JP6068700B2 - Radiation image detection apparatus and operation method thereof - Google Patents

Description

  The present invention relates to a radiological image detection apparatus and an operation method thereof.

  In the medical field, X-ray imaging systems using radiation, such as X-rays, are known. The X-ray imaging system includes an X-ray generation apparatus that generates X-rays and an X-ray imaging apparatus that captures an X-ray image formed by X-rays transmitted through a subject (patient). The X-ray generation apparatus inputs an X-ray source that irradiates an X-ray toward a subject, a radiation source control apparatus that controls driving of the X-ray source, and an instruction for operating the X-ray source to the radiation source control apparatus. It has an irradiation switch. The X-ray imaging apparatus has an X-ray image detection apparatus that receives an X-ray transmitted through a subject and detects an X-ray image, and a console that performs drive control of the X-ray image detection apparatus and storage and display of the X-ray image. ing.

  An X-ray image detection apparatus using a flat panel detector (FPD) as an image detection unit has become widespread. The FPD has an imaging region in which a plurality of pixels that accumulate signal charges according to the amount of incident X-rays are arranged. The pixel includes a photoelectric conversion unit that generates and accumulates charges, and a switching element such as a TFT. The FPD sequentially reads out the signal charges stored in the photoelectric conversion unit of each pixel to the signal processing circuit one row at a time through the signal line provided for each pixel column in response to the ON operation of the switching element. An X-ray image is electrically detected by converting the signal charge into a voltage signal.

  The signal processing circuit includes an integrating amplifier, a CDS circuit, an A / D converter, and the like. One integrating amplifier is provided for each column of signal lines, and integrates signal charges input from the signal lines to convert them into analog voltage signals. One CDS circuit is also provided for each signal line in each column, and is connected to the output terminal of each integrating amplifier. The CDS circuit has a sample-and-hold circuit that performs correlated double sampling on the analog voltage signal converted by the integrating amplifier to remove noise, and holds the analog voltage signal for a predetermined period of time (sample-and-hold). To do. The A / D converter converts the analog voltage signal held in the sample hold circuit into a digital voltage signal and outputs it to the memory. This memory is a frame memory capable of recording a digital voltage signal representing one X-ray image. In addition to this, the signal processing circuit further selects an amplifier for amplifying an analog voltage signal and a sample hold circuit of the CDS circuit of each column by an electronic switch, and converts the analog voltage signal of the sample hold circuit to an A / D converter. A multiplexer is provided for selectively supplying to.

  Patent Document 1 discloses an integration amplifier (readout partial circuit), an A / D converter (conversion partial circuit), and first and second circuits arranged between the A / D converter and a memory (collection subsystem). A signal processing circuit including a buffer memory (line buffer) is described. The first and second buffer memories are line memories capable of recording a digital voltage signal representing an X-ray image for one row of array pixels. An image reading operation for reading out an image for one screen using this signal processing circuit is performed as shown in FIG. In this signal processing circuit, as shown in FIG. 16, the signal of a pixel in a row that is integrated by an integration amplifier in one cycle (normal cycle) in which a digital voltage signal for one row is read, for example, in the Nth cycle. The image signal P (N), which is an analog voltage signal obtained by converting the charge p (N), is sampled and held by the CDS circuit immediately before the next cycle, and the sampled and held image signal P (N) is A in the next cycle. The digital signal is converted by the / D converter into the digital image signal Pd (N) and temporarily stored in the first buffer memory (see the AD Data row), and the image signal Pd (N) is read from the first buffer memory in the next cycle. A so-called pipeline process is executed to output (refer to the Data Out line). In one cycle, the digital conversion of the analog image signal P by the A / D converter and temporary storage (signal input) in one buffer memory, and the digital image temporarily stored in the other buffer memory in the previous cycle The signal Pd (N-1) is simultaneously output to the memory (signal output).

  When a voltage signal based on the charge of a pixel is read through the signal processing circuit, if there is only one buffer memory, the voltage of the next row must be output unless the voltage signal for one row temporarily stored in the buffer memory is output to the memory. The signal cannot be temporarily stored in the buffer memory. That is, the read time of one cycle from the start of reading of the analog voltage signal from the integrating amplifier to the writing of the digital voltage signal to the memory always takes the time of signal input to the buffer memory + signal output from the buffer memory.

  On the other hand, in the case of a pipeline-type signal processing circuit as described in Patent Document 1, since it has two first and second buffer memories that hold signals for two times that are read continuously from the integrating amplifier, In this cycle, the signal output to the memory of the voltage signal temporarily stored in one buffer memory and the signal input to the other buffer memory of the voltage signal sampled and held in the CDS circuit in this cycle can be processed in parallel. That is, since signal input for one row and signal output for the previous row are performed simultaneously in one cycle, the readout time for one screen (frame) is about half that of a single buffer memory. That's it. However, since the signal input and the signal output of a voltage signal in a certain row are shifted by one cycle, the signal charge p in a certain row is integrated to obtain the analog image signal P, and then the digital image signal Pd is output to the memory. The interval between the cycle times T is vacant.

  The Sync signal is sampled and held in each part of the signal processing circuit, such as sample hold of the analog image signal P, A / D conversion into the digital image signal Pd, temporary storage of the digital image signal Pd in the buffer memory, and output to the memory. Synchronous control signal for cycle execution, Internal Reset signal selects integration amplifier charge reset, CDS circuit sample hold state reset, or shift register for operating multiplexer electronic switch to select CDS circuit selection state A signal for executing at least one of selective resets to be returned at the start, the Analog Clock signal is a signal for determining operation control timing of an analog signal processing circuit (analog front end) including an integration amplifier and a CDS circuit, specifically, Integrating amplifier The reference clock signal that defines the timing of charge integration, the output of the voltage signal to the CDS circuit, the sample hold, and the like. The ADC Clock signal and the Buffer Data Clock signal are operation control signals for the A / D converter and the buffer memory, respectively.

  The operation buffer indicates a buffer memory in which the image signal Pd digitally converted by the A / D converter in the cycle is written out of the first and second buffer memories. The first and second buffer memories are alternately switched to the operation buffer. For example, the image signal Pd (N) is temporarily stored in the first buffer memory, and the image signal Pd (N + 1) of the next row is temporarily stored in the second buffer memory.

  By the way, in the X-ray image detection apparatus using the FPD, in order to minimize the influence of dark charge noise on the X-ray image, a reset operation for sweeping out unnecessary accumulated charge of the pixel due to dark current, residual charge of the previous imaging, or the like. Is regularly conducted by FPD. Therefore, in general, in an X-ray imaging system including an X-ray image detection apparatus using FPD, the X-ray irradiation start timing and the accumulation operation in which the FPD finishes the reset operation and accumulates signal charges in the pixels are started. Synchronized with timing. For example, an interface (I / F) capable of mutual communication is provided between the radiation source control device and the X-ray image detection device, and a synchronization signal is sent from the radiation source control device to the X-ray image detection device in accordance with the timing of starting X-ray irradiation. There is a system in which the X-ray image detection apparatus performs processing for shifting to an accumulation operation triggered by reception of a synchronization signal.

  Alternatively, the X-ray image detection apparatus is not connected to the radiation source control apparatus and does not exchange synchronization signals. Instead, the X-ray dose is detected by the dose detection sensor, and the detected dose is preset. Some devices have an irradiation start detection function that compares the irradiation start threshold values, determines that X-ray irradiation has started when the dose exceeds the irradiation start threshold value, and starts the accumulation operation in the FPD. Similarly, it has an irradiation end detection function that determines the end of X-ray irradiation by comparing the dose detected by the dose detection sensor and the irradiation end threshold, and shifts the FPD operation from the accumulation operation to the readout operation when it is determined that the irradiation has ended. There is also.

  In addition, in the X-ray imaging system, in order to obtain an X-ray image with appropriate image quality while suppressing the exposure to the subject, the X-ray dose is detected by a dose detection sensor during X-ray imaging (during irradiation). Then, when the integrated value (cumulative dose) of the dose detected by the dose detection sensor reaches the preset target dose, the X-ray irradiation by the X-ray source is stopped, and the X-ray image detection apparatus reads out from the accumulation operation. In some cases, automatic exposure control (AEC) to be shifted to is performed. The dose irradiated by the X-ray source is determined by the tube current time product (mAs value) which is the product of the X-ray irradiation time and the tube current defining the dose irradiated by the X-ray source per unit time. Although there are approximate recommended values for the imaging conditions such as irradiation time and tube current depending on the subject's imaging site (chest and head), gender, age, etc., the X-ray transmittance changes depending on individual differences such as the physique of the subject. AEC is performed to obtain a more appropriate image quality.

  Conventionally, an ion chamber or the like has been used as a dose detection sensor, but recently, a technique has been proposed in which an FPD pixel is simply modified to operate as a dose detection sensor. In Patent Document 2, a part of pixels (hereinafter referred to as detection pixels) is connected to a radiation detection wiring instead of a signal line without a switching element, and the charge generated in the detection pixel regardless of the on / off operation of the switching element. Flows out to the radiation detection wiring. Then, in the signal processing circuit connected to the radiation detection wiring, a voltage signal corresponding to the charge generated in the detection pixel (hereinafter referred to as a dose detection signal) is sampled at a predetermined period and input to the control unit, The control unit detects the start or end of AEC or X-ray irradiation based on the dose detection signal.

JP 2004-000564 A JP 2011-174908 A

  In Patent Document 2, a dedicated radiation detection wiring different from the signal line is connected to a detection pixel which is a dose detection sensor, and a signal processing circuit dedicated to dose detection is further connected to the wiring. It has been proposed to extract a dose detection signal from a normal pixel signal line or signal processing circuit without providing the above configuration. A plurality of detection pixels may be connected to one signal line. In one sampling in the dose detection operation, the charges of a plurality of detection pixels are collectively detected for each signal line.

