JP5775812B2 - 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.

  Patent Documents 1 and 2 describe FPDs in which one pixel is configured with a photodiode (main pixel) having a large opening area and a photodiode (subpixel) having a small opening area, and each can read out stored charges separately. ing. By mixing or adding the accumulated charges from the main pixel and the sub-pixel to obtain an image signal of one pixel, the dynamic range of the pixel can be expanded.

  By the way, in order to minimize the influence of noise on the X-ray image in the FPD, 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 dose of X-rays 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, In the X-ray image detection apparatus, automatic exposure control (AEC: Automatic Exposure Control) for shifting from the accumulation operation to the readout operation 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 3 describes an X-ray diagnostic apparatus 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.

JP 2009-018154 A JP 2009-078143 A Japanese Patent Application Laid-Open No. 07-201490

  It has been proposed that the subpixels described in Patent Documents 1 and 2 are pixels for dose detection. When the sub-pixel is used for dose detection, the stored charge of the sub-pixel does not contribute to image generation, so that the image quality is deteriorated. This problem can be solved once by preparing a configuration for adding and temporarily storing signals of pixels for dose detection as in the invention described in Patent Document 3. However, the above configuration is necessary and the cost is increased. In addition, since charge accumulation is interrupted when reading a signal, it becomes a dead time for the pixel while reading the signal, and if the infinite number of times of reading is added until the threshold value is exceeded, the sum of the signals will be greater than the actual value depending on the number of times of reading. Since it is detected low, the image quality may deteriorate.

  The present invention has been made in view of the above problems, and an object of the present invention is to provide a radiological image detection apparatus and a driving method thereof that can prevent image quality deterioration due to use of pixels for dose detection with a simple configuration and processing. And

  The radiological image detection apparatus according to the present invention accumulates charges according to the arrival dose of radiation emitted from a radiation source, and allows a plurality of adjacent pixels to read main charges and sub-pixels separately from the charges. A detection panel in which pixels including composite pixels that are regarded as one pixel are arranged, and an integrated value of a dose detection signal that is a voltage signal based on the accumulated charge of the sub-pixel and a preset threshold value Control means for controlling the operation of the detection panel according to a comparison result, and during radiation irradiation, charges are accumulated in pixels other than the composite pixel and the main pixel, and dose detection signals are periodically output from the sub-pixels. Read-out control means for reading out an image signal that is a voltage signal based on accumulated charges from all the pixels of the composite pixel and the pixels other than the composite pixel after completion of radiation irradiation, and the image signal of the composite pixel First correction means for correcting to be the same as that output from a pixel other than the composite pixel, and an image signal of the pixel other than the composite pixel and an image signal of the composite pixel corrected by the first correction means A radiographic image is generated based on the above.

  The first correction unit multiplies the image signal of the main pixel by a ratio of the total aperture area of the composite pixel, which is the sum of the main pixel and the sub-pixel, and the aperture area of the main pixel. Alternatively, the image signal of the main pixel is multiplied by the ratio of the total output of the composite pixel obtained by combining the main pixel and the sub-pixel with respect to radiation irradiated under a certain condition and the output of the main pixel.

  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 first correction unit is configured to read the image signal of the main pixel when the integration amplifier is used. The ratio of the total aperture area of the composite pixel combined with the main pixel and the sub-pixel and the aperture area of the main pixel, or the main pixel and the sub-pixel with respect to radiation irradiated under a certain condition. A value corresponding to the ratio of the combined output of the combined pixels and the output of the main pixel is set.

  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 from the detection panel by irradiating with radiation in the absence of the subject, the first The correction means is configured to determine the ratio of the total aperture area of the composite pixel including the main pixel and the sub pixel to the aperture area of the main pixel, or the main pixel and the sub pixel with respect to radiation irradiated under a certain condition. The combined ratio of the total output of the composite pixel and the output of the main pixel is incorporated into the composite pixel portion of the sensitivity correction data.

  It is preferable that selection means for selecting the composite pixel that outputs a dose detection signal from the sub-pixel is provided. The first correction unit operates only on the image signal of the main pixel of the composite pixel selected by the selection unit. The control unit treats the sub-pixels of the composite pixel not selected by the selection unit in the same manner as the pixels other than the composite pixel and the main pixel.

