JP5840947B2 - Radiation image detection apparatus and driving method thereof - Google Patents

Description

  The present invention relates to a radiological image detection apparatus and a driving 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 receives an X-ray and captures an X-ray image. The X-ray generator has an X-ray source that irradiates X-rays toward a subject, a radiation source control device that controls driving of the X-ray source, and an irradiation switch for inputting an X-ray irradiation start instruction. ing. An X-ray imaging apparatus is 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 controls driving of the X-ray image detection apparatus and performs various image processing on the X-ray image. have.

  In the recent field of X-ray imaging systems, X-ray image detection apparatuses using a flat panel detector (FPD) as a detection panel are widely used in place of X-ray films and imaging plates (IP). In the FPD, pixels that accumulate signal charges corresponding to the X-ray arrival dose are arranged in a matrix. The FPD accumulates signal charge for each pixel, converts the accumulated signal charge into a voltage signal by a signal processing circuit, detects an X-ray image representing the image information of the subject, and uses this as digital image data Output.

  An electronic cassette (portable X-ray image detection apparatus) in which an FPD is built in a rectangular parallelepiped housing has also been put into practical use. Unlike the type that cannot be removed because the electronic cassette is installed on the photographic stand, it can be used detachably attached to an existing photographic stand for film cassettes and IP cassettes or a dedicated photographic stand. It is used by placing it on a bed or holding it on the subject itself in order to take an image of a particular part. In addition, in order to take pictures of elderly people who are being treated at home or those who are suddenly ill due to accidents, disasters, etc., they may be taken out of hospitals where there is no equipment for taking pictures.

  In the FPD, in order to minimize the influence of noise on the X-ray image, a reset operation for sweeping out unnecessary accumulated charges of pixels due to dark current, residual charges of the previous imaging, or the like is periodically performed. Therefore, in general, in the case of an X-ray image detection apparatus having an FPD, it is necessary to synchronize the X-ray irradiation start timing with the timing to end the reset operation and start the accumulation operation. An interface (I / F) capable of mutual communication is provided in the X-ray image detection apparatus, and the timing at which the radiation source control apparatus starts X-ray irradiation is sent to the X-ray image detection apparatus as an irradiation start signal. Then, processing for shifting to the accumulation operation is performed using the irradiation start signal as a trigger.

  In addition, a dose detection sensor that detects the X-ray dose transmitted through the subject is provided, and when the integrated value (cumulative dose) of the dose detected by the dose detection sensor reaches a preset threshold value, or the cumulative dose becomes the threshold value. When the expected time has elapsed after the expected time has been calculated, automatic exposure control (AEC: Automatic Exposure Control) stops X-ray irradiation from the X-ray source and shifts from the accumulation operation to the readout operation in the X-ray image detection apparatus. Control) is also performed.

  Furthermore, in order to cope with the case where there is no communication function with the radiation source control device, when the integrated value of the dose detected by the dose detection sensor reaches a threshold value, it is determined that X-ray irradiation starts and / or ends, and irradiation starts. There is also an X-ray image detection apparatus having a function of starting an accumulation operation when it is determined that the irradiation operation is completed and shifting from the accumulation operation to a reading operation when it is determined that the irradiation is completed.

  Patent Document 1 describes a radiographic imaging device that uses part of an array pixel as a dose detection sensor. During X-ray irradiation, signals are repeatedly read from the dose detection pixels, and AEC is performed based on the signals. The signal of the pixel for dose detection is added every time it is read out and stored in the memory, and an X-ray image is generated from the stored signal and the signal read from all the pixels after the X-ray irradiation is stopped. The charge accumulated in the pixels for dose detection between the last readout of the signal and the irradiation stop is read at the time of readout of all the pixels, and the signal is also added to the signal stored in the memory to generate an X-ray image. ing.

  In Patent Document 2, as in Patent Document 1, a part of the pixels is used for dose detection, the gate pulse in the row of pixels for dose detection is turned on from the start of X-ray irradiation, and the output is monitored to perform AEC. Is going. The row of pixels for dose detection is first read out when all the pixels are read out, and this is used to generate an X-ray image.

Japanese Patent Application Laid-Open No. 2000-1000059 WO2007 / 037121

  In both Patent Documents 1 and 2, the output of a pixel for dose detection is used not only for AEC but also for generation of an X-ray image, and image quality deterioration due to the use of a part of the pixels for dose detection is prevented. However, in Patent Document 1, a configuration for adding and temporarily storing the signals of the pixels for dose detection is required, which increases the cost. In addition, since charge accumulation is interrupted when reading a signal, there is a dead time for the pixels while the signal is read, and the total sum of the signals added each time is detected lower than the actual value, so the image quality may deteriorate. is there.

  In Patent Document 2, a special readout operation of first reading out a row of pixels for dose detection must be performed, and the control becomes complicated. In addition, the specification of the gate driver must be changed for a special read operation, and the cost becomes high as in Patent Document 1.

  The present invention has been made in view of the above problems, and provides a radiological image detection apparatus and a driving method thereof that can use an output of a pixel for dose detection with a simple configuration and control without waste. With the goal.

  The radiological image detection apparatus of the present invention accumulates charges according to the arrival dose of radiation irradiated from a radiation source, and outputs the charges to a signal line according to driving of the first switching element, and the normal pixels This is a voltage signal based on a detection panel in which a second switching element that is driven separately from the pixel is connected and in which detection pixels for detecting the arrival dose are arranged during irradiation of radiation and an accumulated charge of the detection pixel. Control means for controlling the operation of the detection panel in accordance with a comparison result between the integrated value of the dose detection signal and a preset threshold value, and during normal irradiation, the first switching element is turned off and the normal pixel is turned off. The charge is accumulated and the second switching element is turned on to periodically read out the dose detection signal. When the comparison result is output, stop reading out the dose detection signal. The second switching element is turned off to accumulate charge in the detection pixel, and after the radiation irradiation is completed, the first switching element and the second switching element are turned on to accumulate from the normal pixel and the detection pixel. A control unit that reads an image signal that is a voltage signal based on electric charge; and a correction unit that corrects the image signal of the detection pixel to be the same as that output from the normal pixel. A radiation image is generated based on the image signal of the detection pixel corrected by the correction unit.

  It is preferable that a timing unit that measures the charge accumulation time Ta of the image signal of the normal pixel and the charge accumulation time Tb of the image signal of the detection pixel is provided. The correction means multiplies the image signal of the detection pixel by the ratio Ta / Tb of these charge accumulation times.

  In the case of including a variable gain type integration amplifier that integrates charges input via the signal line and converts them into an analog voltage signal, the correction means gains the integration amplifier when reading the image signal of the detection pixel Is set to a value corresponding to the ratio Ta / Tb of the charge accumulation time of the image signal of the normal pixel and the detection pixel.

  In the case of including sensitivity correction means for correcting characteristic variations of each part of the detection panel based on sensitivity correction data generated based on an image read out from the detection panel by irradiating with radiation in the absence of a subject, the correction means Includes the ratio Ta / Tb of the charge accumulation time of the image signal of the normal pixel and the detection pixel in the detection pixel portion of the sensitivity correction data.

  The correction means includes a charge accumulation time Ta of the image signal of the normal pixel and a charge accumulation time Tb of the image signal of the detection pixel in consideration of a dose gradient from the start of radiation irradiation until the dose is saturated and settles to a constant value. Corrections are made based on the cumulative dose ratio of radiation.

  It is preferable to provide defect correction means for interpolating the image signal of the defective pixel with the image signal of the surrounding normal pixels. When the charge accumulation time Tb of the image signal of the detection pixel is equal to 0, the correction by the correction unit is not performed, the detection pixel is handled in the same manner as the defective pixel, and the image of the detection pixel is processed by the defect correction unit. Interpolate the signal. Note that the case of Tb≈0 refers to the case where the value of the image signal of the detection pixel is extremely small or close to 0 and the S / N of the image signal of the detection pixel does not fall within the allowable range.