In this way, when the dose detection signal is extracted from the signal line for normal pixels and the signal processing circuit and the pipeline type signal processing circuit of Patent Document 1 is used, the control timing of X-ray irradiation stop by AEC and X Judgment of the start or end of X-ray irradiation is delayed, the transition to the FPD accumulation operation and readout operation cannot be performed smoothly, the subject is exposed to unnecessary exposure, and artifacts due to the FPD operation delay occur on the X-ray image There is a concern that the image quality deteriorates.
This is because, in the pipeline type signal processing circuit, as described above, the time for one cycle from the integration of the charge in a certain row by the integration amplifier until the output of the corresponding digital voltage signal (dose detection signal) to the memory is performed. For example, when the start of X-ray irradiation is detected, when the X-ray irradiation is actually started at the timing when the integration of the charge is started by the integrating amplifier, a digital voltage signal (dose detection signal) This is because there is a delay of about 3 cycles until the output of the signal is output to the memory and it is determined that the irradiation is started.

  Specifically, when a dose detection operation is performed in a constant cycle using a pipeline type signal processing circuit as in the image readout operation shown in FIG. 16, the timing chart is as shown in FIG. FIG. 17 is substantially the same as FIG. 16 showing the timing chart of the image reading operation. In FIG. 17, the image signal (pixel signal charge “p”, analog image signal “P”, digital image signal “Pd” is shown. )) Instead of a dose detection signal (detection pixel charge “s”, analog dose detection signal “S”, digital dose detection signal “Sd”), and the time for one cycle is also Ts. It has become.

  In the image readout operation, readout of one row is performed at time T of one cycle, and it takes about twice the number of rows of T to read out one frame (all pixels), whereas the dose detection operation In FIG. 5, since the charges of a plurality of detection pixels are read out for each signal line, sampling of all the detection pixels is performed once in time Ts for one cycle. When detecting the start of irradiation, the dose detection signal at the time of the (N-1) th sampling and the dose detection signal at the time of the Nth sampling are compared to determine whether or not the dose detection signal has increased. The dose detection signal at the time of the previous sampling and the current dose detection signal read out in (1) are compared.

  As described above, in the pipeline type signal processing circuit, a time Ts corresponding to one cycle is left after the charge is integrated by the integrating amplifier until the digital voltage signal is output to the memory. Therefore, when detecting the start of X-ray irradiation, as shown in FIG. 17, the actual sampling timing (timing for integrating the charge s (N) of the detected pixel with the integrating amplifier) is actually started as shown in FIG. When X-ray irradiation is started, a delay of about 3 cycles (3Ts) occurs until the digital dose detection signal Sd (N) obtained by the Nth sampling is output from the buffer memory. The timing of irradiation start detection will also be delayed.

  In particular, immediately after the start of X-ray irradiation, the response of the X-ray source is slow and the amount of change in dose per unit time is small. Therefore, in order to ensure S / N when sampling the dose detection signal, the image signal P It is desirable that the time Ts of one cycle of the dose detection operation is longer than the time T of one cycle of the reading operation (50 to 500 μsec). In that case, the detection delay of 3 cycles caused by the pipeline processing becomes a problem that cannot be cleared up at the error level. When the X-ray irradiation time is as short as several milliseconds for reducing the exposure dose, this detection delay is a problem that cannot be ignored.

  However, Patent Documents 1 and 2 do not deal with the above-described problem that occurs when the dose detection signal of the dose detection sensor is extracted from the pipeline type signal processing circuit.

  The present invention has been made in view of the above problems, and even when the output of the dose detection sensor is extracted from the pipeline type signal processing circuit, the radiation image can be quickly processed based on the output of the dose detection sensor. An object of the present invention is to provide a detection device and a method of operating the same.

  The radiological image detection apparatus of the present invention includes an FPD that detects a radiographic image of a subject, a pipeline type signal processing circuit, a memory, an FPD, a signal processing circuit, and a control unit that controls the operation timing of the memory. . The FPD includes a plurality of pixels that accumulate charges according to the radiation dose emitted from the radiation source, a dose detection sensor that detects charges according to the radiation dose, and a plurality of signals provided for each column of pixels. A line is connected to the pixel and the dose detection sensor, and the charge accumulated in the pixel is output as an image signal, and a plurality of signal lines are arranged to output the charge detected by the dose detection sensor as a dose detection signal. Image pickup area. The signal processing circuit is provided for each of a plurality of signal lines, and a plurality of integration amplifiers that integrate electric charges and convert them into voltage signals, and a second voltage signal that is read continuously from the integration amplifiers are temporarily held. The first and second signal holding units have a voltage signal input to one of the first and second signal holding units and a signal output from the other is performed in parallel. The memory stores an image signal and a dose detection signal output as voltage signals from the signal processing circuit. In the read operation for outputting the image signal to the memory, the control unit performs signal integration and signal output by the first and second signal holding units for one integration until the integration amplifier starts resetting charge integration. Control is repeated in a regular cycle corresponding to the period. On the other hand, in the dose detection operation for outputting the dose detection signal to the memory, signal input and signal output by the first and second signal holding units are performed in two types of cycles: a first cycle and a second cycle shorter than the first cycle. Then, control is performed to sandwich the second cycle at least once between the two first cycles.

  The dose detection signal includes an irradiation start detection for detecting that radiation irradiation has started, an irradiation end detection for detecting that radiation irradiation has ended, an automatic exposure control for radiation images, and a gain to be given to the image signal during a readout operation. Used for at least one of the settings.

  In the normal cycle, the plurality of integrating amplifiers integrate the charges for one row of the plurality of pixels.

  A plurality of dose detection sensors are arranged in a dispersed manner in the imaging region, and in the first cycle and the second cycle, the plurality of integrating amplifiers integrate the charges detected by the plurality of dose detection sensors at once.

  The first cycle is preferably longer than the normal cycle.

  In the first cycle or the second cycle, only one voltage signal input from the integrating amplifier to the first signal holding unit or the second signal holding unit is treated as a dose detection signal, and the other voltage signal is used as a dose detection signal. It is not handled and is treated as a dummy signal that has no meaning in terms of data.

  In the first cycle and the second cycle, one integration period of the integration amplifier set in the dose detection operation is divided and assigned. Alternatively, an integration period corresponding to each time is assigned to the first cycle and the second cycle.

  The signal processing circuit is connected to the subsequent stage of the plurality of integration amplifiers, and samples and holds an analog voltage signal integrated by the integration amplifier, and converts the analog voltage signal sampled and held by the CDS circuit into a digital voltage signal. An A / D converter for conversion.

  The first and second signal holding units are two buffer memories connected in parallel between the A / D converter and the memory. Alternatively, two CDS circuits are connected in parallel at the subsequent stage of the integrating amplifier, and the first and second signal holding units are two CDS circuits. In these cases, the control unit alternately performs the first cycle and the second cycle.

  Two first and second signal holding units are provided, respectively, one is two buffer memories connected in parallel between the A / D converter and the memory, and the other is an integrating amplifier It is a CDS circuit that is connected in parallel in the latter stage. In this case, the control unit sandwiches the second cycle twice between the two first cycles.

  The control unit shortens the time of the second cycle with respect to the first cycle by controlling the number and period of the operation control signals given to the signal processing circuit.

  The dose detection sensor uses a part of the pixel. Specifically, the pixel is directly connected to the signal line without passing through the switching element, and the normal pixel that accumulates the signal charge upon receiving radiation and outputs the signal charge to the signal line according to the driving of the switching element. There is a detection pixel, and the detection pixel is used as a dose detection sensor. Or you may use the detection pixel provided with the switching element driven separately from a normal pixel as a dose detection sensor.

  In addition, it is preferable that the radiographic image detection apparatus of this invention is an electronic cassette with which FPD was accommodated in the portable housing | casing.

  Further, the operation method of the radiological image detection apparatus of the present invention controls the operation timing of the FPD for detecting the radiographic image of the subject, the pipeline type signal processing circuit, the memory, the FPD, the signal processing circuit, and the memory. It is an operating method of a radiographic image detection apparatus provided with a control part. The FPD includes a plurality of pixels that accumulate charges according to the radiation dose emitted from the radiation source, a dose detection sensor that detects charges according to the radiation dose, and a plurality of signals provided for each column of pixels. A line is connected to the pixel and the dose detection sensor, and the charge accumulated in the pixel is output as an image signal, and a plurality of signal lines are arranged to output the charge detected by the dose detection sensor as a dose detection signal. Image pickup area. The signal processing circuit is provided for each of a plurality of signal lines, and a plurality of integration amplifiers that integrate electric charges and convert them into voltage signals, and a second voltage signal that is read continuously from the integration amplifiers are temporarily held. The first and second signal holding units have a voltage signal input to one of the first and second signal holding units and a signal output from the other is performed in parallel. The memory stores an image signal and a dose detection signal output as voltage signals from the signal processing circuit. In the read operation for outputting the image signal to the memory, the signal input and the signal output by the first and second signal holding units correspond to one integration period from when the integration amplifier starts to integrate the charge and is reset. In the dose detection operation in which the control unit executes control repeated in a constant normal cycle, while the dose detection signal is output to the memory, signal input and signal output by the first and second signal holding units are performed in the first cycle. The control is performed in two types of cycles, ie, a second cycle shorter than the first cycle, and the control unit executes control for sandwiching the second cycle at least once between the two first cycles.

  In the present invention, when the dose detection signal that is the output of the dose detection sensor is extracted from the pipeline type signal processing circuit, the signal input and signal output of the dose detection signal by the first and second signal holding units are Based on two dose cycles, one cycle and a second cycle shorter than the first cycle, the second cycle is controlled at least once between the two first cycles. Processing can be performed.

