JP2013094454A - Radiographic apparatus, radiographic system, and radiographic method - Google Patents

Abstract

An FPD capable of non-destructive reading is prevented from deteriorating and failing to extend its life.
Non-destructive pixel detection for detecting pixel values of non-destructive pixels from an FPD capable of non-destructively reading an X-ray image and an X-ray image read nondestructively during X-ray irradiation 82, an expected arrival charge amount calculation unit 83 that calculates an expected arrival charge amount of the missing pixel at the end of X-ray irradiation based on the pixel value of the missing pixel, and the like. In addition, when the expected charge amount exceeds the predetermined threshold value, new imaging for controlling the X-ray source so that the signal charge accumulated in the unexposed pixels at the X-ray irradiation end time becomes the predetermined threshold value. An imaging condition calculation unit 84 that calculates the conditions, and a radiation source that feedback-controls the X-ray source so that X-rays based on the new imaging conditions are emitted during X-ray irradiation based on the first imaging conditions And a control device 19.
[Selection] Figure 6

Description

  The present invention relates to a radiation imaging apparatus, a radiation imaging system, and a radiation imaging method for capturing a radiation image of a subject using a radiation image detector.

  In recent years, in the field of radiography, for example, X-ray imaging, an X-ray image detection apparatus using a flat panel detector (hereinafter referred to as FPD) using a semiconductor element as a detector instead of an X-ray film or an imaging plate (IP). (Hereinafter referred to as electronic cassette) has become widespread. The FPD uses an image sensor composed of a so-called solid-state image sensor formed using a semiconductor element, and pixels that accumulate signal charges according to the amount of incident X-rays are arranged in a matrix. The FPD detects an X-ray image representing image information of a subject by converting a signal charge accumulated for each pixel upon incidence of X-rays into a voltage signal. An X-ray image detected by the FPD is output as digital image data.

  As the FPD, in addition to the TFT type in which a pixel including a TFT (Thin Film Transistor) is formed on a glass substrate, a pixel is formed on a silicon substrate by a CMOS manufacturing process as described in Patent Document 1. There is a CMOS type. One of the major differences between the TFT type and the CMOS type is the voltage signal readout method. A TFT type FPD is a method of transferring a signal charge accumulated in a pixel to an integration amplifier via a signal line and reading out a voltage signal corresponding to the signal charge integrated by the integration amplifier. This is a so-called destructive readout method in which the signal charge is emptied. On the other hand, a CMOS type FPD is provided with an amplifier for converting a signal charge into a voltage signal for each pixel, and reads out a voltage signal corresponding to the signal charge while holding the signal charge in the pixel. This is a destructive readout method.

  Of course, in both TFT type and CMOS type FPDs, there is an upper limit to the amount of signal charge accumulated in the pixel, so if the dose of X-rays incident on the pixel is excessive, the pixel is saturated, so-called Overflow occurs. In this case, in the X-ray image, the pixel is crushed black. On the other hand, when the so-called underflow occurs where the X-ray dose is too small, the pixel becomes white in the X-ray image. In X-ray imaging, imaging conditions such as a tube current and an irradiation time that determine the dose irradiated by the X-ray source are set in advance according to the imaging region and physique of the subject so that overflow and underflow do not occur.

  Furthermore, Patent Document 1 discloses pixel values of pixels in a region (a so-called region of interest) in which a region to be observed is copied from an X-ray image obtained by nondestructive readout during X-ray irradiation. Obtain X-ray irradiation immediately when the pixel in the region of interest is likely to be saturated, and control the X-ray source to increase the X-ray irradiation intensity if the pixel value in the region of interest is small Thus, a technique for easily obtaining an X-ray image suitable for observation without overflow or underflow is disclosed.

JP 2004-344249 A

  Incidentally, a CMOS FPD capable of nondestructive reading is formed using a crystalline semiconductor (for example, single crystal silicon). An element formed from a crystalline semiconductor has an order of magnitude higher charge mobility and ON resistance of a switching element than an element formed from an amorphous semiconductor (eg, amorphous silicon). For this reason, comparing a TFT type FPD formed from an amorphous semiconductor with a CMOS type FPD formed from a crystalline semiconductor, the CMOS type FPD is advantageous in increasing the reading speed and the pixel density.