  It is preferable to include selection means for selecting the detection pixel that outputs a dose detection signal. The detection pixels not selected by the selection unit are handled in the same manner as the normal pixels.

  The selection means accepts manual input of the detection pixel that outputs a dose detection signal. In the case where a storage unit that stores the detection pixel that outputs the dose detection signal for each imaging region is provided, the imaging region is designated by the selection unit.

  The selection means compares the integrated value of the dose detection signal with a preset threshold value, and based on the comparison result, a blank region where the detection panel is directly irradiated without radiation passing through the subject, or diagnosis At least one of the regions of interest that is sometimes the most notable is identified, and the detection pixels present in the identified region are selected.

  The selection unit specifies a region in a period in which the arrival dose is increasing immediately after radiation irradiation is started from a radiation source. Alternatively, the region may be specified after the reaching dose reaches a certain value.

  The communication means for communicating with the control device of the radiation source, the integrated value of the dose detection signal and a preset threshold value are compared, and whether or not the cumulative value of the arrival dose has reached the target value based on the comparison result It is preferable to include automatic exposure control means for determining. When the automatic exposure control means calculates the time when the cumulative value of the arrival dose is expected to reach a target value, the control means stops reading the dose detection signal and turns off the second switching element. Charge is accumulated in the detection pixel. The communication means transmits an irradiation stop signal for stopping the irradiation of radiation from the radiation source to the radiation source control device when the expected time calculated by the automatic exposure control means has elapsed. In this case, the detection pixel present in the region of interest that is most noticeable at the time of diagnosis is selected as a detection pixel that outputs a dose detection signal.

  Alternatively, it is preferable to provide an irradiation start detection unit that compares the integrated value of the dose detection signal with a preset threshold value and detects that irradiation of radiation from the radiation source is started based on the comparison result. When the irradiation start detection unit detects the start of radiation irradiation, the control unit stops reading the dose detection signal, turns off the second switching element, and accumulates charges in the detection pixel. In this case, the detection pixel existing in the blank region where the detection panel is directly irradiated with radiation without passing through the subject is selected as a detection pixel that outputs a dose detection signal.

  Further, the gain variable integration amplifier that integrates the charges input via the signal line and converts them into an analog voltage signal is compared with the integrated value of the dose detection signal and a preset threshold, and the comparison It is preferable that a gain setting unit is provided for setting a gain of the integration amplifier when an image signal is read from the normal pixel and the detection pixel based on the result. When the gain setting unit finishes setting the gain, the control unit stops reading the dose detection signal, turns off the second switching element, and accumulates charges in the detection pixel. In this case, the detection pixel present in the region of interest that is most noticeable at the time of diagnosis is selected as a detection pixel that outputs a dose detection signal.

  The normal pixel and the detection pixel have the same configuration, such as the size of the photoelectric conversion element, except that the drive source is different. Alternatively, a part of the photoelectric conversion element of the normal pixel may be separated and used as the detection pixel.

  It is preferable that the detection panel is an electronic cassette housed in a portable housing.

  Further, the driving method of the radiological image detection apparatus of the present invention normally stores charges according to the arrival dose of radiation irradiated from the radiation source, and outputs the charges to the signal line according to driving of the first switching element. A driving method of a radiographic image detection apparatus comprising: a detection panel in which a pixel and a second switching element that is driven separately from the normal pixel are connected, and detection pixels for detecting the arrival dose during radiation irradiation are arranged A control step for controlling the operation of the detection panel according to a comparison result between a cumulative value of a dose detection signal, which is a voltage signal based on the accumulated charge of the detection pixel, and a preset threshold value, During irradiation, the first switching element is turned off to accumulate charges in the normal pixels, and the second switching element is turned on to periodically read the dose detection signal. When the comparison result is output, reading of the dose detection signal is stopped, the second switching element is turned off to accumulate charges in the detection pixel, and after the radiation irradiation ends, the first switching element and the A control step of turning on the second switching element and reading out an image signal that is a voltage signal based on accumulated charges from the normal pixel and the detection pixel is the same as when the image signal of the detection pixel is output from the normal pixel And a correction step for correcting the image signal, and an image generation step for generating a radiation image based on the image signal of the normal pixel and the image signal of the detection pixel corrected by the correction means.

  The present invention accumulates electric charge in the detection pixel at the time when the role of dose detection is finished, corrects the image signal obtained thereby to be the same as that output from the normal pixel, Since the radiation image is generated from the image signal of the normal pixel, the output of the pixel for dose detection can be used for image generation without waste with a simple configuration and control.

It is the schematic which shows the structure of a X-ray imaging system. It is a figure which shows the internal structure of a radiation source control apparatus, and the connection relation of a radiation source control apparatus and another apparatus. It is a block diagram which shows the internal structure of an electronic cassette. It is a figure for demonstrating arrangement | positioning of a detection pixel. It is a block diagram which shows the internal structure of the control part of an electronic cassette. It is a block diagram which shows the internal structure of the AEC part and communication part of an electronic cassette. It is a figure which shows the imaging conditions set with a console. It is a figure which shows transition of operation | movement of FPD in X-ray imaging. It is a flowchart which shows the flow of a process of X-ray imaging. It is a block diagram which shows the internal structure of a lighting field selection circuit. It is a figure for demonstrating the timing which specifies each area | region. It is a figure which shows transition of operation | movement of FPD in X-ray imaging. It is a figure which shows the example which provided the irradiation start detection part. It is a figure which shows the structure in the case of setting the gain of the integral amplifier at the time of read-out operation | movement. It is a figure which shows another aspect of FPD.

  In FIG. 1, an X-ray imaging system 2 includes an X-ray source 10 including an X-ray tube that emits X-rays, a radiation source controller 11 that controls the operation of the X-ray source 10, and an X-ray irradiation start. An irradiation switch 12 for instructing, an electronic cassette 13 for detecting an X-ray transmitted through the subject and outputting an X-ray image, a console 14 for controlling the operation of the electronic cassette 13 and image processing of the X-ray image, It has a standing imaging table 15 for imaging the subject in a standing posture and a lying imaging table 16 for imaging 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. In addition to this, a radiation source moving device (not shown) for setting the X-ray source 10 in a desired direction and position is provided.

  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. The irradiation field limiter has, for example, a plurality of lead plates that shield X-rays arranged in a cross pattern, and an irradiation opening that transmits X-rays is formed in the center, and the position of the lead plate is moved. The irradiation field is limited by changing the size of the irradiation opening.

  As shown in FIG. 2, the radiation source control device 11 boosts an input voltage with a transformer to generate a high-voltage tube voltage, and supplies the X-ray source 10 through a high-voltage cable. Main information and signals of the console 14 and the control unit 21 that controls the tube voltage that determines the energy spectrum of the X-rays irradiated by the source 10, the tube current that determines the irradiation amount per unit time, and the X-ray irradiation time. And a communication I / F 22 that mediates transmission and reception.

  An irradiation switch 12, a memory 23, and a touch panel 24 are connected to the control unit 21. The irradiation switch 12 is, for example, a two-stage push switch operated by an operator such as a radiographer, and generates a warm-up start signal for starting the warm-up of the X-ray source 10 by one-stage push. An irradiation start signal for causing the X-ray source 10 to start irradiation is generated. 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 21 starts supplying power from the high voltage generator 20 to the X-ray source 10.

  The memory 23 stores several types of imaging conditions such as tube voltage and tube current in advance. The shooting conditions are manually set by the operator through the touch panel 24. The radiation source control device 11 tries to irradiate X-rays with the tube voltage or tube current irradiation time product of the set imaging conditions. When the AEC detects that a necessary and sufficient dose has been reached, the X-ray irradiation is stopped even if it is less than the tube current irradiation time product (irradiation time) to be irradiated on the radiation source control device 11 side. To function. In order to prevent the X-ray irradiation from ending before the target dose is reached and the AEC is determined to stop irradiation, the tube current irradiation time product (irradiation time) is included in the imaging conditions of the X-ray source 10. But is also possible). Note that the set tube current irradiation time product is preferably a value corresponding to the imaging region.