1 is a schematic diagram of an X-ray imaging system. It is an external appearance perspective view which shows an electronic cassette. It is a block diagram which shows the internal structure of an electronic cassette. It is a timing chart which shows the dose detection operation | movement of a pipeline type signal processing circuit. It is a figure which shows transition of operation | movement of FPD in X-ray imaging. It is a block diagram which shows the internal structure of the electronic cassette which provided two CDS in the back | latter stage of each integration amplifier. 7 is a timing chart showing a dose detection operation of the pipeline type signal processing circuit of the example shown in FIG. 6. It is a block diagram which shows the internal structure of the electronic cassette which provided two buffer memories and provided two CDS in the back | latter stage of each integration amplifier. It is a timing chart which shows the dose detection operation | movement of the pipeline type signal processing circuit of the example shown in FIG. It is a block diagram which shows the internal structure of the electronic cassette which has an irradiation completion | finish detection part. 1 is a schematic diagram of an X-ray imaging system having a communication function between a radiation source control device and an electronic cassette. It is a block diagram which shows the internal structure of the electronic cassette which has an AEC part. It is a figure which shows transition of operation | movement of FPD in the X-ray imaging of the example shown in FIG. 11, FIG. It is a block diagram which shows a part of internal structure of the electronic cassette which has a gain setting part. It is a block diagram which shows another example of a detection pixel. It is a timing chart which shows the image reading operation | movement of a pipeline type signal processing circuit. It is a timing chart when performing a dose detection operation | movement by a fixed cycle similarly to an image read-out operation | movement using a pipeline type signal processing circuit.

[First Embodiment]
In FIG. 1, an X-ray imaging system 2 instructs an X-ray source 10, a radiation source controller 11 that controls the operation of the X-ray source 10, and a warm-up start and an X-ray irradiation start to the X-ray source 10. An irradiation switch 12 to detect the X-ray transmitted through the subject and output an X-ray image, a console 14 for controlling the operation of the electronic cassette 13 and displaying the X-ray image, and a subject. It has a standing photographing table 15 for photographing in a standing posture and a lying photographing table 16 for photographing in a lying posture. The X-ray source 10, the radiation source control device 11, and the irradiation switch 12 constitute an X-ray generator 2a, the electronic cassette 13, and the console 14 constitute an X-ray imaging device 2b, respectively. The X-ray generator 2a and the X-ray imaging apparatus 2b do not have a mutual communication function, and the electronic cassette 13 has an X-ray irradiation start detection function. In addition to this, a radiation source moving device or the like for setting the X-ray source 10 in a desired direction and position is provided, and the X-ray source 10 is shared by the standing imaging table 15 and the standing imaging table 16. .

  The X-ray source 10 includes an X-ray tube that emits X-rays, and an irradiation field limiter (collimator) that limits an X-ray irradiation field emitted by the X-ray tube. The X-ray tube has a cathode made of a filament that emits thermoelectrons, and an anode (target) that emits X-rays when the thermoelectrons emitted from the cathode collide. When there is an instruction to start warm-up, the anode starts rotating, and the warm-up ends when the specified rotational speed is reached. The irradiation field limiter has, for example, four lead plates that shield X-rays arranged on each side of a square, and a rectangular irradiation opening that transmits X-rays is formed in the center. By moving the position, the size of the irradiation aperture is changed to limit the irradiation field.

  The X-ray source controller 11 boosts the input voltage with a transformer to generate a high-voltage tube voltage and supplies the X-ray source 10 to the X-ray source 10 through a high-voltage cable, and the X-ray source 10 emits X-rays. A tube voltage that determines an energy spectrum, a tube current that determines an irradiation amount per unit time, and a control unit that controls an irradiation time of X-rays. Imaging conditions such as tube voltage, tube current, and irradiation time are manually set by an operator such as a radiographer through the operation panel of the radiation source control device 14.

  The irradiation switch 12 is, for example, a two-stage push switch operated by an operator at the start of X-ray imaging, and generates a warm-up start signal for starting warm-up of the X-ray source 10 by one-stage push (half press). Then, an irradiation start signal for starting the irradiation of the X-ray source 10 is generated by two-stage pressing (full pressing). These signals are input to the radiation source control device 11 through a signal cable.

  When receiving the irradiation start signal from the irradiation switch 12, the control unit of the radiation source control device 11 starts power supply from the high voltage generator to the X-ray source 10, and simultaneously operates a timer to Measure the irradiation time. Then, when the irradiation time set in the imaging conditions has elapsed, the power supply from the high voltage generator to the X-ray source 10 is stopped. Although the X-ray irradiation time varies depending on the imaging conditions, in the case of still image shooting, the maximum X-ray irradiation time is often set in the range of about 500 msec to about 2 s, and the irradiation time is the maximum. The irradiation time is set as the upper limit.

  The console 14 is communicably connected to the electronic cassette 13 by a wired method or a wireless method, and controls the operation of the electronic cassette 13 in accordance with an input operation from an operator via an input device 14a such as a keyboard. The X-ray image sent from the electronic cassette 13 is displayed on the display 14b of the console 14, and the data is stored in a storage device such as a hard disk or memory in the console 14, or a data storage such as an image storage server connected to the console 14 over a network. Is remembered.

  The console 14 receives an input of an inspection order including information such as the subject's sex, age, imaging region, and imaging purpose, and displays the inspection order on the display 14b. The examination order is input from an external system that manages patient information such as HIS (Hospital Information System) and RIS (Radiation Information System) and examination information related to radiation examination, or is manually input by an operator. The examination order includes imaging directions such as imaging parts such as the head, chest, and abdomen, front, side, oblique, PA (X-rays are irradiated from the back of the subject), and AP (X-rays are irradiated from the front of the subject). included. The operator confirms the contents of the inspection order on the display 14b, and inputs photographing conditions corresponding to the contents through the operation screen displayed on the display 14b.

  The imaging conditions input at the console 14 include various parameters such as tube voltage, tube current, and irradiation time, as in the case of the radiation source control device 11. These various parameters are stored in the console 14 for each imaging region, and by specifying a desired imaging region, it is possible to set imaging conditions according to this. Since the cumulative dose of X-rays is determined by the product of the tube current and the irradiation time, the tube current and the irradiation time are individually input as imaging conditions set in the radiation source control device 11 and the console 14. Instead, a tube current time product (mAs value), which is the product of both, may be input. The imaging conditions of the radiation source controller 11 are manually set by the operator with reference to the imaging conditions of the console 14.

  In FIG. 2, the electronic cassette 13 includes an FPD 35 and a portable box 20 having a flat box shape that accommodates the FPD 35. The housing 20 is made of, for example, a conductive resin. A rectangular opening is formed in the front surface 20a of the housing 20 on which X-rays enter, and a transmission plate 21 is attached to the opening as a top plate. The transmission plate 21 is made of a carbon material that is lightweight, has high rigidity, and has high X-ray permeability. The housing 20 also functions as an electromagnetic shield that prevents electromagnetic noise from entering the electronic cassette 13 and radiating electromagnetic noise from the electronic cassette 13 to the outside.

  The housing 20 has a size compliant with the international standard ISO 4090: 2001 similar to, for example, a film cassette or an IP cassette (also referred to as a CR cassette). A plurality of electronic cassettes 13 are provided in one room for the X-ray imaging system 2, for example, two for the standing imaging table 15 and the standing imaging table 16. The electronic cassette 13 is detachably attached to the holders 15a and 16a (see FIG. 1) of the standing position imaging table 15 and the standing position imaging table 16 so that the front surface 20a of the housing 20 is held in a posture facing the X-ray source 10. Set to Then, the X-ray source 10 is moved by the radiation source moving device according to the imaging table to be used. Further, the electronic cassette 13 may be used alone as it is placed on the bed on which the subject lies, or is held by the subject itself, in addition to being set on the standing position photographing stand 15 or the standing position photographing stand 16. . Note that the electronic cassette 13 is approximately the same size as the film cassette or the IP cassette, and therefore can be attached to an existing photographing stand for the film cassette or the IP cassette.

  In FIG. 3, the electronic cassette 13 includes a communication unit 30 and a battery 31 for communicating with the console 14 in a wired or wireless manner. The communication unit 30 mediates transmission / reception of various information and signals including image data of the console 14 and the control unit 32. The battery 31 supplies power of a predetermined voltage to each part of the electronic cassette 13. A relatively small battery 31 is used so as to fit in the thin electronic cassette 13. The battery 31 can be taken out from the electronic cassette 13 and set in a dedicated cradle for charging. Furthermore, the battery 31 may be configured to be capable of wireless power feeding.

  The communication unit 30 is wired to the console 14 when wireless communication between the electronic cassette 13 and the console 14 becomes impossible due to a shortage of the remaining battery 31 or the like. When a cable from the console 14 is connected to the communication unit 30, wired communication with the console 14 is possible. At this time, power may be supplied from the console 14 to the electronic cassette 13.

  The FPD 35 has a TFT active matrix substrate, and an imaging region 37 is formed on this substrate. In the imaging region 37, a plurality of pixels 36 that accumulate charges according to the X-ray arrival dose are arranged in a matrix of n rows (x direction) × m columns (y direction) at a predetermined pitch. n and m are integers of 2 or more, for example, n and m≈2000. Note that the arrangement of the pixels 36 may not be a square arrangement, but may be a honeycomb arrangement.

The FPD 35 has a scintillator (phosphor, not shown) that converts X-rays into visible light, and is an indirect conversion type in which visible light converted by the scintillator is photoelectrically converted by the pixels 36. The scintillator is made of CsI: Tl (thallium activated cesium iodide), GOS (Gd ₂ O ₂ S: Tb, terbium activated gadolinium oxysulfide), or the like, and is opposed to the entire surface of the imaging region 37 in which the pixels 36 are arranged. Is arranged. Note that the scintillator and the TFT active matrix substrate may be a PSS (Penetration Side Sampling) system in which the scintillator and the substrate are arranged in this order when viewed from the X-ray incident side. Side sampling) method may be used. Alternatively, a direct conversion type FPD using a conversion layer (such as amorphous selenium) that directly converts X-rays into charges may be used without using a scintillator.