  It is preferable to prepare first sensitivity correction data that does not incorporate the ratio of the opening area or the output ratio and second sensitivity correction data that incorporates the ratio. The first correction unit applies the first sensitivity correction data to an image signal of the main pixel of the composite pixel that has not been selected by the selection unit, and the first correction unit selects the composite pixel selected by the selection unit. The second sensitivity correction data is applied to the image signal of the main pixel.

  The selection unit receives manual input of the composite pixel that outputs a dose detection signal from the sub-pixel. In a case where a storage unit that stores the composite pixel that outputs a dose detection signal from the sub-pixel 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 composite pixel existing in the identified region is 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 control unit accumulates electric charges in the sub-pixels of the composite pixel not selected by the selection unit after the region is specified by the selection unit. In this case, the second correction unit corrects the image signal of the sub-pixel of the composite pixel not selected by the selection unit to be the same as when the charge is accumulated immediately after the start of radiation irradiation. It is preferable to provide.

  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. The communication means transmits an irradiation stop signal for stopping irradiation of radiation from the radiation source to the control apparatus of the radiation source when the automatic exposure control means determines that the cumulative value of the arrival dose has reached a target value. To do. In this case, the composite pixel present in the region of interest that is most noticeable at the time of diagnosis is selected as a composite pixel that outputs a dose detection signal from the sub-pixel.

  Alternatively, irradiation start and / or end detection is performed in which the integrated value of the dose detection signal is compared with a preset threshold value, and based on the comparison result, the start and / or end of irradiation of radiation from the radiation source is detected. Preferably means are provided. In this case, the composite pixel existing in the blank region where the detection panel is directly irradiated with radiation without passing through the subject is selected as a composite pixel that outputs a dose detection signal from the sub-pixel.

  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 the gain of the integration amplifier when the image signal is read based on the result. In this case, the composite pixel present in the region of interest that is most noticeable at the time of diagnosis is selected as a composite pixel that outputs a dose detection signal from the sub-pixel.

  The main pixel has an opening area larger than that of the sub-pixel, and the total opening area of the composite pixel including the main pixel and the sub-pixel is substantially equal to that of pixels other than the composite pixel. Alternatively, the composite pixel is composed of a plurality of pixels having the same opening area.

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

  In addition, according to the driving method of the radiation image detection apparatus of the present invention, charges corresponding to the arrival dose of radiation irradiated from a radiation source can be accumulated, and charges can be read out separately to a plurality of adjacent pixels to signal lines. A method for driving a radiation image detection apparatus including a detection panel in which pixels including a composite pixel that is regarded as one pixel divided into main pixels and sub-pixels are arranged, and a voltage signal based on accumulated charges of the sub-pixels Is a control step of controlling the operation of the detection panel according to a comparison result between the integrated value of the dose detection signal and a preset threshold value, and during irradiation with radiation, the pixels other than the composite pixel and the main pixel are controlled. A voltage based on the accumulated charge from all the pixels other than the composite pixel and the composite pixel after the radiation detection is completed and the dose detection signal is periodically read out from the sub-pixel after the radiation irradiation is completed. A control step for reading out an image signal as a signal, a correction step for correcting the image signal of the composite pixel to be the same as that output from a pixel other than the composite pixel, an image signal of a pixel other than the composite pixel, and the And an image generation step of generating a radiation image based on the image signal of the composite pixel corrected by the first correction means.

  In the present invention, when a sub-pixel of a composite pixel is used for dose detection, correction is performed so that the image signal of the composite pixel is the same as that output from a pixel other than the composite pixel, and the corrected image signal and composite pixel Since the radiation image is generated from the image signals of the pixels other than those, it is possible to prevent image quality deterioration due to the use of the pixels for dose detection with a simple configuration and processing.

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 the example which comprised the composite pixel with the some pixel with the same opening area.

  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 sub-pixel 58b (see FIG. 3) of the composite pixel 58 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 corresponds to a pixel other than the composite pixel, and includes a photodiode 38 that is a photoelectric conversion element that generates charges (electron-hole pairs) upon incidence of visible light, and a capacitor that accumulates charges generated by the photodiode 38 (see FIG. And a thin film transistor (TFT) 39.