  The irradiation signal I / F 25 is connected to the electronic cassette 13 when the X-ray irradiation stop timing is defined based on the output of the detection pixel 58 (see FIG. 3) of the electronic cassette 13. In this case, when receiving a warm-up start signal from the irradiation switch 12, the control unit 21 transmits an inquiry signal to the electronic cassette 13 via the irradiation signal I / F 25. When the electronic cassette 13 receives the inquiry signal, the electronic cassette 13 checks whether or not it can shoot, and if it is in a shootable state, transmits an irradiation permission signal. The control unit 21 receives the irradiation permission signal with the irradiation signal I / F 25, and further starts the power supply from the high voltage generator 20 to the X-ray source 10 when receiving the irradiation start signal from the irradiation switch 12. The control unit 21 stops the power supply from the high-voltage generator 20 to the X-ray source 10 when receiving the irradiation stop signal emitted from the electronic cassette 13 by the irradiation signal I / F 25, and X-ray irradiation. Stop.

  The electronic cassette 13 includes an FPD 35 (see FIG. 3) and a portable housing (not shown) that houses the FPD 35. The casing of the electronic cassette 13 is substantially rectangular and has a flat shape, and the plane size is the same size as a film cassette or an IP cassette (also called a CR cassette) (a size conforming to the international standard ISO 4090: 2001). is there. For this reason, it can be attached to an existing photographing stand for a film cassette or an IP 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 15 a and 16 a of the standing imaging stand 15 and the standing imaging stand 16 so that the imaging surface 37 (see FIG. 3) of the FPD 35 is held in a posture facing the X-ray source 10. Set. The electronic cassette 13 can be used alone as it is placed on the bed on which the subject lies, or held by the subject itself, instead of being set on the standing position imaging stand 15 or the supine position imaging stand 16. is there.

  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. Specifically, control such as power on / off of the electronic cassette 13 and mode switching to a standby mode or a photographing mode is performed.

  In addition to being displayed on the display 14b of the console 14, the X-ray image from the electronic cassette 13 is stored in a data storage such as a storage device or memory in the console 14 or an image storage server connected to the console 14 over a network. .

  The console 14 receives an input of an examination order including information such as the patient's sex, age, imaging region, and imaging purpose, and displays the examination 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 radiographs 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). Directions are 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.

  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 for operating 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. 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 includes a TFT active matrix substrate, and includes an imaging surface 37 formed by arranging a plurality of pixels 36 that accumulate charges according to the X-ray arrival dose on the substrate. The plurality of pixels 36 are two-dimensionally arranged in a matrix of n rows (x direction) × m columns (y direction) at a predetermined pitch.

The FPD 35 has a scintillator (phosphor) 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, gadolinium oxysulfide) or the like, and is disposed so as to face the entire imaging surface 37 on which the pixels 36 are arranged. Has been. 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 photodiode 38 that is a photoelectric conversion element that generates charges (electron-hole pairs) upon incidence of visible light, a capacitor (not shown) that accumulates charges generated by the photodiode 38, and a first A thin film transistor (TFT) 39 is provided as a switching element.

  The photodiode 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. In the photodiode 38, a TFT 39 is connected to the lower electrode, and a bias line is connected to the upper electrode. The bias lines are provided for the number of rows (n rows) of the pixels 36 in the imaging surface 37 and are bound to one connection. The connection is connected to the bias power supply. A bias voltage is applied to the upper electrode of the photodiode 38 from the bias power source through the connection and the 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 of which is positive and the other is negative. As a result, charge is accumulated in the capacitor.

  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 photodiode 38. The scanning lines 40 and the signal lines 41 are wired in a grid pattern. The scanning lines 40 are the number of rows of the pixels 36 in the imaging surface 37 (n rows), and the signal lines 41 are the number of columns of the pixels 36 (m columns). Min) each is provided. The scanning line 40 is connected to the gate driver 42, and the signal line 41 is connected to the signal processing circuit 45.

  The gate driver 42 drives the TFT 39 to accumulate a signal charge corresponding to the X-ray arrival dose in the pixel 36, a read (main reading) operation for reading the signal charge from the pixel 36, and a reset (empty reading). ) Make an action. 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 in sequence 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 each row. Turn on. The charge accumulated in the capacitor of the pixel 36 is read out to the signal line 41 and input to the signal processing circuit 45 when the TFT 39 is turned on.

  The signal processing circuit 45 includes an integrating amplifier 46, a CDS circuit (CDS) 47, a multiplexer (MUX) 48, an A / D converter (A / D) 49, and the like. The integrating amplifier 46 is individually connected to each signal line 41. The integrating amplifier 46 includes an operational amplifier 46a and a capacitor 46b connected between the input and output terminals of the operational amplifier 46a, and the signal line 41 is connected to one input terminal of the operational amplifier 46a. The other input terminal of the operational amplifier 46a is connected to the ground (GND). A reset switch 46c is connected in parallel to the capacitor 46b. The integrating amplifier 46 integrates the charges input from the signal line 41, converts them into analog voltage signals V1 to Vm, and outputs them. The MUX 48 is connected to the output terminal of the operational amplifier 46a in each column via the amplifier 50 and the CDS 47. An A / D 49 is connected to the output side of the MUX 48.

  The CDS 47 has a sample and hold circuit, performs correlated double sampling on the output voltage signal of the integration amplifier 46 to remove noise, and holds the output voltage signal of the integration amplifier 46 for a predetermined period (sample hold). ) The MUX 48 selects one CDS 47 by an electronic switch in order from the CDS 47 of each column connected in parallel based on an operation control signal from a shift register (not shown), and outputs voltage signals V1 to V1 output from the selected CDS 47. Vm is serially input to the A / D 49. The A / D 49 converts the input voltage signals V <b> 1 to Vm into digital voltage signals, and outputs the digital voltage signals to the memory 51 or the AEC unit 52 built in the electronic cassette 13. An amplifier may be connected between the MUX 48 and the A / D 49.

  When the voltage signals V1 to Vm for one row are read from the integrating amplifier 46 by the MUX 48, the control unit 32 outputs a reset pulse RST to the integrating amplifier 46 and turns on the reset switch 46c. As a result, the signal charge for one row stored in the capacitor 46b is discharged and reset. After resetting the integrating amplifier 46, the reset switch 46c is turned off again, and one sample hold circuit of the CDS 47 is held after a predetermined time has elapsed, and the kTC noise component of the integrating amplifier 46 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 hold circuit of the CDS 47 after a predetermined time has elapsed. These operations are sequentially repeated to read the signal charges of the pixels 36 in all rows.

  When the reading of all rows is completed, image data representing an X-ray image for one screen is recorded in the memory 51. This image data is read from the memory 51 and output to the console 14 through the communication unit 30. Thus, an X-ray image of the subject is detected.

  Dark charges are generated in the semiconductor layer of the photodiode 38 regardless of whether X-rays are incident. This dark charge is accumulated in the capacitor of the pixel 36 because the 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. While the TFT 39 is on, dark charge flows from the pixel 36 to the capacitor 46 b of the integrating amplifier 46 through the signal line 41. In the reset operation, unlike the read operation, the charge accumulated in the capacitor 46b is not read out by the MUX 48, and the reset pulse RST is output from the control unit 32 in synchronization with the generation of the gate pulses G1 to Gn. 46c is turned on, the charge accumulated in the capacitor 46b is discharged, and the integrating amplifier 46 is reset.