  The pixel 36 includes a photoelectric conversion unit 38 that generates and accumulates charges (electron-hole pairs) by incidence of visible light, and a thin film transistor (TFT) 39 that is a switching element.

  The photoelectric conversion unit 38 has a structure in which a semiconductor layer (for example, PIN type) that generates electric charges and an upper electrode and a lower electrode are arranged above and below the semiconductor layer. The photoelectric conversion unit 38 has a TFT 39 connected to the lower electrode and a bias line (not shown) connected to the upper electrode. The bias lines are provided for the number of rows (n rows) of the pixels 36 in the imaging region 37 and are bound to one bus. The bus is connected to a bias power source. A bias voltage is applied from the bias power source to the upper electrode of the photoelectric conversion unit 38 through the bus line and its bias line. An electric field is generated in the semiconductor layer by applying a bias voltage, and charges (electron-hole pairs) generated in the semiconductor layer by photoelectric conversion move to the upper and lower electrodes, one having a positive polarity and the other having a negative polarity. Then, charges are accumulated in the photoelectric conversion unit 38.

  The TFT 39 has a gate electrode connected to the scanning line 40, a source electrode connected to the signal line 41, and a drain electrode connected to the photoelectric conversion unit 38. The scanning lines 40 and the signal lines 41 are wired in a grid pattern. The scanning lines 40 are common to the pixels 36 for one row, and the signal lines 41 correspond to the number of rows (n rows) of the pixels 36. Are provided in common for one column of pixels 41, one for each column of pixels 36 (m columns). The scanning line 40 is connected to the gate driver 42, and the signal line 41 is connected to the signal processing circuit 43.

  The gate driver 42 drives the TFT 39 to perform an accumulation operation for accumulating signal charges corresponding to the X-ray arrival dose in the pixel 36, a read operation for reading the signal charges from the pixel 36, and a reset operation. The control unit 32 controls the start timing of each of the operations executed by the gate driver 42.

  In the accumulation operation, the TFT 39 is turned off, and signal charges are accumulated in the pixel 36 during that time. In the read operation, gate pulses G1 to Gn for simultaneously driving the TFTs 39 in the same row are generated sequentially from the gate driver 42, the scanning lines 40 are sequentially activated one by one, and the TFTs 39 connected to the scanning lines 40 are provided for one row. Turn on one by one. The charge accumulated in the pixel 36 is read out to the signal line 41 and input to the signal processing circuit 43 when the TFT 39 is turned on. In order to shorten the readout operation time, thinning readout for reading out charges every predetermined row by generating a gate pulse, or binning readout for simultaneously reading out charges from a plurality of rows by giving gate pulses to a plurality of rows at the same time May be executed.

  Dark charges are generated in the semiconductor layer of the photoelectric conversion unit 38 regardless of whether X-rays are incident. This dark charge is accumulated in the photoelectric conversion unit 38 because a bias voltage is applied. Since the dark charge generated in the pixel 36 becomes a noise component for the image data, a reset operation is performed at predetermined time intervals in order to remove this. The reset operation is an operation for sweeping out dark charges generated in the pixels 36 through the signal line 41.

  The reset operation is performed, for example, by a sequential reset method in which the pixels 36 are reset row by row. In the sequential reset method, as in the signal charge reading operation, gate pulses G1 to Gn are sequentially generated from the gate driver 42 to the scanning line 40, and the TFTs 39 of the pixels 36 are turned on line by line.

  Instead of the sequential reset method, multiple rows of pixels are grouped as one group, and the reset is performed sequentially within the group, and the dark charge of the number of groups of rows is simultaneously swept away. An all-pixel reset method that simultaneously sweeps out dark charges may be used. The reset operation can be speeded up by a parallel reset method or an all-pixel reset method.

  The signal processing circuit 43 is an application specific integrated circuit (ASIC) that performs pipeline processing based on various operation control signals (see FIG. 4 and the like) such as a Sync signal, an Internal Reset signal, and an Analog Clock signal given from the control unit 32. In the image reading operation, the signal processing circuit 43 performs pipeline processing that repeats a normal cycle with a constant cycle time T, as in the conventional example of FIG. On the other hand, in the dose detection operation for detecting the dose detection signal Sd used for detecting the start of irradiation shown in FIG. 4, for example, the signal processing circuit 43 is used as described later in order to speed up the output of the dose detection signal Sd. During the first cycle at time T1, pipeline processing is performed with a second cycle at time T2 shorter than time T1 interposed therebetween.

  The signal processing circuit 43 includes an integrating amplifier 50, a CDS circuit (CDS) 51, a multiplexer (MUX) 52, an A / D converter (A / D) 53, and first and second buffer memories 54a and 54b.

  The integrating amplifier 50 is individually connected to each signal line 41. The integrating amplifier 50 includes an operational amplifier 50a and a capacitor 50b connected between the input and output terminals of the operational amplifier 50a, and the signal line 41 is connected to one input terminal of the operational amplifier 50a. The other input terminal of the operational amplifier 50a is connected to the ground (GND). A reset switch 50c is connected in parallel to the capacitor 50b. The integrating amplifier 50 integrates the charge input from the signal line 41, converts it into analog voltage signals V1 to Vm, and outputs it. The MUX 52 is connected to the output terminal of the operational amplifier 50a in each column via the amplifier 55 and the CDS 51. An A / D 53 is connected to the output side of the MUX 52.

  The CDS 51 has two sample and hold circuits, performs correlated double sampling on the output voltage signal of the integrating amplifier 50 to remove noise, and holds the output voltage signal of the integrating amplifier 50 for a predetermined period by the sample and hold circuit. (Sample hold). Based on an operation control signal from a shift register (not shown) in the control unit 32, the MUX 52 selects one CDS 51 with the electronic switch in order from the CDS 51 of each column connected in parallel, and is output from the selected CDS 51. Voltage signals V1 to Vm are serially input to the A / D 53. The A / D 53 converts the input voltage signals V1 to Vm into digital voltage signals, and alternately outputs them to the first buffer memory 54a or the second buffer memory 54b connected in parallel.

  The first and second buffer memories 54a and 54b are signal holding units for realizing pipeline processing, and are line memories capable of recording digital voltage signals representing one row of X-ray images. The first and second buffer memories 54a and 54b have m memory cells in which the digital voltage signals V1 to Vm of each column are respectively stored. The first and second buffer memories 54a and 54b temporarily store (signal input) one row of digital voltage signals output from the A / D 53 and output them to the memory 60 built in the electronic cassette 13 (signals). Output. In the memory 60, a digital voltage signal for one row is recorded as image data representing an X-ray image for one row in association with the coordinates of each pixel 36. Thus, reading for one row is completed. An amplifier may be connected between the MUX 52 and the A / D 53.

  When the reading for one row is completed, the control unit 32 outputs a reset pulse RST to the integrating amplifier 50 and turns on the reset switch 50c. As a result, the signal charge for one row accumulated in the capacitor 50b is discharged and reset. After resetting the integrating amplifier 50, the reset switch 50c is turned off again, and one of the sample hold circuits of the CDS 51 is held after a predetermined time has elapsed, and the kTC noise component of the integrating amplifier 50 is sampled. Thereafter, a gate pulse of the next row is output from the gate driver 42 to start reading signal charges of the pixels 36 of the next row. Further, after the gate pulse is output, the signal charge of the pixel 36 in the next row is held by another sample and hold circuit of the CDS 51 after a predetermined time elapses. These operations are sequentially repeated to read the signal charges of the pixels 36 in all rows.

  When reading of all rows is completed, image data representing one X-ray image is recorded in the memory 60. The image data is read from the memory 60 to the control unit 32, subjected to various image processing by the control unit 32, and then output to the console 14 through the communication unit 30. Thus, an X-ray image of the subject is detected.

  In the reset operation, dark charges flow from the pixel 36 to the capacitor 50b of the integrating amplifier 50 through the signal line 41 while the TFT 39 is in the ON state. Unlike the read operation, the MUX 52 does not read out the electric charge accumulated in the capacitor 50b, and in synchronization with the generation of the gate pulses G1 to Gn, the reset pulse RST is output from the control unit 32 and the reset switch 50c is turned on. The charge accumulated in the capacitor 50b is discharged, and the integrating amplifier 50 is reset.

  The control unit 32 is provided with a circuit (not shown) that performs various image processing such as offset correction, sensitivity correction, and defect correction on the X-ray image data in the memory 60. The offset correction circuit subtracts the offset correction image acquired from the FPD 35 without irradiating the X-ray in units of pixels from the X-ray image, thereby removing fixed pattern noise caused by individual differences in the signal processing circuit 43 and the imaging environment. . The sensitivity correction circuit is also called a gain correction circuit, and corrects variations in sensitivity of the photoelectric conversion unit 38 of each pixel 36, variations in output characteristics of the signal processing circuit 43, and the like. The defect correction circuit linearly interpolates the pixel value of the defective pixel with the pixel values of the surrounding normal pixels based on the defective pixel information generated at the time of shipment or regular inspection. The defect correction circuit also interpolates the pixel values of the pixels 36 in the column where the detection pixels 65 described later are arranged. The various image processing circuits described above may be provided in the console 14 and various image processing may be performed by the console 14.

  In addition to the normal pixel 36 connected to the signal line 41 via the TFT 39 as described above, the FPD 35 includes the detection pixel 65 connected to the signal line 41 without passing through the TFT 39 in the same imaging region 37. There are several. The detection pixel 65 is a pixel used for detecting the arrival dose of X-rays to the imaging region 37, and functions as a dose detection sensor for detecting the start of irradiation. In the figure, the detection pixel 65 is hatched to be distinguished from the normal pixel 36.

