area 27 first gate driver (first driving means)
29 Control unit (control means)
31, 47, 48 Photodiode 32, 49 TFT
41 element part for detection 42 second gate driver (second driving means)
43 operational amplifier 44 first element (first detection element)
46 Second element (second detection element)
53 Reset switch 61 Imaging panel (imaging means)

Claims (13)

An imaging means for detecting a radiation image of a subject by imaging radiation emitted from a radiation source and passing through the subject, and a plurality of pixels that accumulate signal charges according to the amount of incident radiation are arranged in two dimensions An imaging means having an imaging region of the emitted radiation;
A first driving means for driving the pixel, a reset operation for resetting a dark charge generated in the pixel regardless of the presence or absence of irradiation of the radiation source, an accumulation operation for accumulating the signal charge in the pixel, First driving means for performing three operations of a reading operation of reading out the signal charge from the pixel;
Irradiation detection means for detecting the start of irradiation by the radiation source;
Similar to the pixel, at least one set of detection elements including a first detection element that accumulates the signal charge and a second detection element that is disposed in the vicinity of the first detection element and accumulates only the dark charge. A detecting element portion having an element,
Second driving means for driving the first detection element and the second detection element to simultaneously read out the respective accumulated charges;
An irradiation detection means for detecting the start of the irradiation based on a difference in the amount of accumulated charge simultaneously read from the first detection element and the second detection element;
Control means for controlling the first driving means so that the pixel shifts from the reset operation to the accumulation operation when the irradiation detection means detects the start of the irradiation. Radiation image detection device.   The irradiation detection means includes a differential amplifier that outputs a difference between analog signals output from the first detection element and the second detection element, and is based on the difference between the analog signals output from the differential amplifier. The radiation image detection apparatus according to claim 1, wherein the start of irradiation is detected.   The radiological image detection apparatus according to claim 1, wherein the first detection element and the second detection element are arranged in the imaging region. The first detection element and the pixel have the same structure,
The second detection element has a shielding member that shields radiation or visible light, and has the same structure as the first detection element and the pixel except that the second detection element has the shielding member. The radiographic image detection apparatus according to claim 3. The imaging means includes a scintillator that converts the radiation into visible light, and the pixels and the first and second detection elements are photoelectric elements that are formed on a substrate and convert visible light into an electrical signal. It consists of conversion elements,
The radiographic image detection apparatus according to claim 4, wherein the scintillator is disposed so as to face the entire array surface of the photoelectric conversion elements. The photoelectric conversion element has a layer structure in which a semiconductor layer which is a photoelectric conversion layer is sandwiched between a lower electrode disposed on the substrate side and an upper electrode disposed on the scintillator side,
The first detection element and the pixel are configured by the photoelectric conversion element in which the upper electrode is transparent, and the second detection element is configured by the photoelectric conversion element in which the upper electrode is opaque. The radiographic image detection apparatus according to claim 5, wherein the opaque upper electrode functions as the shielding member for shielding visible light from the scintillator toward the semiconductor layer.   The radiographic image detection apparatus according to claim 3, wherein a plurality of the detection element units are arranged in the imaging region.   The radiological image detection apparatus according to claim 7, wherein the detection element units are distributed at different positions in the imaging region.   The radiological image detection apparatus according to claim 3, wherein the detection element unit includes a plurality of sets of the first detection element and the second detection element.   10. The radiological image detection apparatus according to claim 3, wherein the size of one detection element unit is substantially the same as the size of one pixel. 11.   11. The radiological image detection apparatus according to claim 3, wherein the detection element unit is also used for measuring an irradiation amount irradiated to a subject from the radiation source.   The irradiation detection unit detects the stop of the irradiation based on the difference in the accumulated charge amount, and the control unit shifts from the accumulation operation to the read operation when the stop is detected. The radiographic image detection apparatus of any one of Claims 1-11 to do.   The image pickup unit, the first drive unit, the second drive unit, the irradiation detection unit, and the control unit are electronic cassettes housed in a portable housing. The radiographic image detection apparatus of any one of Claims.

Images