  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 plurality of composite pixels 58 in the same imaging surface 37 in addition to the normal pixels 36 to which the TFTs 39 driven by the gate drivers 42 and the scanning lines 40 are connected as described above. The composite pixel 58 occupies about several ppm to several percent of the pixel 36 in the imaging surface 37. The composite pixel 58 includes two photodiodes 59a and 59b and TFTs 60a and 60b. The two photodiodes 59a and 59b have a size such that the opening area is almost the same as the photodiode 38 of the pixel 36, and the photodiode 59a is formed larger than the photodiode 59b. The TFTs 60 a and 60 b have drain electrodes connected to the photodiodes 59 a and 59 b, respectively, and source electrodes connected to the signal line 41. On the other hand, the TFT 60a is connected to the scanning line 40 in the same manner as the TFT 39 while the TFT 60b is connected to a scanning line 61 different from the scanning line 40. The scanning line 61 is connected to a gate driver 62 different from the gate driver 42, and the TFT 60 b is turned on by gate pulses g 1 to gn from the gate driver 62. Hereinafter, a set of the photodiode 59a and the TFT 60a is referred to as a main pixel 58a, and a set of the photodiode 59b and the TFT 60b is referred to as a sub-pixel 58b.

  The sub-pixel 58b can read the accumulated charge from the signal line 41 independently of the pixel 36 and the main pixel 58a. In the reset operation and the read operation, the gate pulse of the same row is emitted from the gate driver 62 in synchronization with the gate pulse of the gate driver 42. As a result, the sum of the charges accumulated in the main pixel 58a and the sub-pixel 58b flows to the signal line 41. The sub-pixel 58b is a pixel that is used to detect an X-ray arrival dose on the imaging surface 37, and functions as an AEC sensor.

  As shown in FIG. 4, the composite pixels 58 follow a waveform trajectory 63 indicated by a dotted line symmetrical with respect to the center of the imaging surface 37 so as to be evenly distributed in the imaging surface 37 without being locally biased in the imaging surface 37. Is provided. The composite pixels 58 are provided one by one in the column of the pixels 36 to which the same signal line 41 is connected, and the column in which the composite pixel 58 is provided is provided by sandwiching, for example, two to three columns in which the composite pixel 58 is not provided. It is done. The position of the composite pixel 58 is known at the time of manufacturing the FPD 35, and the FPD 35 stores the positions (coordinates) of all the composite pixels 58 in a nonvolatile memory (not shown) in advance. It should be noted that a plurality of composite pixels 58 may be provided in the same column, or the composite pixels 58 may be arranged in a concentrated manner on the contrary to the present embodiment, and the arrangement of the composite pixels 58 can be changed as appropriate. For example, in a mammography apparatus that captures an image of the breast, the composite pixels 58 may be arranged concentrated on the chest wall side.

  When a gate pulse is generated from the gate driver 62 and the TFT 60 b is turned on, the signal charge generated in the sub-pixel 58 b of the composite pixel 58 is read out to the signal line 41. Since the pixel 36 and the main pixel 58a are separate drive sources from the pixel 36 and the main pixel 58a, the sub-pixel 58b is not connected even when the pixel 36 and the main pixel 58a in the same column have the TFTs 39 and 60a turned off and the signal charge is accumulated. Can be read out. At this time, the charge generated in the sub-pixel 58b flows into the capacitor 46b of the integrating amplifier 46 on the signal line 41 to which the sub-pixel 58b is connected. During the accumulation operation of the pixel 36 and the main pixel 58a, the electric charge from the sub pixel 58b accumulated in the integrating amplifier 46 with the TFT 60b 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 called a gain correction circuit, and corrects variations in sensitivity of the photodiode 38 of each pixel 36 including the composite 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. The offset correction image and sensitivity correction data of the composite pixel 58 are not divided into the main pixel 58a and the sub-pixel 58b, but the composite pixel 58 is regarded as one pixel and assigned to each composite pixel 58 one by one. Yes. 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 multiplication circuit 68 in addition to the various image processing circuits 65 to 67. The multiplication circuit 68 adds the main pixel 58 a and the sub-pixel 58 b to the image signal of the main pixel 58 a of the composite pixel 58 in the lighting field sent from the sensitivity correction circuit 66 based on the shooting field lighting information sent from the console 14. Multiplying the ratio of the total aperture area of the combined composite pixel 58 and the aperture area of the main pixel 58a. More specifically, when the opening areas Sa and Sb of the main pixel 58a and the sub-pixel 58b are set, (Sa + Sb) / Sa is multiplied. The multiplication circuit 68 does not operate on the image signal of the composite pixel 58 outside the lighting field.