  The detection pixels 65 are arranged so as to be evenly distributed in the imaging region 37 without being locally biased in the imaging region 37. One detection pixel 65 is provided for each column of pixels 36 to which the same signal line 41 is connected, and the columns where the detection pixels 65 are provided sandwich, for example, two to three columns where the detection pixels 65 are not provided. Provided. The positions of the detection pixels 65 are known when the FPD 35 is manufactured, and the FPD 35 stores the positions (coordinates) of all the detection pixels 65 in a non-volatile memory (not shown) in advance. In contrast to the present embodiment, the detection pixels 65 may be locally concentrated and the arrangement of the detection pixels 65 can be changed as appropriate. For example, in a mammography apparatus that captures an image of the breast, the detection pixels 65 may be arranged concentrated on the chest wall side. A plurality of detection pixels 65 may be provided in each column.

  The basic configuration of the pixel 36 and the detection pixel 65, such as the photoelectric conversion unit 38, is exactly the same. Therefore, it can be formed by substantially the same manufacturing process. The detection pixel 65 is not provided with the TFT 39 between the signal line 41 and is directly connected to the signal line 41. For this reason, the signal charge generated in the photoelectric conversion unit 38 of the detection pixel 65 flows out to the signal line 41 regardless of whether the TFT 39 is on or off. The normal pixel 36 can turn off the TFT 39 and read out the charge even during the accumulation operation for accumulating the signal charge. Therefore, the charge s (see FIG. 4) generated in the detection pixel 65 always flows into the capacitor 50b of the integrating amplifier 50 on the signal line 41 to which the detection pixel 65 is connected. The charge s from the detection pixel 65 accumulated in the integrating amplifier 50 is output to the A / D 53 and converted into a digital voltage signal (hereinafter referred to as a dose detection signal Sd) by the A / D 53. The dose detection signal Sd is input to the first buffer memory 54 a and then output to the memory 60. In addition, a dummy signal Dd, which will be described later, is input to the second buffer memory 54b while the dose detection signal Sd is output from the first buffer memory 54a, and the next dose detection signal Sd is input to the first buffer memory 54a. During this time, the dummy signal Dd is output from the second buffer memory 54b. The FPD 35 repeats such a dose detection operation a plurality of times until the irradiation start detection unit 61 determines that X-ray irradiation has started based on the dose detection signal Sd.

  The irradiation start detector 61 is driven and controlled by the controller 32. The irradiation start detection unit 61 reads out the dose detection signal Sd from the memory 60 and performs X-ray irradiation start detection based on the read out dose detection signal Sd. Specifically, the maximum value among the dose detection signals Sd read out from the memory 60 and a preset irradiation start threshold value are compared for each output of the dose detection signal Sd. The irradiation start detection unit 61 determines that X-ray irradiation has started when the maximum value of the dose detection signal Sd is equal to or greater than the irradiation start threshold, and outputs an irradiation start detection signal to the control unit 32. The same value is set as the irradiation start threshold regardless of the imaging conditions.

  During the dose detection operation, the signal processing circuit 43 operates as shown in FIG. Note that the meanings of the legends of the Sync signal, Analog Clock signal, ADC Clock signal, and Buffer Clock signal are substantially the same as those in FIGS. 16 and 17, and the description thereof will be omitted (the same applies to FIG. 7 and the like). The Internal Reset signal is also a signal used for resetting the integration amplifier 50 as in the case of FIGS. 16 and 17. However, in the example of FIG. 4, the integration amplifier 50 is not reset at the end of the first cycle. The integration amplifier 50 is reset only when the second cycle following the first cycle is completed. “Second” in the row of the operation buffer represents the second buffer memory 54b (the same applies to FIG. 9).

  In FIG. 4, the signal processing circuit 43 alternately performs the first cycle and the second cycle as one set. The point that the first cycle and the second cycle that is extremely shorter than the first cycle are alternately performed is different from the read operation in which a constant normal cycle is repeated. The time T1 of the first cycle is set longer than the time T of one cycle during the read operation in order to ensure the S / N of the dose detection signal Sd (T1> T). The time T1 of the first cycle is, for example, 50 to 500 μsec, which is several times to several tens of times the time T of one cycle of the read operation. The total time T1 + T2 of the first cycle and the second cycle corresponds to one integration period of the integration amplifier 50 set in the dose detection operation. The operation buffer for the first cycle is the first buffer memory 54a, and the operation buffer for the second cycle is the second buffer memory 54b. If the S / N of the dose detection signal Sd can be secured sufficiently, the time T1 of the first cycle may be the same as or shorter than the time T of the normal cycle.

  Here, paying attention to the dose detection signal Sd (N), in the first cycle, the charge s (N) from the detection pixel 65 integrated by the integration amplifier 50 at the time of the previous set of the first cycle and the second cycle. The analog dose detection signal S (N) based on the above is sampled and held by the CDS 51 (see CDS row), and the dose detection signal S (N) is converted to the digital dose detection signal Sd (N) by the A / D 53 (AD The data is temporarily stored in the first buffer memory 54a, and at the same time, the digital dummy signal Dd (N-1) temporarily stored in the second buffer memory 54b in the previous second cycle is stored in the second buffer. The data is output from the memory 54b to the memory 60 (see the Data Out line).

  In the subsequent second cycle, the analog dummy signal D (N) is sampled and held, converted to the dummy signal Dd (N), and the dummy signal Dd (N) is temporarily stored in the second buffer memory 54b. The dose detection signal Sd (N) temporarily stored in the first buffer memory 54a is simultaneously output to the memory 60 in one cycle. By doing so, the A / D conversion of the dose detection signal S (N) and the temporary storage (signal input) of the dose detection signal Sd (N) in the first buffer memory 54a in one set of the first cycle and the second cycle. Then, the output (signal output) of the dose detection signal Sd (N) from the first buffer memory 54a to the memory 60 can be completed. In the dose detection operation, since the charges of all the detection pixels 65 are collectively read for each cycle, the dose detection signal Sd from all the detection pixels 65 is stored in the memory 60 in one set of the first cycle and the second cycle. To be recorded.

  The dummy signal Dd is merely a signal that is used for convenience in order to output the digital dose detection signal Sd temporarily stored in the first buffer memory 54a in the first cycle to the memory 60, and is used for detecting the start of X-ray irradiation. It is a signal that does not contribute and is meaningless in terms of data. For this reason, the memory 60 receives only the dose detection signal Sd output from the first buffer memory 54a in the second cycle, and discards the dummy signal Dd output from the second buffer memory 54b in the first cycle without receiving it.

  As indicated by a dot-and-dash circle A, in the second cycle, the period of the Analog Clock signal is made extremely shorter than in the first cycle in order to shorten the time for sample holding of the dummy signal D. . By making the time T2 of the second cycle much shorter than the time T1 of the first cycle (T1 >> T2), the dose detection signal Sd is output to the memory 60 substantially immediately after the end of the A / D conversion to detect the start of irradiation. To be served.

  Next, with reference to FIGS. 4, 5, 16, and 17, a procedure for performing X-ray imaging in the X-ray imaging system 2 will be described.

  First, the subject is placed at a predetermined position in front of the standing position imaging stand 15 or placed on the supine position imaging stand 16, and the electronic cassette 13 set on the standing position or standing position imaging stand 15, 16 is set. Adjust the position and position of the subject by adjusting the height and horizontal position. Further, the height, horizontal position, and size of the irradiation field of the X-ray source 10 are adjusted according to the position of the electronic cassette 13 and the size of the imaging region. Next, imaging conditions are set in the radiation source control device 11 and the console 14.

  In FIG. 5, the controller 32 repeatedly causes the FPD 35 to perform a reset operation before X-ray imaging. After the positions of the subject and the electronic cassette 13 are adjusted and the imaging conditions are set in the radiation source control device 11, when an imaging preparation instruction is transmitted from the console 14 by an operator's input operation, the control unit 32 controls the FPD 35. The operation is shifted from the reset operation to the dose detection operation. Then, the charge generated in the detection pixel 65 flows from the signal line 41 to the capacitor 50b of the integration amplifier 50. In the dose detection operation, since the TFT 39 is turned off, charges are accumulated in the normal pixel 36, but this charge is discarded by the reset operation after detection of irradiation start.

  In the dose detection operation, the signal processing circuit 43 repeats pipeline processing with the first cycle and the subsequent second cycle as one set. In the first cycle, the analog dose detection signal S based on the charge s from the detection pixel 65 integrated in the integration amplifier 50 in the previous set is sample-held by the CDS 51, and converted into the dose detection signal Sd by the A / D 53. In addition, the dose detection signal Sd is temporarily stored in the first buffer memory 54a. The dummy signal Dd is output from the second buffer memory 54b to the memory 60 simultaneously with the input of the dose detection signal Sd to the first buffer memory 54a, but is discarded as unnecessary data. In the second cycle, the dose detection signal Sd of the first buffer memory 54 a is output to the memory 60. The dose detection signal Sd output to the memory 60 is read by the irradiation start detection unit 61, and the maximum value among them is compared with the irradiation start threshold value.

  When the irradiation switch 12 is pushed by the operator and X-rays are irradiated from the X-ray source 10, the maximum value of the dose detection signal Sd soon becomes equal to or greater than the irradiation start threshold, and the irradiation start detector 61 starts X-ray irradiation. The irradiation start detection signal is output to the control unit 32. Upon receiving the irradiation start detection signal, the control unit 32 causes the FPD 35 to perform a reset operation once and then starts an accumulation operation. As a result, the X-ray irradiation start timing and the accumulation operation start timing of the FPD 35 are synchronized.