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

  The FPD 35 includes a TFT 57 driven by a gate driver 55 and a scanning line 56 other than the normal pixel 36 in addition to the normal pixel 36 to which the TFT 39 driven by the gate driver 42 and the scanning line 40 is connected as described above. A plurality of detection pixels 58 connected to each other (corresponding to the second switching element) are provided in the same imaging surface 37. The TFT 57 is turned on by gate pulses g 1 to gn from the gate driver 55. The basic configuration of the detection pixel 58 such as the photodiode 38 is exactly the same as that of the pixel 36 except for the driving source, and the stored charge can be read from the signal line 41 independently of the pixel 36. In the reset operation and the read operation, after completing the normal operation of the pixel 36, the gate driver 55 issues gate pulses g1 to gn in the same manner to perform the reset operation or the read operation of the detection pixel 58. Alternatively, in synchronization with the operation of the gate driver 42, the reset operation or the read operation of the pixels 36 and the detection pixels 58 in the same row are simultaneously performed. The detection pixel 58 is a pixel used for detecting the arrival dose of X-rays on the imaging surface 37, and functions as an AEC sensor. The detection pixels 58 occupy about several ppm to several percent of the pixels 36 in the imaging surface 37.

  As shown in FIG. 4, the detection pixels 58 follow a waveform trajectory 59 indicated by a dotted line symmetrical with respect to the center of the imaging surface 37 so that the detection pixels 58 are evenly distributed in the imaging surface 37 without being locally biased in the imaging surface 37. Is provided. The detection pixels 58 are provided one by one in the column of pixels 36 to which the same signal line 41 is connected, and the columns in which the detection pixels 58 are provided are provided by sandwiching, for example, two to three columns in which the detection pixels 58 are not provided. It is done. The position of the detection pixel 58 is known when the FPD 35 is manufactured, and the FPD 35 stores the position (coordinates) of all the detection pixels 58 in a nonvolatile memory (not shown) in advance. In contrast to the present embodiment, the detection pixels 58 may be locally concentrated and the arrangement of the detection pixels 58 can be changed as appropriate. For example, in a mammography apparatus that captures an image of the breast, the detection pixels 58 may be arranged concentrated on the chest wall side.

  When a gate pulse is generated from the gate driver 55 and the TFT 57 is turned on, the signal charge generated in the detection pixel 58 is read out to the signal line 41. Since the pixel 36 is a driving source different from the pixel 36, the pixel 36 in the same column can turn off the TFT 39, and the signal charge of the detection pixel 58 can be read even during the accumulation operation for accumulating the signal charge. . At this time, the charge generated in the detection pixel 58 flows into the capacitor 46b of the integration amplifier 46 on the signal line 41 to which the detection pixel 58 is connected. During the accumulation operation of the pixel 36, the charge from the detection pixel 58 that is accumulated in the integrating amplifier 46 with the TFT 57 turned on is output to the A / D 49 at a predetermined sampling period.

  In FIG. 5, the control unit 32 is provided with circuits 65, 66, and 67 that perform various image processing such as offset correction, sensitivity correction, and defect correction on the X-ray image data in the memory 51. The offset correction circuit 65 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 45 and the imaging environment. To do.

  The sensitivity correction circuit 66 is also referred to as a gain correction circuit, and corrects variations in sensitivity of the photodiode 38 of each pixel 36 including the detection pixel 58, variations in output characteristics of the signal processing circuit 45, and the like. Sensitivity correction is performed based on sensitivity correction data generated based on an image obtained by subtracting the offset correction image from an image obtained by irradiating a predetermined dose of X-rays in the absence of the subject. Sensitivity correction data is shifted from the reference value so that the output of each pixel is uniformly the same by multiplying the X-ray image after offset correction when a predetermined dose of X-rays is irradiated in the absence of the subject. Has a coefficient for correcting each pixel. For example, when the output of the pixel A is the reference 1 but the output of the pixel B is 0.8, the coefficient of the pixel B is 1.25 (1 / 0.8 = 1.25).

  The defect correction circuit 67 linearly interpolates the pixel value of the defective pixel with the pixel values of the surrounding normal pixels based on the defective pixel information attached at the time of shipment.

  The offset correction image and the sensitivity correction data are acquired, for example, at the time of shipment of the electronic cassette 13, or are acquired by a manufacturer's service person at the time of regular maintenance or by an operator at the start time of the hospital, and are recorded in the internal memory of the control unit 32. And read out during correction. Note that the above-described various image processing circuits 65 to 67 may be provided in the console 14 and various image processing may be performed by the console 14.

  The control unit 32 is provided with a timer 68 and a multiplication circuit 69 in addition to the various image processing circuits 65 to 67. The timer 68 measures the charge accumulation time Ta (see FIG. 8) of the image signal of the pixel 36 and the charge accumulation time Tb (see FIG. 8) of the image signal of the detection pixel 58. The multiplication circuit 69 multiplies the image signal of the detection pixel 58 sent from the sensitivity correction circuit 66 by the charge accumulation time Ta / Tb ratio Ta / Tb.

  The AEC unit 52 is driven and controlled by the control unit 32. The AEC unit 52 acquires a digital voltage signal (hereinafter referred to as a dose detection signal) from the signal line 41 to which the detection pixel 58 is connected from the A / D 49, and performs AEC based on the acquired dose detection signal.

  In FIG. 6, the AEC unit 52 includes an integration circuit 75, a comparison circuit 76, and a threshold value generation circuit 77. The integration circuit 75 integrates the average value, the maximum value, the mode value, or the total value of the dose detection signals from the detection pixels 58 in the lighting field. The comparison circuit 76 starts monitoring the integrated value of the dose detection signal from the integration circuit 75 when switching from the standby mode in which the reset operation is repeated to the imaging mode in which the accumulation operation is started. Then, the integrated value and the irradiation stop threshold given from the threshold generation circuit 77 are compared with each other at an appropriate timing, and the time when the accumulated X-ray dose is expected to reach the target value is calculated. When the calculated expected time is reached, the comparison circuit 76 outputs an irradiation stop signal. Although the irradiation stop signal is output after the expected time has elapsed, the predicted time itself may be output to the radiation source control device 11.

  The communication unit 30 is provided with an irradiation signal I / F 78. The irradiation signal I / F 25 of the radiation source control device 11 is connected to the irradiation signal I / F 78. The irradiation signal I / F 78 receives an inquiry signal, transmits an irradiation permission signal in response to the inquiry signal, receives an irradiation start signal, and outputs a comparison circuit 76, that is, an irradiation stop signal.

  As shown in FIG. 7, in the console 14, an imaging condition can be set for each imaging region by the input device 14a. The imaging conditions include a tube voltage, a tube current, a lighting field of the detection pixel 58, and an irradiation stop threshold value for determining the stop of X-ray irradiation compared with the integrated value of the dose detection signal of the detection pixel 58. ing. The imaging condition information is stored in the storage device, and the imaging condition corresponding to the imaging region designated by the input device 14 a is read from the storage device and provided to the electronic cassette 13. 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.

  The daylighting field indicates the region of the detection pixel 58 used for AEC, and a portion that is the region of interest most noticeable at the time of diagnosis and that can stably obtain a dose detection signal is set for each imaging region. For example, when the imaging region is the chest, the left and right lung field portions are set as the daylighting fields as indicated by a and b surrounded by dotted lines in FIG. The daylighting field is represented by xy coordinates. When the daylighting field is rectangular as in this example, for example, two xy coordinates connected by diagonal lines are stored. The xy coordinates correspond to the positions in the imaging surface 37 of the pixels 36 including the detection pixels 58 of the electronic cassette 13. The direction parallel to the scanning line 40 is the x axis and the direction parallel to the signal line 41 is the y axis. The coordinates of the upper left pixel 36 are expressed at the origin (0, 0).

  In FIG. 8, before photographing, the FPD 35 operates in a standby mode in which the reset operation is repeatedly executed regardless of whether the pixel 36 or the detection pixel 58 is used. When the irradiation start signal is received as the irradiation signal I / F 78, the control unit 32 causes the FPD 35 to finish the reset operation, start the accumulation operation, and switch from the standby mode to the imaging mode. However, for only the detection pixels 58 in the daylighting field set in the imaging conditions, a dose detection operation for turning on the TFT 57 and outputting a dose detection signal is started.