  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 sub-pixel 58b 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 integrating circuit 75 integrates the average value, maximum value, mode value, or total value of the dose detection signals from the sub-pixels 58b 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 is compared with the irradiation stop threshold given from the threshold generation circuit 77 at an appropriate timing. When the integrated value reaches the threshold value, the comparison circuit 77 outputs an irradiation stop signal.

  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, and an irradiation stop threshold value for determining the X-ray irradiation stop compared with the integrated value of the dose detection signal of the sub-pixel 58b. 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 composite 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 coordinate corresponds to the position in the imaging surface 37 of the pixel 36 including the composite pixel 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 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 composite 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, a dose detection operation for turning on the TFT 60b and outputting a dose detection signal is started only for the sub-pixel 58b of the composite pixel 58 in the daylighting field set in the imaging condition.

  When the comparison circuit 76 of the AEC unit 52 determines that the integrated value of the dose detection signal has reached the threshold value, the control unit 32 sets the FPD 35 regardless of the pixel 36, the composite pixel 58 outside the lighting field, and the composite pixel 58 inside the lighting field. The operation is shifted from the accumulation operation to the read operation. This completes one shooting. The FPD 35 returns to the standby mode.

  However, in this case, in the composite pixel 58 outside the lighting field, charges corresponding to the X-ray arrival dose are accumulated in both the main pixel 58a and the sub-pixel 58b, and a voltage signal equivalent to that of the normal pixel 36 is output in the readout operation. In the voltage signal output by the reading operation from the composite pixel 58 in the lighting field, the accumulated charge of the sub-pixel 58b is output as a dose detection signal for AEC, so only the charge generated in the main pixel 58a is reflected. . Accordingly, the value is decreased by the amount of the sub-pixel 58b than in the case of the normal pixel 36. Therefore, the multiplication circuit 68 multiplies the image signal output from the main pixel 58a of the composite pixel 58 in the daylighting field by the read operation after the X-ray irradiation is stopped by the multiplication circuit 68 and corrects it. For example, when Sa: Sb = 4: 1 (the opening area of the main pixel 58a is four times that of the sub-pixel 58b), the image signal output from the main pixel 58a in the lighting field is (4 + 1) /4=1.25. Double. The composite pixel 58 in the daylight field as well as the composite pixel 58 in the daylight field can be used to generate an X-ray image.

  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 pressed 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 main pixel 36 and the composite pixel 58 outside the lighting field are detected. The pixel 58a, the sub-pixel 58b, and the main pixel 58a of the composite pixel 58 in the lighting field are shifted from the reset operation to the accumulation operation, and switched to the photographing mode. On the other hand, the sub-pixel 58b of the composite pixel 58 in the daylighting field is shifted to the dose detection operation with the TFT 60b turned on (S12).

  In the case of the pixel 36, the composite pixel 58 outside the lighting field, and the main pixel 58a of the composite pixel 58 inside the lighting field, the charges generated by the X-ray irradiation by the X-ray source 10 are accumulated in the photodiodes 38, 59a, and 59b. In the case of the sub-pixel 58b of the composite pixel 58 in the lighting field, it flows into the integrating amplifier 46 through the signal line 41, and is output as a dose detection signal from the integrating amplifier 46 to the A / D 49 and AEC unit 52 at a predetermined sampling period.

  The dose detection signal from the sub-pixel 58b 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 generation circuit 77 (S14), and stops irradiation when the integrated value reaches the threshold value (YES in S15). Output a signal. 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 (S16).

  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, based on the information on the daylighting field, the image signal corresponding to the composite pixel 58 in the daylighting field is multiplied by (Sa + Sb) / Sa by the multiplication circuit 68 and corrected. The image signal of the composite pixel 58 outside the lighting field passes through the multiplication circuit 68. Thus, one X-ray image is generated (S17). 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 output of the sub-pixel 58b is used for AEC, the image signal of the composite pixel 58 is corrected by multiplying the area ratio and used for generating an X-ray image. It is possible to prevent image quality degradation caused by using the output of the sub-pixel 58b for AEC. Since the correction is completed simply by multiplying the image signal by the area ratio, a large-scale configuration is not required and it is simple.

  Since charges are also accumulated in the sub-pixels 58b of the composite pixels 58 outside the daylighting field for use in image generation, it can be said that there is almost no deterioration in image quality due to the composite pixels 58 being arranged.

  Since the operation of the multiplication circuit 68 is dynamically changed in accordance with the information on the daylighting field, it is possible to correct each of the image signals from the composite field 58 in the daylighting field and outside, so It is possible to generate an X-ray image with uniform image quality with little difference.