  The control unit 32 causes the FPD 35 to perform the accumulation operation until the time set in the shooting conditions, and then shifts to an image reading operation. In the read operation, the same pipeline processing as in the conventional example of FIG. 16 is performed by the signal processing circuit 43, and the digital image signal Pd is alternately output to the memory 60 from the first and second buffer memories 54a and 54b. After the read operation is completed, the FPD 35 returns to the reset operation. The image data output to the memory 60 by the read operation is transmitted to the console 14 via the communication unit 30 after various image processing, and is displayed on the display 14b as an X-ray image.

  When the dose detection operation is performed in a constant cycle as in the example of FIG. 17, the pipeline processing is repeated in the first cycle without performing the second cycle, and therefore temporarily stored in the buffer memory 54a or 54b. It takes time T1 for one cycle until the dose detection signal Sd is output to the memory 60. When the dose detection signal Sd is output, the time T1 is set several times to several tens of times longer than the time T of one cycle when the image signal Pd is output in order to ensure S / N. In this embodiment, the delay of one cycle in the pipeline processing becomes so large that it cannot be ignored, but in this embodiment, the output of the dose detection signal Sd can be speeded up by executing the second cycle. Specifically, when the X-ray irradiation is actually started at the timing at which the integration of the electric charge is started by the integrating amplifier 50, the irradiation start detection unit 61 determines that the irradiation starts from about 3 in the example of FIG. Although a delay corresponding to the cycle occurs, in this example, since the second cycle is an extremely short time, a delay corresponding to two cycles is substantially sufficient. Therefore, the time from when the X-ray is actually irradiated to when the irradiation start detection unit 61 detects the irradiation start can be made shorter than before. As a result, unnecessary exposure to the subject can be reduced, and a special effect that the irradiated X-ray can be reflected in the X-ray image without waste can be obtained.

  In this embodiment, the example is described in which the integration amplifier 50 is not reset at the end of the first cycle, but is reset only at the end of the second cycle. However, the reset is performed at the end of the first cycle. May be. In this case, the integration time of the dose detection signal is the time T1 of the first cycle, and the buffer memories 54a and 54b in which the dose detection signal and the dummy signal are temporarily stored are reversed. That is, the charge s integrated during the first cycle time T1 used as the dose detection signal is read to the CDS 51 at the end of the first cycle, and in the subsequent second cycle, the second buffer memory 54b. Is temporarily stored. Then, it is output from the second buffer memory 54b to the memory 60 in the next first cycle. On the contrary, the dummy signal is temporarily stored in the first buffer memory 54a in the first cycle, and is output from the first buffer memory 54a to the memory 60 in the next second cycle.

[Second Embodiment]
In the first embodiment, the first and second buffer memories 54a and 54b are provided as signal holding units to realize pipeline processing. However, as in the FPD 70 shown in FIG. A signal processing circuit 72 may be used as a pipeline processing signal holding unit by connecting two CDSs 71a and 71b connected in parallel to each other via an amplifier 55.

  In this case, during the dose detection operation, the signal processing circuit 72 operates as shown in FIG. The row of the operation CDS indicates the first and second CDSs 71a and 71b that sample and hold the dose detection signal S or the dummy signal D in the cycle. The first CDS 71a starts sampling and holding the dose detection signal S during the first cycle, and the second CDS 71b starts sampling and holding the dummy signal D during the second cycle (see the rows of the first and second CDS).

  In FIG. 7, in the present embodiment, the first cycle and the second cycle are alternately performed as one set, similarly to the first embodiment. In the first cycle, the dose detection signal S (N) based on the charge s (N) integrated by the integrating amplifier 50 at the previous set is sampled and held by the first CDS 71a (refer to the row of the first CDS). The dummy signal D (N−1) sampled and held in the second CDS 71b in the second cycle is digitally converted by the A / D 53 and output to the memory 60 (see the AD Data row). In the present embodiment, there is no buffer memory in the subsequent stage of the A / D 53, and the A / D 53 and the memory 60 are directly connected. Therefore, the AD Data row corresponds to the Data Out in the first embodiment.

  In the subsequent second cycle, the dummy signal D (N) is sampled and held in the second CDS 71b, and the digital dose detection signal Sd (N) of the dose detection signal S (N) sampled and held in the first CDS 71a in the previous first cycle. Conversion to memory and output to the memory 60 are performed simultaneously. Also in the present embodiment, as indicated by the dashed-dotted circle B, the time for sample holding of the dummy signal D is shortened to shorten the time T2 of the second cycle, as in the first embodiment. In the second cycle, the period of the Analog Clock signal is shortened. Also in this embodiment, the same effect as the first embodiment can be obtained. Since the configuration and operation other than substituting the buffer memory with CDS are the same as those in the first embodiment, the description thereof is omitted.

[Third Embodiment]
Further, a signal processing circuit in which the first and second CDSs 81a and 81b and the first and second buffer memories 82a and 82b are provided as signal holding units as in the FPD 80 shown in FIG. 8 by combining the first and second embodiments. 83 may be used. In this case, during the dose detection operation, the signal processing circuit 83 operates as shown in FIG.

In the dose detection operation, the signal processing circuit 83 repeats the pipeline processing in which the short second cycle is continuously performed twice after the first cycle as one set. In the first cycle, the dose detection signal S (N) based on the charge s (N) integrated by the integrating amplifier 50 at the previous set is sampled and held by the first CDS 81a (see the first CDS row), and the previous cycle The dummy signal D ₂ (N−1) sampled and held in the second CDS 81b in the second cycle of the second set is digitally converted by the A / D 53 (see the AD Data row) and the dummy signal D ₂ d (N− 1), which is temporarily stored in the first buffer memory 82a. At the same time, the dummy signal D ₁ d (N−1) temporarily stored in the second buffer memory 82b in the second cycle of the second time of the previous set is output from the second buffer memory 82b to the memory 60 (see Data Out row). To do.

In the subsequent second cycle, the dummy signal D ₁ (N) is sampled and held in the second CDS 81b, the dose detection signal S (N) sampled and held in the first CDS 81a in the previous first cycle, and digital Temporary storage of the converted dose detection signal Sd (N) in the second buffer memory 82b, to the memory 60 of the dummy signal D ₂ d (N−1) temporarily stored in the first buffer memory 82a in the first cycle. Are simultaneously output. In the second cycle, the dummy signal D ₂ (N) is sampled and held in the first CDS 81a. In the second cycle, the dummy signal D ₁ (N) sampled and held in the second CDS 81b is converted to digital. The dummy signal D ₁ d (N) is temporarily stored in the first buffer memory 82a, and the dose detection signal Sd (N) temporarily stored in the second buffer memory 82b in the first second cycle is stored in the memory 60. Output simultaneously. In this case as well, as in the first and second embodiments, the period of the Analog Clock signal is shortened in the second cycle, as indicated by the dot-dash circle C. Thus, by shortening the time T2 of the second cycle, the dose detection signal Sd can be output to the memory 60 at a shorter interval than in the conventional pipeline processing.

  In the second cycle of the first and second embodiments, the time required to output the dose detection signal Sd from the first buffer memory 54a to the memory 60 (first embodiment), or the dose sampled and held in the first CDS 71a. It is only necessary to secure the time (second embodiment) required for digital conversion of the detection signal S (N). Further, since the dummy signal Dd is a signal that does not contribute to the detection of the start of X-ray irradiation as described above, sample hold and A / D conversion of the dummy signal D do not need to be performed exactly as in the first cycle. . In the first and second embodiments, the period of the Analog Clock signal is shortened in the second cycle in order to shorten the time of the second cycle, but the period is the same as that in the first cycle, and the number of clocks of the Analog Clock signal is The sample hold and A / D conversion of the dummy signal D may be rounded up before the number of the first cycle is less, and only these may be performed formally, or only the Sync signal without inputting the Analog Clock signal. It is also possible to input only the dose detection signal Sd to the memory 60 without performing the sample hold and the A / D conversion of the dummy signal D by inputting with an accelerated period. Thereby, the time T2 of the second cycle can be further shortened, and the output of the dose detection signal Sd can be further increased.

  Similarly, the output of the dummy signal Dd from the buffer memory in the first second cycle and the A / D conversion of the dummy signal D in the second cycle in the third embodiment do not need to be performed strictly. By rounding up these processes before the number of clocks of the Buffer Data Clock signal in the second cycle of the second cycle and the number of clocks of the ADC Clock signal in the second cycle of the second cycle are less than the number in the first cycle, the second cycle is further accelerated. You may plan. Specifically, the output of the dummy signal Dd from the buffer memory and the A / D conversion of the dummy signal D are not performed for all the columns but are stopped halfway. Alternatively, the processing may be thinned out by skipping j columns by performing odd columns or even columns, every plural columns, or i columns continuously. Furthermore, it is not necessary to perform output from each memory cell of the buffer memory or A / D conversion without providing the Buffer Data Clock signal and the ADC Clock signal itself. Similarly, for the operation control signal given to the MUX 52, in the second cycle, the cycle is shortened, the output of the operation control signal is stopped halfway, or the number of clocks is made smaller than in the first cycle by giving it thinned out. Alternatively, the operation control signal may not be output to the MUX 52 itself. A combination of shortening the period of the Analog Clock signal, selecting a sample and hold circuit by MUX, A / D conversion, stopping output from the buffer memory, and thinning may be performed.

  Instead of detecting the start of X-ray irradiation based on the dose detection signal Sd, the end of X-ray irradiation may be detected based on the dose detection signal Sd. In this case, an irradiation end detection unit 86 is provided instead of the irradiation start detection unit 61 as in the FPD 85 shown in FIG. Similar to the irradiation start detection, the irradiation end detection unit 86 compares the maximum value of the dose detection signal Sd obtained by the dose detection operation that intersects the first cycle and the second cycle with the irradiation end threshold, and the maximum value is If it falls below the irradiation end threshold, it is determined that X-ray irradiation has ended. When the irradiation end detection unit 86 determines the irradiation end, the control unit 32 shifts the FPD 85 from the accumulation operation to the reading operation. The FPD 85 is the same as the FPD 35 in FIG. 3 except that an irradiation end detection unit 86 is provided, and thus the description thereof is omitted. Moreover, both the irradiation start detection part 61 and the irradiation end detection part 86 may be provided, and the X-ray irradiation start and end may be detected based on the dose detection signal Sd.