  When the comparison circuit 76 of the AEC unit 52 calculates the time when the accumulated X-ray dose is expected to reach the target value, the control unit 32 turns off the TFT 57 of the detection pixel 58 in the lighting field and shifts to the accumulation operation. Let The normal pixel 36 and the detection pixel 58 outside the lighting field are continuously subjected to the accumulation operation. When the expected time is reached and an irradiation stop signal is output from the irradiation signal I / F 78, thereby stopping the X-ray irradiation by the X-ray source 10, the control unit 32 detects the pixel 36, the detection pixel 58 outside the lighting field, Regardless of the detection pixel 58 in the lighting field, the operation of the FPD 35 is shifted from the accumulation operation to the reading operation. This completes one shooting. The FPD 35 returns to the standby mode.

  However, in this case, as indicated by hatching, the voltage signal (image signal) output from the detection pixel 58 in the daylighting field is output at the time Tb after the expected time is calculated and the accumulation operation is performed. Only the charge is reflected, and the accumulated charge generated during the dose detection operation is not reflected because it is used for AEC. Therefore, the value is decreased by the amount of hatching compared to the case of the normal pixel 36. Therefore, after calculating the expected time and shifting to the accumulation operation, the time from when the X-ray irradiation is stopped until the reading operation is started, that is, the charge accumulation time Tb of the image signal of the detection pixel 58 in the lighting field, and the pixel 36 and the charge accumulation time Ta of the image signal of the detection pixel 58 outside the daylighting field is measured by the timer 68 of the control unit 32, and the image signal output from the detection pixel 58 in the daylighting field in the read operation after the X-ray irradiation is stopped The ratio Ta / Tb is multiplied by the multiplication circuit 69 to be corrected. Not only the detection pixels 58 outside the lighting field but also the detection pixels 58 inside the lighting field can be used to generate an X-ray image.

  As can be seen from the X-ray dose graph in FIG. 8, it takes some time from the start of X-ray irradiation until the dose saturates and settles to a constant value, so the cumulative dose at time Ta falls below Ta. The trapezoidal area at the bottom and the accumulated dose at time Tb are rectangular areas with Tb as one side. For this reason, if the image signal is simply multiplied by the ratio Ta / Tb, the accumulated dose at the times Ta and Tb is approximated to a rectangular area, and an error occurs in correction. Therefore, in order to further improve the accuracy of correction, the dose gradient (dose increase rate per unit time) from the start of X-ray irradiation until the dose saturates and settles to a constant value is tabulated for each tube voltage and displayed on the console 14. May be stored in the storage device or the like, and correction may be performed based on this table. Specifically, the trapezoidal area which is the cumulative dose at time Ta and the rectangular area which is the cumulative dose at time Tb are accurately obtained, and the ratio is multiplied by the image signal.

  Next, a procedure for performing X-ray imaging in the X-ray imaging system 2 will be described with reference to a flowchart of FIG. First, the subject is placed at a predetermined position in front of the standing imaging table 15 or placed on the supine imaging table 16, and the electronic cassette 13 set on the imaging table 15, 16 in either the standing position or the lying position. The height and horizontal position of the subject are adjusted to match the position of the subject to be imaged. 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 step 10 (S10), in the standby mode before X-ray imaging, the control unit 32 causes the FPD 35 to repeatedly perform a reset operation. When the irradiation switch 12 is pushed in two steps and an irradiation start signal is output from the radiation source control device 11 and received by the irradiation signal I / F 78 (YES in S11), the pixel 36 and the detection pixel 58 outside the lighting field are reset. The operation is shifted from the operation to the accumulation operation and switched to the photographing mode. On the other hand, the detection pixel 58 in the daylighting field is shifted to the dose detection operation with the TFT 57 turned on (S12).

  X-ray irradiation by the X-ray source 10 is started by pressing the irradiation switch 12 in two steps. The charges generated thereby are accumulated in the photodiode 38 in the case of the pixel 36 and the detection pixel 58 outside the lighting field, and flow into the integrating amplifier 46 through the signal line 41 in the case of the detection pixel 58 in the lighting field, and as a dose detection signal. The signal is output from the integrating amplifier 46 to the A / D 49 and the AEC unit 52 at a predetermined sampling period.

  The dose detection signal from the detection pixel 58 in the daylighting field is output to the integration circuit 75 of the AEC unit 52 and integrated by the integration circuit 75 (S13). The threshold generation circuit 77 generates an irradiation stop threshold given from the console 14 and outputs this to the comparison circuit 76. The comparison circuit 76 compares the integrated value of the dose detection signal from the integration circuit 75 with the irradiation stop threshold value from the threshold value generation circuit 77 (S14), and calculates the time when the accumulated X-ray dose is expected to reach the target value. calculate. When the calculation of the expected time is completed (YES in S15), the detection pixel 58 in the lighting field is shifted to the accumulation operation (S16).

  After reaching the expected time (YES in S17), an irradiation stop signal is output from the comparison circuit 76. The irradiation stop signal is transmitted toward the irradiation signal I / F 25 of the radiation source control device 11 via the irradiation signal I / F 78. Further, the operation of the FPD 35 is shifted from the accumulation operation to the read operation (S18).

  When receiving the irradiation stop signal with the irradiation signal I / F 25, in the radiation source control device 11, the control unit 21 stops the power supply from the high voltage generator 20 to the X-ray source 10, thereby stopping the X-ray irradiation. Is done.

  Various image processing is performed on the X-ray image data in the memory 51 by the various image processing circuits 65 to 67 of the control unit 32. At this time, Ta / Tb is multiplied by the multiplication circuit 69 to correct the image signal corresponding to the detection pixel 58 in the daylighting field. Thus, one X-ray image is generated (S19). The X-ray image is wired or wirelessly transmitted to the console 14 via the communication unit 30, and is displayed on the display 14b for diagnosis.

  As described above, according to the present invention, when the expected time for stopping the X-ray irradiation in the AEC unit 52 is calculated, the TFT 57 of the detection pixel 58 in the lighting field is turned off to shift to the accumulation operation. Since the image signal based on the charge is corrected and used to generate an X-ray image, the output of the detection pixel 58 is added and temporarily stored in the memory, or it is detected with a simple configuration and control without performing a special readout operation. The output of the pixel 58 can be used for image generation without waste.

  Note that the X-ray irradiation time is extremely short in low-dose imaging and the like, and the X-ray irradiation is stopped almost simultaneously with the end of the calculation of the expected time. That is, the charge accumulation time Tb of the image signal of the detection pixel 58 in the lighting field is approximately 0. It may become. In this case, the ratio Ta / Tb increases for a small value of the image signal output from the detection pixel 58 in the daylighting field, but the S / N is bad. Therefore, in such a case, it is preferable that the detection pixel 58 in the daylighting field is treated as a defective pixel and the defect correction circuit 67 interpolates with pixel values of surrounding normal pixels.

  When the image signals of the detection pixels are read and added a plurality of times as in the prior art, the total sum of the image signals is surely lower than the actual image signal in accordance with the number of times of reading, compared to the case of reading them once. In the present invention, since the image signal of the detection pixel 58 is read out by a single read operation, there is no influence of a decrease in the image signal due to the number of read operations. When the charge accumulation time Tb≈0 and the S / N deteriorates, there is no problem because it is dealt with as a defective pixel as described above.

  In the above embodiment, the daylighting field is stored in advance in the imaging conditions, but the daylighting field and other areas may be specified by monitoring the dose detection signal. In this case, the lighting field selection circuit 80 shown in FIG. 10 is provided in the front stage of the integration circuit 75 of the AEC unit 52. The lighting field selection circuit 80 selects which detection pixel 58 to use for the AEC, among all the detection pixels 58 in the imaging surface 37. It is undefined which detection pixel 58 is used for AEC before the start of imaging, and after receiving the irradiation start signal, all the detection pixels 58 shift to the dose detection operation (see FIG. 12). A dose detection signal for all the detection pixels 58 is output to the lighting field selection circuit 80.