  Instead of the area ratio of the above embodiment, the image signal of the main pixel 58a in the lighting field is multiplied by the ratio of the overall output of the composite pixel 58 including the main pixel 58a and the sub-pixel 58b and the output of the main pixel 58a. It may be corrected. In the same manner as when the sensitivity correction data is acquired, X-rays of a predetermined dose are irradiated in the absence of the subject, and the output ratio based on the output voltage signals Va and Vb of the main pixel 58a and the sub-pixel 58b at that time. Is calculated. Specifically, (Va + Vb) / Va is multiplied by the image signal of the main pixel 58a in the lighting field. The output ratio may be the average value or the mode value of all the composite pixels 58, or the output ratio is calculated for each composite pixel 58 and the result is stocked together with the position information of the composite pixel 58 to obtain information on the lighting field. And the output ratio corresponding to the composite pixel 58 may be read out.

  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 daylighting field selection circuit 80 selects which composite pixel 58 of all composite pixels 58 in the imaging plane 37 uses the dose detection signal of the sub-pixel 58b of the composite pixel 58 for AEC. It is undefined which subpixel 58b is used for AEC before the start of imaging, and after receiving the irradiation start signal, the subpixels 58b of all the composite pixels 58 shift to the dose detection operation. The dose detection signal of the sub-pixel 58b of all the composite 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 sub-pixel 58b existing in the irradiation field among the dose detection signals output from the sub-pixel 58b. In other words, the sub-pixel 58b existing in the non-irradiation field portion is excluded from the lighting field candidates.

  The subject region specifying unit 82 among the dose detection signals output from the sub-pixel 58b existing in the irradiation field, the dose detection signal of the sub-pixel 58b existing in the subject region irradiated with the X-ray transmitted through the subject. To pick up. That is, the sub-pixel 58b existing in the blank region where the X-ray is directly incident without passing through the subject is 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 subpixel 58b whose dose detection signal is equal to or greater than the threshold th1 as the subpixel 58b existing in the unexposed region, and specifies the other as existing in the subject region. Alternatively, the sub-pixel 58b 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 sub-pixel 58b existing in the background region. In this way, the irradiation field specifying unit 81 and the subject area specifying unit 82 exclude the sub-pixels 58b existing in the non-irradiation field and the unexposed area from the sub-pixels 58b of all the composite pixels 58 from the candidates for the lighting field.

  The lighting field region specifying unit 83 compares the threshold th2 with the magnitude of the dose detection signal from the sub-pixel 58b 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 sub-pixel 58b in which the dose detection signal falls within a predetermined range (threshold value th2 ± α) centering on the threshold value th2 as the sub-pixel 58b existing in the desired lighting field region. The identification of these irradiation field, subject area, and lighting field area (exclusion of the non-irradiation field and the blank area from the sampling field candidates) is performed on the dose detection signal from the sub-pixel 58b sent in the dose detection operation. For real time. The lighting field selection circuit 80 finally outputs to the integrating circuit 75 the dose detection signal of the sub-pixel 58b specified by each of the specifying units 81, 82, 83 as existing in the lighting field region.

  In FIG. 10, when the chest is imaged, the dose detection signals of the sub-pixels 58b in the non-irradiated 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 sub-pixel 58b in the unexposed region is excluded by the subject region specifying unit 82, and finally the sub-pixel 58b existing in the left and right lung fields as the lighting 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. The period Ta during which the X-ray arrival dose increases or the period Tb 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 Ta and Tb, it is possible to specify each region without any problem in any period.

  When specifying each area during the period Ta when the arrival dose is increasing, the value of the dose detection signal is relatively small, and it is easily affected by noise. However, each area 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.

  When each region is specified in the period Tb 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, the output of the dose detection signal is more stable and the S / N is better in the period Tb than in the period Ta, so that the reliability of the identification result of each region can be improved. .

  In the case where each region is specified in the period Ta and the case where each region is specified in the period Tb, a threshold value th0 for specifying the non-irradiation field region, a threshold value th1 for specifying the background missing 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 the two-dot chain line, the case where each region is specified in the period Ta is set as compared to the case where each region is specified in the period Tb shown by the one-dot chain line. The value is low.

  The sub-pixels 58b in the areas other than those identified as existing in the daylight field area by the daylight field selection circuit 80 continue the dose detection operation even after each area is identified. Therefore, in this case, the image signal from the main pixel 58a of all the composite pixels 58 is multiplied by the area ratio or the output ratio by the multiplication circuit 68 regardless of the inside or outside of the lighting field and corrected.