  Further, as described below, an automatic exposure control (AEC) operation may be performed based on the dose detection signal Sd.

  In FIG. 11, an X-ray imaging system 90 includes a radiation source controller 91 and an electronic cassette 92 having a communication function for exchanging various synchronization signals with each other. As shown in FIG. 12, the APD unit 96 is provided in the FPD 95 of the electronic cassette 92 in place of the irradiation start detection unit 61. Communication with the radiation source control device 91 is performed by the communication unit 97. The communication unit 97 communicates with the console 14 in the same manner as the communication unit 30. The X-ray imaging system 90 is different from the X-ray imaging system 2 of FIG. 1 only in that the radiation source control device 91 and the electronic cassette 92 have a communication function. Since it is the same, description is abbreviate | omitted. Similarly, the FPD 95 is the same as the FPD 35 in FIG. 3 except that the AEC unit 96 and the like are provided.

  In this case, in the radiation source control device 91, in order to prevent the X-ray irradiation from ending and falling short before the target dose is reached and the AEC unit 96 determines to stop the irradiation, the X-ray irradiation time is set. A value with a margin is set. In the X-ray source 10, the maximum value of the irradiation time set according to the imaging region for safety regulation may be set. The radiation source control device 91 performs X-ray irradiation control with the tube voltage, tube current, and irradiation time under the set imaging conditions. On the other hand, when the AEC function determines that the accumulated dose of X-rays has reached a necessary and sufficient target dose, the X-ray irradiation is stopped even if the irradiation time is shorter than the irradiation time set in the radiation source control device 11. To work.

  In this case, the imaging conditions set by the console 14 store information on the lighting field for measuring the accumulated dose by the AEC unit 96, and when the imaging conditions are set by the console 14, the information on the lighting field is stored. Provided to the electronic cassette 92. The daylighting field corresponds to a region of interest that is most noticeable at the time of diagnosis, and is a portion where the dose detection signal Sd can be stably obtained. For example, when the imaging region is the chest, a portion corresponding to the left and right lung fields is set.

  As shown in FIG. 13, when the radiation source control device 91 receives a warm-up start signal from the irradiation switch 12, an irradiation start request signal that is a synchronization signal that inquires whether or not to start X-ray irradiation. Is transmitted to the electronic cassette 92. When the electronic cassette 92 receives the irradiation start request signal by the communication unit 97, the electronic cassette 92 causes the FPD 95 to finish the reset operation and start the accumulation operation. Then, an irradiation permission signal that is a response to the irradiation start request signal is transmitted from the communication unit 97 to the radiation source control device 91. The radiation source control device 91 receives the irradiation permission signal from the electronic cassette 92, and further starts the power supply from the high voltage generator to the X-ray source 10 when receiving the irradiation start signal from the irradiation switch 12. X-rays are emitted from the source 10. Further, when receiving the irradiation stop signal emitted from the communication unit 97, the radiation source control device 91 stops the power supply from the high voltage generator to the X-ray source 10 and stops the X-ray irradiation.

  The AEC unit 96 operates when the FPD 95 starts an accumulation operation. The signal processing circuit 43 performs the dose detection operation of each of the above embodiments, and periodically outputs the dose detection signal Sd to the memory 60.

  The AEC unit 96 measures the cumulative dose of X-rays reaching the imaging region 37 by sequentially adding, for each coordinate, the dose detection signal Sd read from the memory 60 by a plurality of outputs. More specifically, the AEC unit 96 represents the representative values (average value, maximum value, mode) of the dose detection signal Sd from the detection pixel 65 existing in the lighting field based on the lighting field information given from the console 14. Value, total value, etc.), and the representative values are added up to determine the cumulative dose in the lighting field.

  As a method for determining the daylighting field, the representative value of the dose detection signal may be integrated for each block obtained by dividing the imaging region 37 into a plurality of blocks, and the block having the lowest integrated value may be set as the daylighting field. An arbitrary part of the imaging region 37 may be designated as a lighting field.

  The AEC unit 96 compares the calculated accumulated dose in the lighting field with a preset irradiation stop threshold (target dose) each time the dose detection signal Sd is output, and the accumulated dose has reached the irradiation stop threshold. It is determined whether or not. The AEC unit 96 outputs an irradiation stop signal to the control unit 32 when it is determined that the accumulated dose in the lighting field exceeds the irradiation stop threshold and the accumulated dose of X-rays has reached the target dose.

  When it is determined by the AEC unit 96 that the accumulated dose of the X-rays reaching the lighting field has reached the target value, and an irradiation stop signal is output, the control unit 32 sends the radiation source control device 91 via the communication unit 97. Send an irradiation stop signal. Thereby, X-ray irradiation by the X-ray source 10 is stopped. At the same time, the control unit 32 shifts the operation of the FPD 95 from the accumulation operation to the read operation. After the read operation is completed, the FPD 95 returns to the reset operation.

  The AEC unit 96 may compare the integrated value of the dose detection signal Sd with the irradiation stop threshold value to calculate the time when the accumulated X-ray dose is expected to reach the target value. In this case, an irradiation stop signal is transmitted to the radiation source control device 91 after the predicted time has elapsed, or the predicted time is transmitted to the radiation source control device 91. When transmitting the predicted time to the radiation source control device 91, the radiation source control device 91 measures the expected time and stops the X-ray irradiation when the expected time is reached. The controller 32 shifts the operation of the FPD 95 from the accumulation operation to the read operation after the predicted time has elapsed.

  Since the signal processing circuit 43 performs the dose detection operation of each embodiment to speed up the output of the dose detection signal Sd, the X-ray irradiation stop timing is delayed and the X-rays are irradiated more than the target value, and the object is detected. The special effect that the situation where an exposure dose increases can be avoided is acquired.

  In addition to using the dose detection signal Sd for X-ray irradiation start or end detection and AEC, the gain of the integrating amplifier during the read operation may be switched based on the dose detection signal Sd. In this case, as shown in FIG. 14, a variable gain type integration amplifier 100 is used in place of the integration amplifier 50.

  In FIG. 14, the integrating amplifier 100 includes an operational amplifier 100 a and a reset switch 100 c as with the integrating amplifier 50. Two capacitors 100b and 100d are connected between the input and output terminals of the operational amplifier 100a, and a gain changeover switch 100e is connected to the capacitor 100d. When the gain switch 100e is on, the output voltage signal V from the integrating amplifier is V = q / (C1 + C2), and when the gain switch 100e is off, V = q / C1. Where q is the accumulated charge, and C1 and C2 are the capacitances of the capacitors 100b and 100d, respectively. Thus, the gain of the integrating amplifier 100 can be changed by switching the gain changeover switch 100e on / off. Here, an example is shown in which two capacitors are connected and the gain is switched in two stages. However, it is preferable that two or more capacitors are connected and the gain can be changed in two or more stages.

  The gain setting unit 101 is provided in the FPD instead of the irradiation start detection unit 61, the irradiation end detection unit 86, and the AEC unit 96. The gain setting unit 101 operates when the FPD starts an accumulation operation, and controls the operation of the gain selector switch 100e during the read operation. The signal processing circuit performs the dose detection operation of each of the above embodiments, and periodically outputs the dose detection signal Sd to the memory 60. The gain setting unit 101 sets the gain of the integrating amplifier 100 to the minimum value in the dose detection operation so that the dose detection signal Sd is not saturated. In this example, the gain changeover switch 100e is turned on.

  Similar to the AEC unit 96, the gain setting unit 101 integrates the representative value of the dose detection signal Sd from the detection pixel 65 in the lighting field a predetermined number of times, and compares the integrated value with a preset threshold value. When the integrated value is larger than the threshold value, the gain setting unit 101 turns on the gain changeover switch 100e during the reading operation. On the other hand, when the accumulated accumulated dose to the portion corresponding to the region of interest in the imaging region 37 is low and the integrated value is less than or equal to the threshold value, the gain changeover switch 100e is turned off during the reading operation to increase the gain of the integrating amplifier 100. More specifically, the gain of the integrating amplifier 100 is set so that the maximum value and the minimum value of the output voltage signal V of the lighting field match the maximum value and the minimum value of the A / D conversion range.

  In radiography with a low X-ray cumulative dose, the maximum and minimum widths of the voltage signal V are narrower than the A / D conversion range, and the X-ray image obtained in such a case is unclear and conspicuous in noise. However, if the gain of the integrating amplifier is increased when the accumulated dose reaching the lighting field is low as described above, it is possible to obtain an X-ray image with good image quality in which noise is not noticeable. can get.

  Since the signal processing circuit performs the dose detection operation of each of the above-described embodiments to increase the output of the dose detection signal Sd, the X-ray irradiation time is several msec or less in low-dose imaging and the like, and the FPD accumulation operation is extremely short. Even when the process ends in time and shifts to the reading operation, the gain setting of the integrating amplifier can be made in time.

  Note that the gain of the amplifier may be set instead of the integrating amplifier. Further, detection of the start or end of X-ray irradiation, AEC, and gain setting of the integrating amplifier may be combined.

  In each of the above embodiments, the detection pixel 65 short-circuited to the signal line 41 without using the TFT 39 is used as the AEC sensor. However, the pixel 36 in which the source electrode and the drain electrode of the TFT 39 are short-circuited is used as the detection pixel. Also good. Further, for example, as in the detection pixel 105 shown in FIG. 15, a gate driver 106 different from the normal pixel 36 and a TFT 108 driven by the scanning line 107 are connected, and the accumulated charge is read out independently of the normal pixel 36. It is good also as a structure which can do. The gate driver 106 can selectively apply a gate pulse to the detection pixel 105 corresponding to the daylighting field to read out the accumulated charge of the detection pixel 105 corresponding to the daylighting field.