  The lighting field selection circuit 80 includes an irradiation field specifying unit 81, a subject region specifying unit 82, and a lighting field region specifying unit 83. The irradiation field specifying unit 81 irradiates X-rays on the imaging surface 37 determined by the irradiation aperture of the irradiation field limiter and the distance SID (Source Image Distance) from the imaging surface 37 to the focal point of the X-ray tube of the X-ray source 10. Identify the field. Specifically, the irradiation field is specified by comparing the dose detection signal with the threshold th0. Since the dose detection signal is almost zero in a non-irradiated field that is not irradiated with X-rays, the threshold th0 is set to a value close to zero. Then, an area where the dose detection signal is equal to or less than the threshold th0 is identified as non-irradiation, and the remaining area is identified as an irradiation field. The irradiation field specifying unit 81 picks up the dose detection signal of the detection pixel 58 existing in the irradiation field among the dose detection signals output from the detection pixel 58. In other words, the detection pixels 58 existing in the non-irradiated field portion are excluded from the lighting field candidates.

  The subject region specifying unit 82 is a dose detection signal of the detection pixel 58 existing in the subject region irradiated with X-rays transmitted through the subject out of the dose detection signals output from the detection pixels 58 existing in the irradiation field. To pick up. In other words, the detection pixels 58 existing in the unexposed area where the X-rays are directly incident without passing through the subject are excluded from the candidates for the lighting field.

The subject region specifying unit 82 specifies the subject region by comparing the dose detection signal with the threshold value th1. The threshold th1 is obtained using, for example, an area dose calculation formula based on an NDD method (Numerical Dose Determination method). The constant determined according to the rectification method (inverter, single phase, three phase) of the high voltage generator 20 is T, the tube voltage correction coefficient is C_kV, and the tube voltage with respect to the thickness of various filters installed in the X-ray source 10 When the correction coefficient is C_Fil, the tube current irradiation time product is mAs, the backscattering correction coefficient is BSF, and the X-ray irradiation field on the imaging surface 37 is AREA, the area dose D is
D = T × C_kV × C_Fil × mAs × (1 / SID) ² × BSF × AREA (1)
Is obtained by calculating. The above T, C_kV, etc. are stored in the storage device of the console 14 in the form of a data table. The values according to the specifications of the X-ray source 10 and the high voltage generator 20 are read from the storage device and the threshold th1 is calculated. Is provided to the electronic cassette 13.

  The subject region specifying unit 82 specifies the detection pixels 58 whose dose detection signal is equal to or greater than the threshold th1 as the detection pixels 58 existing in the unexposed region, and specifies the other pixels existing in the subject region. Alternatively, the detection pixel 58 in which the dose detection signal falls within a predetermined range (threshold th1 ± α) having the threshold value th1 as the center value may be specified as the detection pixel 58 existing in the background region. In this way, the irradiation field specifying unit 81 and the subject region specifying unit 82 exclude the detection pixels 58 existing in the non-irradiation field and the unexposed area from the detection field candidates among all the detection pixels 58.

  The lighting field region specifying unit 83 compares the threshold value th2 with the magnitude of the dose detection signal from the detection pixel 58 existing in the irradiation field and the subject region. The threshold th2 is an X-ray dose that will reach the daylighting field, and is obtained in advance by experiments and simulations based on the body shape of a typical adult male as a subject. The threshold th2 is stored together with the imaging conditions in the storage device of the console 14, for example, and the value is changed according to the set imaging region. The lighting field region specifying unit 83 specifies the detection pixels 58 in which the dose detection signal falls within a predetermined range (threshold value th2 ± α) centering on the threshold value th2 as the detection pixels 58 existing in the desired lighting field region. The identification of these irradiation fields, subject areas, and lighting field areas (exclusion of non-irradiation fields and unexposed areas from the sampling field candidates) is based on the dose detection signal from the detection pixel 58 sent in the dose detection operation. For real time. The daylight selection circuit 80 finally outputs a dose detection signal of the detection pixel 58 specified as existing in the daylight field by each of the specifying units 81, 82, 83 to the integration circuit 75.

  In FIG. 10, when the chest is imaged, the dose detection signals of the detection pixels 58 in the non-irradiation field portions at both ends of the imaging surface 37 are excluded by the irradiation field specifying unit 81, and further between the arm and torso of the subject. The dose detection signal of the detection pixel 58 in the unexposed region is excluded by the subject region specifying unit 82, and finally the detection pixel 58 existing in the left and right lung fields that are the sampling field region is specified by the lighting field region specifying unit 83. It shows a state.

  The timing for specifying the irradiation field, the subject region, and the lighting field region by each of the specifying units 81, 82, 83 of the lighting field selection circuit 80 is as follows. As shown in FIG. Either the period Tc during which the X-ray arrival dose increases or the period Td after the driving of the X-ray source 10 is stabilized and the arrival dose is saturated and settled to a constant value may be used. Since the method of changing the arrival dose in each region is different in the periods Tc and Td, it is possible to specify each region without any problem in any period.

  When specifying each region in the period Tc during which the arrival dose is increasing, the value of the dose detection signal is relatively small, and it is easily affected by noise. However, each region is specified almost simultaneously with the start of X-ray irradiation. The process can be completed, and the calculation of the expected AEC time can proceed smoothly. In addition, charge loss due to prolonged dose detection operation can be suppressed.

  When each region is specified in the period Td after the arrival dose is saturated, the dose detection signal obtained by the previous sampling is temporarily stored and compared with the dose detection signal obtained this time. Then, when the previous and present dose detection signals are equal, it is determined that the arrival dose is saturated, and identification of each region is started. Although it takes time to wait until the arrival dose is saturated, since the output of the dose detection signal is more stable and the S / N is better in the period Td than in the period Tc, the reliability of the identification result of each region can be improved. .

  In the case where each region is specified in the period Tc and the case where each region is specified in the period Td, a threshold value th0 for specifying a non-irradiation field region, a threshold value th1 for specifying a background region, and a lighting field region are set. Naturally, the threshold value th2 for specifying is set to a different value, and as shown by a two-dot chain line, the case where each region is specified in the period Tc is set rather than the case where each region is specified in the period Td shown by a one-dot chain line. The value is low.

  As shown in FIG. 12, the detection pixels 58 in the areas other than those specified as being present in the daylight field area by the daylight selection circuit 80 are immediately shifted from the dose detection operation to the accumulation operation. When each region is specified in the period Tc, the charge loss due to the dose detection operation of the detection pixel 58 outside the daylighting field is negligible, so that it is handled in the same way as the normal pixel 36 without correction by the multiplication circuit 69. Also good. Of course, each timer is identified and the time from the dose detection operation to the accumulation operation and the readout operation is started, that is, the charge accumulation time Tb ′ of the image signal of the detection pixel 58 outside the lighting field is counted by the timer 68, The charge loss may be strictly corrected by multiplying the image signal of the detection pixel 58 outside the lighting field by Ta / Tb ′ by the multiplication circuit 69. When each region is specified in the period Td, the charge loss is not negligible. Therefore, the voltage signal of the detection pixel 58 outside the daylighting field is multiplied by Ta / Tb 'for correction.

  In addition, after specifying each area, the detection pixel 58 outside the lighting field is not immediately shifted from the dose detection operation to the accumulation operation, but at the same timing as the detection pixel 58 in the lighting field (after the calculation of the expected time by the comparison circuit 76). The storage operation may be shifted to. In this way, the correction is completed if the voltage signal of the detection pixel 58 is uniformly multiplied by Ta / Tb, and the time for measuring the time Tb 'by the timer 68 can be saved.

  As described above, since the sensitivity correction is a process of multiplying each pixel value of the X-ray image by a coefficient, the multiplication circuit 69 performs correction of charge loss by multiplying the image signal by Ta / Tb or Ta / Tb ′. The same is true. Therefore, the coefficient of sensitivity correction data corresponding to the detection pixel 58 in the lighting field is multiplied by Ta / Tb (the coefficient corresponding to the detection pixel 58 outside the lighting field may be multiplied by Ta / Tb ′), and sensitivity correction and charge loss correction are performed. May be completed at the same time. In this way, the multiplication circuit 69 is not necessary, and the cost can be further reduced. You may multiply the coefficient of the sensitivity correction data which corresponds to the detection pixel 58 in a lighting field by the exact area ratio of the accumulated dose in time Ta and Tb.