  Alternatively, as shown in FIG. 12, after specifying each region, the TFT 60b of the sub-pixel 58b outside the daylighting field may be turned off to immediately shift from the dose detection operation to the accumulation operation. When each region is specified in the period Ta, the charge loss due to the dose detection operation of the sub-pixel 58b outside the lighting field is insignificant, and therefore, equivalent to the composite pixel 58 outside the lighting field of the above embodiment without any particular correction. You may handle it. Of course, the accumulated charges of the main pixel 58a and the sub-pixel 58b are read out at different timings, and the charge accumulation time Tc of the image signal of the pixel 36 and the main pixel 58a and each region are specified to change from the dose detection operation to the accumulation operation. The control unit 32 measures the time until the read operation starts, that is, the charge accumulation time Td of the image signal of the subpixel 58b outside the lighting field, and the multiplication circuit 68 converts the Tc into the image signal of the subpixel 58b outside the lighting field. The charge loss may be strictly corrected by multiplying / Td. When each region is specified in the period Tb, the charge loss becomes a magnitude that cannot be ignored. Therefore, the multiplication circuit 68 multiplies the voltage signal of the sub-pixel 58b outside the lighting field by Tc / Td and corrects it. Then, the corrected image signal of the sub-pixel 58b and the image signal of the main pixel 58a are added to obtain the image signal of the composite pixel 58.

  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 68 multiplies the image signal of the main pixel 58a in the lighting field by the area ratio or the output ratio. It is the same. Therefore, the coefficient of sensitivity correction data corresponding to the composite pixel 58 in the lighting field may be multiplied by the area ratio or the output ratio, and the sensitivity correction and the correction by the multiplication circuit 68 of the above embodiment may be completed at the same time. In this way, the multiplication circuit 68 is not necessary, and the cost can be further reduced.

  Further, there are two types of sensitivity correction data obtained by multiplying the coefficient of the composite pixel 58 by the area ratio or output ratio and second sensitivity correction data by multiplying the coefficient of the composite pixel 58 by the area ratio or output ratio. Depending on the information of the daylighting field, processing may be performed such that the composite pixel 58 outside the daylighting field is switched to the first sensitivity correction data when it is within the daylighting field, and the second sensitivity correction data is switched.

  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 imaging conditions are set on the console 14, the pixel 36 is shifted from the reset operation to the accumulation operation, and the sub-pixel 58b of the composite pixel 58 is shifted from the reset operation to the dose detection operation, respectively. Detection of the dose detection signal is started by the detection unit 85. 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 accumulation operation is started in the pixel 36 and the main pixel 58a. The sub-pixel 58b continues the dose detection 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 sub-pixel 58b of the composite pixel 58 existing in the unexposed region for the determination of the X-ray irradiation start. 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 comparison result between the dose detection signal and the threshold value. 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.

  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.

  As in 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 sub-pixels 58b of the composite pixel 58 in the lighting field a predetermined number of times, and the integration The value is compared with a preset 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 integral amplifier, the gain setting unit may adjust the gain of the integral amplifier to perform correction by the multiplication circuit 68 of the above embodiment instead. Specifically, when the charge from the main pixel 58a of the composite pixel 58 in the lighting field is integrated by the integration amplifier during the read operation, the gain of the integration amplifier is adjusted so as to be a gain corresponding to the area ratio or the output ratio. Similarly to the case where the area ratio or the output ratio is incorporated in the coefficient of the sensitivity correction data, the cost can be reduced by omitting the multiplication circuit 68.

  When each region is specified using the lighting field selection circuit 80 in FIG. 10, the TFT 60b of the sub-pixel 58b outside the lighting field is turned off to immediately shift from the dose detection operation to the accumulation operation as shown in FIG. Instead of multiplying the image signal of the subpixel 58b outside the lighting field by Tc / Td, the gain of the variable gain type integration amplifier may be set to a value corresponding to Tc / Td. Further, as described in paragraph [0098], Tc / Td may be incorporated into the coefficient of the sensitivity correction data.

  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 composite pixels 58 are treated as normal pixels 36. In this case, naturally, the multiplication circuit 68 does not operate.