  In addition, a digital voltage signal based on a leaked charge leaking from the pixel 36 when all the TFTs 39 are turned off may be read out via a signal processing circuit, and a dose may be detected based on the digital voltage signal. In this case, all the pixels 36 are dose detection sensors. In addition, instead of modifying the normal pixels as detection pixels, a dose detection sensor is provided separately from the pixels in the gaps between the pixels, and the output of the dose detection sensor is read out via a signal processing circuit. Also good. In short, the present invention can be applied as long as the dose detection signal is output from the pipeline type signal processing circuit, and the same effects as those of the above-described embodiments are exhibited.

  In each of the above-described embodiments, an example in which the console 14 and the electronic cassette 13 are separate has been described. However, the console 14 does not have to be an independent device, and the function of the console 14 may be mounted on the electronic cassette 13. Similarly, the radiation source control device 11 and the console 14 may be integrated. Further, the present invention is not limited to an electronic cassette that is a portable X-ray image detection device, and may be applied to an X-ray image detection device that is installed on an imaging table.

  Furthermore, the present invention can be applied not only to X-rays but also to an imaging system that uses other radiation such as γ rays.

2, 90 X-ray imaging system 10 X-ray source 11, 91 Radiation source control device 13, 92 Electronic cassette 14 Console 32 Control unit 35, 70, 80, 85, 95 FPD
36 pixels 43, 72, 83 Signal processing circuit 50, 100 Integration amplifier 51, 71a, 71b, 81a, 81b CDS circuit (CDS)
53 A / D converter (A / D)
54a, 54b, 82a, 82b Buffer memory 61 Irradiation start detection unit 65, 105 Detection pixel 86 Irradiation end detection unit 96 AEC unit 101 Gain setting unit

Claims (24)

A plurality of pixels for accumulating charges according to the radiation dose emitted from the radiation source, a dose detection sensor for detecting charges according to the radiation dose, and a plurality of signal lines provided for each column of the pixels. Yes, the pixel and the dose detection sensor are connected, the charge output from the pixel is output as an image signal according to the driving of the switching element, and the switching element is driven separately from the switching element of the pixel. In response, the FPD that has an imaging region in which a plurality of signal lines that output the charges output from the dose detection sensor as dose detection signals are arranged, and that detects a radiographic image of the subject,
Provided for each of the plurality of signal lines, a plurality of integration amplifiers that integrate charges and convert them into voltage signals, and first and second voltage signals that are temporarily read from the integration amplifier and that are read twice. A pipeline-type signal processing circuit having a two-signal holding unit and executing a cycle of performing a signal input of a voltage signal to one of the first and second signal holding units and a signal output from the other in parallel When,
A memory for storing the image signal and the dose detection signal output as voltage signals from the signal processing circuit;
A control unit for controlling operation timing of the FPD, the signal processing circuit, and the memory;
In the read operation for outputting the image signal to the memory, the integration amplifier starts resetting charge and is reset as the cycle in which signal input and signal output by the first and second signal holding units are performed in parallel. Control to repeat a certain normal cycle corresponding to one integration period until
On the other hand, in the dose detection operation for outputting the dose detection signal to the memory, the first cycle and the first cycle are the cycles in which signal input and signal output by the first and second signal holding units are performed in parallel. A radiological image detection apparatus comprising: a control unit that performs two types of cycles that are shorter than the second cycle and executes control that sandwiches the second cycle at least once between the two first cycles.   The radiographic image detection apparatus according to claim 1, wherein in the imaging region, the pixel and the dose detection sensor are arranged at a position where they do not overlap.   The radiation image detection apparatus according to claim 1, wherein the dose detection sensor is disposed in a gap between the plurality of pixels.   4. The radiological image detection apparatus according to claim 1, wherein, in the normal cycle, the plurality of integrating amplifiers integrate charges for one row of the plurality of pixels. 5. A plurality of the dose detection sensors are arranged in a dispersed manner in the imaging region,
5. The radiographic image according to claim 1, wherein in the first cycle and the second cycle, the plurality of integrating amplifiers collectively integrate the charges detected by the plurality of dose detection sensors. Detection device.   The radiographic image detection apparatus according to claim 1, wherein the first cycle is longer than the normal cycle.   In the first cycle or the second cycle, only one voltage signal input from the integrating amplifier to the first signal holding unit or the second signal holding unit is treated as the dose detection signal, and the other voltage signal The radiation image detection apparatus according to claim 1, wherein the radiation image detection apparatus is not handled as the dose detection signal, but is handled as a dummy signal having no meaning in terms of data.   The said 1st cycle and the said 2nd cycle are divided in the integration period for one time of the said integration amplifier set in the said dose detection operation, The allocation of any one of Claim 1 thru | or 7 Radiation image detection device.   The radiological image detection apparatus according to claim 1, wherein an integration period corresponding to each time is assigned to the first cycle and the second cycle. The signal processing circuit is connected to a subsequent stage of the plurality of integrating amplifiers, and samples and holds an analog voltage signal integrated by the integrating amplifiers;
10. The radiological image detection apparatus according to claim 1, further comprising an A / D converter that converts an analog voltage signal sampled and held by the CDS circuit into a digital voltage signal. 11.   The radiographic image detection apparatus according to claim 10, wherein the first and second signal holding units are two buffer memories connected in parallel between the A / D converter and the memory. Two CDS circuits are connected in parallel at the subsequent stage of the integrating amplifier,
The radiographic image detection apparatus according to claim 10, wherein the first and second signal holding units are the two CDS circuits.   The radiological image detection apparatus according to claim 11, wherein the control unit alternately performs the first cycle and the second cycle. Two each of the first and second signal holding units are provided,
One is two buffer memories connected in parallel between the A / D converter and the memory,
11. The radiological image detection apparatus according to claim 10, wherein the other is the two CDS circuits connected in parallel at the subsequent stage of the integrating amplifier.   The radiological image detection apparatus according to claim 14, wherein the control unit sandwiches the second cycle twice between the two first cycles.   16. The control unit according to claim 1, wherein the control unit shortens the time of the second cycle with respect to the first cycle by controlling the number and cycle of operation control signals given to the signal processing circuit. The radiographic image detection apparatus described in 1.   The dose detection signal includes an irradiation start detection for detecting that radiation irradiation has started, an irradiation end detection for detecting that radiation irradiation has ended, an automatic exposure control of a radiographic image, and the image signal during the readout operation. The radiological image detection apparatus according to claim 1, wherein the radiological image detection apparatus is used for at least one of gain settings given to the radiography. In the automatic exposure control, every time the dose detection signal is output to the memory, the cumulative dose of the radiation is obtained by sequentially adding the dose detection signal,
The obtained cumulative dose and a preset irradiation stop threshold value are compared each time the dose detection signal is output to the memory to determine whether or not the cumulative dose has reached the irradiation stop threshold value. The radiological image detection apparatus according to claim 17, wherein a time when the cumulative dose is expected to reach the irradiation stop threshold is calculated.   The radiological image detection apparatus according to claim 18, wherein in the automatic exposure control, the cumulative dose of the daylighting field set in the imaging region is obtained. In the gain setting, using a gain variable amplifier capable of switching the gain to be given to the image signal,
Comparing an integrated value obtained by integrating the dose detection signal a predetermined number of times with a preset threshold value,
If the integrated value is greater than the threshold, set the gain variable gain amplifier low,
The radiographic image according to any one of claims 17 to 19, wherein when the integrated value is equal to or less than the threshold value, the gain of the variable gain amplifier is set higher than when the integrated value is larger than the threshold value. Detection device.   21. The radiological image detection apparatus according to claim 20, wherein, in the gain setting, the integrated value of the daylighting field set in the imaging region is obtained.   The radiological image detection apparatus according to claim 20 or 21, wherein, in the gain setting, when the dose detection signal is output to the memory, the gain of the gain variable amplifier is set to a minimum value.   The radiographic image detection device according to any one of claims 1 to 22, wherein the FPD is an electronic cassette housed in a portable housing. A plurality of pixels for accumulating charges according to the radiation dose emitted from the radiation source, a dose detection sensor for detecting charges according to the radiation dose, and a plurality of signal lines provided for each column of the pixels. Yes, the pixel and the dose detection sensor are connected, the charge output from the pixel is output as an image signal according to the driving of the switching element, and the switching element is driven separately from the switching element of the pixel. In response, the FPD that has an imaging region in which a plurality of signal lines that output the charges output from the dose detection sensor as dose detection signals are arranged, and that detects a radiographic image of the subject,
Provided for each of the plurality of signal lines, a plurality of integration amplifiers that integrate charges and convert them into voltage signals, and first and second voltage signals that are temporarily read from the integration amplifier and that are read twice. A pipeline-type signal processing circuit having a two-signal holding unit and executing a cycle of performing a signal input of a voltage signal to one of the first and second signal holding units and a signal output from the other in parallel When,
A memory for storing the image signal and the dose detection signal output as voltage signals from the signal processing circuit;
An operation method of a radiological image detection apparatus comprising the FPD, the signal processing circuit, and a control unit that controls operation timing of the memory,
In the read operation for outputting the image signal to the memory, the integration amplifier starts resetting charge and is reset as the cycle in which signal input and signal output by the first and second signal holding units are performed in parallel. Control the controller to repeat a certain normal cycle corresponding to one integration period until
On the other hand, in the dose detection operation for outputting the dose detection signal to the memory, the first cycle and the first cycle are the cycles in which signal input and signal output by the first and second signal holding units are performed in parallel. A method of operating a radiation image detection apparatus, wherein two types of shorter second cycles are performed and the control unit executes control to sandwich the second cycle at least once between the two first cycles.

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