  Some X-ray imaging systems do not have a communication function between the radiation source control device 11 and the electronic cassette 13 as in the above embodiment. In this case, as shown in FIG. 13, an irradiation start detection unit 85 may be provided instead of the AEC unit 52, and the detection unit 85 may detect the start of X-ray irradiation based on the dose detection signal. When detecting the start of X-ray irradiation, when an imaging condition is set on the console 14, the pixel 36 is shifted from the reset operation to the accumulation operation, and the detection pixel 58 is shifted from the reset operation to the dose detection operation. The detection of the dose detection signal is started. The dose detection signals are integrated and compared with a preset irradiation start threshold, and when the integrated value reaches the irradiation start threshold, it is determined that X-ray irradiation starts. When it is determined that the irradiation is started, the detection pixel 58 shifts from the dose detection operation to the accumulation operation as in the above embodiment, and the pixel 36 continues to perform the accumulation operation.

  As can be seen from the above description, the irradiation start detection unit 85 has the same basic configuration as the AEC unit 52 except that the threshold value to be compared with the integrated value of the dose detection signal is different. The irradiation start detection unit 85 uses the dose detection signal from the detection pixel 58 present in the unexposed region for determination of the start of X-ray irradiation. Since the unaccompanied region has a higher X-ray arrival dose than the subject region and the amount of change per unit time in the arrival dose is large, an S / N dose detection signal sufficient to determine the start of X-ray irradiation is provided. It can be obtained in a short time. Therefore, accurate and quick judgment can be made.

  The identification of the blank region may depend on manual input by the operator, or the subject region identification unit 82 in FIG. 10 may be used. Or anyway, you may use the dose detection signal which shows the maximum value for judgment of the X-ray irradiation start. In this case, the time required for the determination can be shortened by not specifying the blank region.

  Instead of or in addition to detecting the start of X-ray irradiation, the end of X-ray irradiation may be detected based on the dose detection signal. In this case, when the end of X-ray irradiation is detected, the control unit 32 shifts the FPD 35 from the accumulation operation to the read operation. However, since the detection pixel 58 continues the dose detection operation until the end of the X-ray irradiation is determined, the dose detection operation is not stopped halfway and the accumulation operation is not performed as in the above embodiment.

  In addition to using the dose detection signal 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. In this case, as shown in FIG. 14, instead of the integrating amplifier 46, a variable gain type integrating amplifier 90 is used.

  In FIG. 14, the integrating amplifier 90 includes an operational amplifier 90 a and a reset switch 90 c as in the integrating amplifier 46. Two capacitors 90b and 90d are connected between the input and output terminals of the operational amplifier 90a, and a gain changeover switch 90e is connected to the capacitor 90d. When the gain switch 90e is on, the output voltage signal V from the integrating amplifier is V = q / (C1 + C2), and when the gain switch 90e is off, V = q / C1. Where q is the accumulated charge, and C1 and C2 are the capacitances of the capacitors 90b and 90d, respectively. Thus, the gain of the integrating amplifier 90 can be changed by switching the gain changeover switch 90e on / off. Here, an example is shown in which two capacitors are connected and the gain is switched in two steps. However, the gain can be changed in two or more steps by connecting two or more capacitors or using a variable capacitance capacitor for the capacitor. It is preferable.

  The gain setting unit 91 is provided in the FPD 35 instead of the AEC unit 52 and the irradiation start detection unit 85. The gain setting unit 91 operates when the FPD 35 starts the accumulation operation, and controls the operation of the gain changeover switch 90e during the read operation. A dose detection signal is periodically input from the signal processing circuit 45 to the gain setting unit 91. The gain setting unit 91 sets the gain of the integrating amplifier 90 to the minimum value when outputting the dose detection signal so that the dose detection signal is not saturated. In this example, the gain changeover switch 90e is turned on.

  Similarly to the case of AEC, the gain setting unit 91 integrates the total value, average value, maximum value, or mode value of the dose detection signals from the detection pixels 58 in the lighting field a predetermined number of times, and sets the integrated value and the preset value in advance. Compare with the threshold value. When the integrated value is larger than the threshold value, the gain setting unit 91 turns on the gain changeover switch 90e during the reading operation. On the other hand, when the accumulated accumulated dose to the portion corresponding to the lighting field on the imaging surface 37 is low and the integrated value is equal to or less than the threshold value, the gain selector switch 90e is turned off during the reading operation to increase the gain of the integrating amplifier 90. More specifically, the gain of the integrating amplifier 90 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 cumulative dose to the portion corresponding to the daylighting field is low as described above, an X-ray image with good image quality with no noticeable noise can be obtained. For this reason, the irradiation dose set to an X-ray source can be suppressed, and the special effect that a patient's exposure dose can be decreased as a result is acquired. It is also possible to stop the X-ray irradiation early by setting the AEC irradiation stop threshold low, and to compensate for the shortage by increasing the gain of the integrating amplifier. Can be reduced.

  In the case of using a variable gain type integration amplifier, the gain setting unit may adjust the gain of the integration amplifier to correct the charge loss by the multiplication circuit 69 of the above embodiment instead. Specifically, the gain of the integration amplifier is adjusted so that the gain corresponding to Ta / Tb is obtained when the charges from the detection pixels 58 in the lighting field are integrated by the integration amplifier during the reading operation. It is also possible to integrate the charges from the detection pixels 58 outside the lighting field with a gain of Ta / Tb ′. As in the case of incorporating Ta / Tb or Ta / Tb 'into the coefficient of the sensitivity correction data, the cost can be reduced by omitting the multiplication circuit 69. Note that the gain of the integrating amplifier may be adjusted so that the gain corresponds to an accurate area ratio of the accumulated dose at the times Ta and Tb.

  X-ray irradiation start or end detection, AEC, and integral amplifier gain setting may be combined. You may comprise so that an operator can set which of these functions is performed. When a selection is made not to execute any function, all detection pixels 58 are treated as normal pixels 36.

  In the above embodiment, the sizes and the like of the pixel 36 and the detection pixel 58 are the same, but a part of the photodiode 38 of the pixel 36 may be used as the detection pixel 101 as in the FPD 100 shown in FIG. As in the above-described embodiment, separately from the TFT 39, the scanning line 40, and the gate driver 42 of the pixel 36, the TFT 102, the scanning line 103, and the gate driver 104 are connected to the detection pixel 101. Can be read from the signal line 41. The driving method is the same as that in the above embodiment, and the detection pixel 101 in the daylighting field moves to an accumulation operation after completing the role of dose detection. However, at the time of the reading operation, the gate pulse is simultaneously applied to the scanning lines 40 and 103 in the row where the detection pixel 101 outside the lighting field is present, and the pixel 36 and the detection pixel 101 are read at the same timing. In this way, an image signal in which the accumulated charges of the pixel 36 and the detection pixel 101 are mixed is obtained. This image signal has almost the same value as the pixel 36 without the detection pixel 101. On the other hand, in the row where the detection pixel 101 is in the daylighting field, the pixel 36 and the detection pixel 101 are read at different timings. Then, after correcting the image signal of the detection pixel 101 in the daylighting field by multiplying by Ta / Tb, it is added to the image signal of the pixel 36 to obtain a final value used for the X-ray image.