  In the above embodiment, the sub-pixel 58b of the composite pixel 58 is illustrated as a pixel used for AEC or the like. However, the present invention is also applied to a configuration in which a plurality of adjacent pixels having the same aperture area are binned (pixel addition). It can be applied in the same manner as the form. For example, as shown in FIG. 15, binning is performed by regarding nine adjacent pixels 100 as one pixel, and the six pixels 100 a indicated by a thick frame are used for image detection in the same manner as the main pixel 58 a of the above embodiment. When the three pixels 100b indicated by hatching are used for AEC or the like as in the sub-pixel 58b, the image signals of the six pixels 100a are corrected by 1.5 times (in the case of area ratio).

  By performing binning, the data volume of image data handled in subsequent image processing can be greatly reduced, and the processing speed can be increased. In addition, since the plurality of pixels are regarded as one pixel, the apparent sensitivity of the FPD is also improved. For this reason, it is suitable for the moving image photography etc. which image | photograph several times continuously with low dose.

  Note that, for binning, hardware binning for controlling the operation of an FPD (gate driver) and reading and adding the charges of a plurality of adjacent pixels to a signal line at a time as in the above embodiment, and an image signal after digital conversion However, the present invention can be applied to either of them.

  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 FPD
36, 100, 100a, 100b Pixel 39, 60a, 60b TFT
40, 61 Scan line 41 Signal line 42, 62 Gate driver 46, 90 Integration amplifier 52 AEC unit 58 Composite pixel 58a, 58b Main pixel, sub pixel 66 Sensitivity correction circuit 68 Multiplication circuit 76 Comparison circuit 78 Irradiation signal I / F
80 Lighting field selection circuit 85 Irradiation start detection unit 91 Gain setting unit

Claims (24)