  In the above embodiment, the 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 X-ray imaging system 10 X-ray source 11 Radiation source control device 13 Electronic cassette 14 Console 14a Input device 30 Communication unit 32 Control unit 35, 100 FPD
36 pixels 39, 57, 102 TFT
41 Signal Line 46, 90 Integrating Amplifier 52 AEC Unit 58, 101 Detection Pixel 66 Sensitivity Correction Circuit 67 Defect Correction Circuit 68 Timer 69 Multiplying Circuit 80 Lighting Field Selection Circuit 85 Irradiation Start Detection Unit 91 Gain Setting Unit

Claims (22)

A normal pixel that accumulates electric charge according to the arrival dose of radiation emitted from the radiation source, and outputs the electric charge to the signal line according to driving of the first switching element, and a second pixel that is driven separately from the normal pixel A detection panel to which a switching element is connected and detection pixels for detecting the arrival dose during irradiation of radiation are arranged;
Control means for controlling the operation of the detection panel according to a comparison result between an integrated value of a dose detection signal that is a voltage signal based on the accumulated charge of the detection pixel and a preset threshold value;
During radiation irradiation, the first switching element is turned off to accumulate charges in the normal pixels, and the second switching element is turned on to periodically read out the dose detection signal, and perform processing based on the comparison result. When reading is completed , reading of the dose detection signal is stopped, the second switching element is turned off to accumulate charges in the detection pixel, and after the irradiation of radiation, the first switching element and the second switching element are turned on. Control means for turning on and reading out an image signal that is a voltage signal based on accumulated charges from the normal pixel and the detection pixel;
Correction means for correcting so that the image signal of the detection pixel is the same as that output from the normal pixel,
A radiographic image detection apparatus that generates a radiographic image based on the image signal of the normal pixel and the image signal of the detection pixel corrected by the correction unit. A timing means for measuring the charge accumulation time Ta of the image signal of the normal pixel and the charge accumulation time Tb of the image signal of the detection pixel;
The radiological image detection apparatus according to claim 1, wherein the correction unit multiplies the image signal of the detection pixel by a ratio Ta / Tb of the charge accumulation times. Timing means for measuring the charge accumulation time Ta of the image signal of the normal pixel and the charge accumulation time Tb of the image signal of the detection pixel;
A variable gain type integration amplifier that integrates charges input via the signal line and converts them into an analog voltage signal;
The correction means sets the gain of the integration amplifier when reading the image signal of the detection pixel to a value corresponding to the charge accumulation time ratio Ta / Tb of the image signal of the normal pixel and the detection pixel. The radiographic image detection apparatus according to claim 1, wherein Timing means for measuring the charge accumulation time Ta of the image signal of the normal pixel and the charge accumulation time Tb of the image signal of the detection pixel;
Based on sensitivity correction data generated based on an image read out from the detection panel by irradiating radiation in the absence of a subject, a sensitivity correction means for correcting characteristic variations of each part of the detection panel,
2. The radiographic image according to claim 1, wherein the correction unit weaves a ratio Ta / Tb of charge accumulation times of image signals of the normal pixel and the detection pixel into the detection pixel portion of the sensitivity correction data. Detection device.   The correction means includes a charge accumulation time Ta of the image signal of the normal pixel and a charge accumulation time Tb of the image signal of the detection pixel in consideration of a dose gradient from the start of radiation irradiation until the dose is saturated and settles to a constant value. The radiological image detection apparatus according to claim 1, wherein correction is performed based on a ratio of a cumulative dose of radiation in the apparatus. A defect correcting means for interpolating an image signal of a defective pixel with an image signal of a surrounding normal pixel;
When the charge accumulation time Tb of the image signal of the detection pixel is equal to 0, the correction by the correction unit is not performed, the detection pixel is handled in the same manner as the defective pixel, and the image of the detection pixel is processed by the defect correction unit. 6. The radiological image detection apparatus according to claim 2, wherein the signal is interpolated. Selecting means for selecting the detection pixel that outputs a dose detection signal;
The radiographic image detection apparatus according to claim 1, wherein the detection pixels not selected by the selection unit are handled in the same manner as the normal pixels.   The radiographic image detection apparatus according to claim 7, wherein the selection unit receives manual input of the detection pixels that output a dose detection signal. Storage means for storing the detection pixel for outputting a dose detection signal for each imaging region;
The radiographic image detection apparatus according to claim 8, wherein an imaging region is designated by the selection unit.   The selection means compares the integrated value of the dose detection signal with a preset threshold value, and based on the comparison result, a blank region where the detection panel is directly irradiated without radiation passing through the subject, or diagnosis The radiological image detection apparatus according to claim 7, wherein at least one of the regions of interest that are sometimes most noticeable is specified, and the detection pixels existing in the specified region are selected.   The radiographic image detection apparatus according to claim 10, wherein the selection unit specifies a region in a period in which the arrival dose is increasing immediately after radiation irradiation is started from a radiation source.   The radiation imaging system according to claim 10, wherein the selection unit specifies an area after the arrival dose reaches a constant value. Communication means for communicating with the radiation source control device;
An automatic exposure control means for comparing the integrated value of the dose detection signal with a preset threshold value and determining whether or not the cumulative value of the arrival dose has reached a target value based on the comparison result;
When the automatic exposure control means calculates the time when the cumulative value of the arrival dose is expected to reach a target value, the control means stops reading the dose detection signal and turns off the second switching element. Charge is accumulated in the detection pixel;
The said communication means transmits the irradiation stop signal for stopping irradiation of the radiation by a radiation source to the control apparatus of a radiation source, if the estimated time calculated by the said automatic exposure control means passes. The radiographic image detection apparatus according to any one of 12.   The radiological image detection apparatus according to claim 13, wherein the detection pixel present in a region of interest that is most noticeable at the time of diagnosis is selected as a detection pixel that outputs a dose detection signal. Comparing the integrated value of the dose detection signal with a preset threshold value, and comprising an irradiation start detection means for detecting the start of radiation irradiation from the radiation source based on the comparison result,
When the irradiation start detection unit detects the start of radiation irradiation, the control unit stops reading the dose detection signal, turns off the second switching element, and accumulates charges in the detection pixel. The radiographic image detection apparatus as described in any one of Claims 1 thru | or 14.   16. The detection pixel that is present in a blank region where radiation is directly irradiated on the detection panel without passing through a subject is selected as a detection pixel that outputs a dose detection signal. Radiation image detection device. A variable gain type integration amplifier that integrates charges input via the signal line and converts them into an analog voltage signal;
Gain setting means for comparing the integrated value of the dose detection signal with a preset threshold value and setting the gain of the integration amplifier when reading the image signal from the normal pixel and the detection pixel based on the comparison result ,
The control means, when the gain setting means finishes setting the gain, stops reading out the dose detection signal, turns off the second switching element, and accumulates charges in the detection pixel. Item 17. The radiological image detection apparatus according to any one of Items 1 to 16.   The radiological image detection apparatus according to claim 17, wherein the detection pixel present in a region of interest most noticeable at the time of diagnosis is selected as a detection pixel that outputs a dose detection signal.   The radiographic image detection apparatus according to claim 1, wherein the normal pixel and the detection pixel have the same configuration except that driving sources are different.   The radiographic image detection apparatus according to any one of claims 1 to 18, wherein a part of the photoelectric conversion element of the normal pixel is separated and used as the detection pixel.   21. The radiological image detection apparatus according to claim 1, wherein the detection panel is an electronic cassette housed in a portable housing. A normal pixel that accumulates electric charge according to the arrival dose of radiation emitted from the radiation source, and outputs the electric charge to the signal line according to driving of the first switching element, and a second pixel that is driven separately from the normal pixel A driving method of a radiological image detection apparatus comprising a detection panel to which a switching element is connected and detection pixels for detecting the arrival dose during radiation irradiation are arranged,
A control step of controlling the operation of the detection panel according to a comparison result between an integrated value of a dose detection signal that is a voltage signal based on the accumulated charge of the detection pixel and a preset threshold value;
During radiation irradiation, the first switching element is turned off to accumulate charges in the normal pixels, and the second switching element is turned on to periodically read out the dose detection signal, and perform processing based on the comparison result. When reading is completed , reading of the dose detection signal is stopped, the second switching element is turned off to accumulate charges in the detection pixel, and after the irradiation of radiation, the first switching element and the second switching element are turned on. A control step of turning on and reading out an image signal that is a voltage signal based on accumulated charges from the normal pixel and the detection pixel;
A correction step for correcting the image signal of the detection pixel to be the same as that output from the normal pixel;
An image generation step of generating a radiation image based on the image signal of the normal pixel and the image signal of the detection pixel corrected in the correction step .

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