Charges corresponding to the radiation dose irradiated from the radiation source are accumulated, and a plurality of adjacent pixels are divided into main pixels and sub-pixels that can read out charges to the signal line separately, and are regarded as one pixel. A detection panel in which pixels including composite pixels made 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 sub-pixel and a preset threshold value;
During radiation irradiation, charges are accumulated in pixels other than the composite pixel and the main pixel, and dose detection signals are periodically read out from the sub-pixels. After the radiation irradiation is completed, the pixels other than the composite pixel and the composite pixel are read out. Control means for reading out an image signal that is a voltage signal based on accumulated charges from all pixels;
First correction means for correcting the image signal of the composite pixel to be the same as that output from a pixel other than the composite pixel;
A radiographic image detection apparatus that generates a radiographic image based on an image signal of a pixel other than the composite pixel and an image signal of the composite pixel corrected by the first correction unit.   The first correction means includes a ratio of a total aperture area of the composite pixel including the main pixel and the sub-pixel and an aperture area of the main pixel, or the main pixel with respect to radiation irradiated under a certain condition. The radiological image detection apparatus according to claim 1, wherein an image signal of the main pixel is multiplied by a ratio of a total output of the composite pixel including the sub-pixels and an output of the main pixel. A variable gain type integration amplifier that integrates charges input via the signal line and converts them into an analog voltage signal,
The first correction means determines the gain of the integration amplifier when reading the image signal of the main pixel, and is the total opening area of the composite pixel including the main pixel and the sub-pixel and the opening area of the main pixel. A ratio or a value corresponding to a ratio of a total output of the composite pixel in which the main pixel and the sub-pixel are combined with respect to radiation irradiated under a certain condition and an output of the main pixel is set. Item 2. The radiological image detection apparatus according to Item 1. 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, a sensitivity correction means for correcting characteristic variations of each part of the detection panel is provided,
The first correction means includes a ratio of a total aperture area of the composite pixel including the main pixel and the sub-pixel and an aperture area of the main pixel, or the main pixel with respect to radiation irradiated under a certain condition. The radiological image detection apparatus according to claim 1, wherein a ratio of a total output of the composite pixel combined with sub-pixels and an output of the main pixel is woven into the composite pixel portion of the sensitivity correction data. Selecting means for selecting the composite pixel that outputs a dose detection signal from the subpixel;
5. The radiographic image according to claim 1, wherein the first correction unit operates only on an image signal of the main pixel of the composite pixel selected by the selection unit. 6. Detection device.   The radiological image detection apparatus according to claim 5, wherein the control unit treats the sub-pixel of the composite pixel that has not been selected by the selection unit in the same manner as a pixel other than the composite pixel and the main pixel. . 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, a sensitivity correction means for correcting characteristic variations of each part of the detection panel is provided,
The ratio of the total aperture area of the composite pixel combined with the main pixel and the sub pixel to the aperture area of the main pixel, or the composite including the main pixel and the sub pixel with respect to radiation irradiated under a certain condition Prepare first sensitivity correction data that does not incorporate the ratio of the total output of pixels and the output of the main pixel, and second sensitivity correction data that incorporates,
The first correction unit applies the first sensitivity correction data to an image signal of the main pixel of the composite pixel that has not been selected by the selection unit, and the first correction unit selects the composite pixel selected by the selection unit. The radiological image detection apparatus according to claim 5, wherein the second sensitivity correction data is applied to an image signal of a main pixel.   The radiographic image detection apparatus according to claim 5, wherein the selection unit receives manual input of the composite pixel that outputs a dose detection signal from the subpixel. Storage means for storing the composite pixel that outputs a dose detection signal from the sub-pixel 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 according to any one of claims 5 to 7, characterized in that at least one of the regions of interest that is sometimes the most notable is specified, and the composite pixel existing in the specified region is selected. apparatus.   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.   13. The control unit according to claim 10, wherein after the region is specified by the selection unit, charges are accumulated in the sub-pixels of the composite pixel not selected by the selection unit. The radiographic image detection apparatus described.   A second correction unit that corrects the image signal of the sub-pixel of the composite pixel that has not been selected by the selection unit to be the same as when the charge is accumulated immediately after the start of radiation irradiation; The radiographic image detection apparatus according to claim 13. 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;
The communication means transmits an irradiation stop signal for stopping irradiation of radiation from the radiation source to the control apparatus of the radiation source when the automatic exposure control means determines that the cumulative value of the arrival dose has reached a target value. The radiological image detection apparatus according to claim 1, wherein the radiological image detection apparatus is a radiographic image detection apparatus.   The radiological image detection apparatus according to claim 15, wherein the composite pixel existing in a region of interest that is most noticeable at the time of diagnosis is selected as a composite pixel that outputs a dose detection signal from the sub-pixel.   Irradiation start and / or end detection means for comparing the integrated value of the dose detection signal with a preset threshold value and detecting that irradiation of radiation from the radiation source is started and / or terminated based on the comparison result The radiographic image detection apparatus according to claim 1, further comprising:   2. The composite pixel that is present in a blank region where radiation is directly irradiated onto the detection panel without passing through a subject is selected as a composite pixel that outputs a dose detection signal from the sub-pixel. The radiological image detection apparatus according to 17. A variable gain type integration amplifier that integrates charges input via the signal line and converts them into an analog voltage signal;
2. A gain setting unit that compares an integrated value of the dose detection signal with a preset threshold value and sets a gain of the integration amplifier when an image signal is read based on the comparison result. The radiological image detection apparatus as described in any one of thru | or 18.   The radiological image detection apparatus according to claim 19, wherein the composite pixel existing in a region of interest most noticeable at the time of diagnosis is selected as a composite pixel that outputs a dose detection signal from the sub-pixel.   The main pixel has an opening area larger than that of the sub-pixel, and a total opening area of the composite pixel including the main pixel and the sub-pixel is substantially equal to that of pixels other than the composite pixel. 21. The radiological image detection apparatus according to any one of 1 to 20.   21. The radiological image detection apparatus according to claim 1, wherein the composite pixel is composed of a plurality of pixels having the same opening area.   The radiographic image detection apparatus according to any one of claims 1 to 22, wherein the detection panel is an electronic cassette housed in a portable housing. Charges corresponding to the radiation dose irradiated from the radiation source are accumulated, and a plurality of adjacent pixels are divided into main pixels and sub-pixels that can read out charges to the signal line separately, and are regarded as one pixel. A method of driving a radiological image detection apparatus including a detection panel in which pixels including composite pixels made 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 sub-pixel and a preset threshold value;
During radiation irradiation, charges are accumulated in pixels other than the composite pixel and the main pixel, and dose detection signals are periodically read out from the sub-pixels. After the radiation irradiation is completed, the pixels other than the composite pixel and the composite pixel are read out. A control step of reading out an image signal which is a voltage signal based on accumulated charges from all pixels;
A correction step for correcting the image signal of the composite pixel to be the same as that output from a pixel other than the composite pixel;
An image generation step of generating a radiation image based on an image signal of a pixel other than the composite pixel and an image signal of the composite pixel corrected by the first correction unit, and driving the radiation image detection apparatus Method.

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