  On the other hand, since a CMOS type FPD is formed of a crystalline semiconductor, it is more susceptible to damage due to radiation than a TFT type FPD formed of an amorphous semiconductor. For this reason, the CMOS type FPD has a problem that it deteriorates earlier than the TFT type FPD and it is difficult to extend the life.

  For example, in the case of a CMOS type FPD formed from single crystal silicon, noise (offset noise, etc.) related to pixels starts to become remarkable from an irradiation dose of about several thousand R, and when the total irradiation dose further accumulates about several thousand R, Deterioration or failure occurs, for example, the gain of an amplifier used for image reading becomes unstable. This is because charge pairs are generated by the incidence of radiation and accumulated at the interface between the single crystal silicon substrate and the insulating film (silicon oxide film, etc.). This is because the bond of other materials is broken and the material properties deteriorate.

  Further, conventionally, a radiation image is obtained by arranging a scintillator and an FPD in this order from the radiation incident surface side, converting the radiation into light by the scintillator, and receiving the light emitted from the scintillator by the FPD with the detection surface facing the front side. Indirect-conversion electronic cassettes with PSS (Penetration Side Sampling) method are known, but in recent years, the arrangement order of scintillators and FPDs has been reversed in order to further improve image quality. 2. Description of the Related Art An indirect conversion electronic cassette of a backside incidence type (ISS: Irradiation Side Sampling) in which the detection surface is directed to the backside is known.

  In the case of the PSS method and the ISS method, the deterioration due to radiation irradiation is not as fast as the TFT type FPD as described above, but in the case of the PSS method, the scintillator is arranged on the front surface of the FPD. The radiation is attenuated by the scintillator, and even when the CMOS type FPD is used, the deterioration of the FPD due to the irradiation of radiation is somewhat suppressed. On the other hand, in the case of the ISS system, there is no member that attenuates incident radiation on the front surface of the FPD, and all incident radiation passes through the FPD and reaches the scintillator. Therefore, when using a CMOS type FPD in an ISS electronic cassette, Deterioration of FPD due to irradiation of selenium is a particularly serious problem.

  Furthermore, a direct conversion FPD that directly converts incident radiation into signal charges using an X-ray conversion layer without using a scintillator is also known. Amorphous selenium (a-Se), which is a typical X-ray conversion layer in a direct conversion FPD, has an atomic number smaller than that of a typical scintillator material, and particularly has little absorption of X-rays on the high energy side. Since it reaches the FPD with a small attenuation, if it is formed in the CMOS type in order to enable non-destructive reading by the direct conversion type FPD, deterioration due to radiation irradiation becomes a particularly serious problem as described above.

  The present invention has been made in view of the above-described problems, and an object of the present invention is to prevent deterioration and failure of an FPD capable of nondestructive readout and to extend the life of a radiographic apparatus using the same. is there.

  The radiographic apparatus of the present invention is a radiographic image detecting means for detecting a radiographic image of a subject by receiving radiation irradiated from a radiation source according to imaging conditions and transmitted through the subject, and is formed using a single crystal semiconductor. A radiation image having a plurality of pixels for accumulating signal charges according to the amount of incident radiation, and capable of nondestructively reading data representing the radiation image based on the amount of accumulation of signal charges from the pixels Detecting means; and a missing pixel detecting means for detecting a pixel value of a missing pixel that the radiation reaches directly without passing through the subject from the radiation image read nondestructively during the irradiation of the radiation; , The pixel value of the missing pixels, the elapsed time from the irradiation start time of the radiation to the acquisition time of the radiation image read nondestructively during the radiation irradiation, and the radiation A predicted arrival charge amount calculating means for calculating the amount of signal charge accumulated in the unexposed pixels at the end of irradiation of radiation as an expected arrival charge amount based on an irradiation end time; and Compared with the above, when the expected amount of charges exceeds the predetermined threshold value, the radiation source is controlled so that the signal charge accumulated in the missing pixels at the radiation irradiation end time becomes the predetermined threshold value. An imaging condition calculation unit for calculating the new imaging condition; and the radiation based on the new imaging condition calculated by the imaging condition calculation unit during the irradiation of the radiation based on the first imaging condition. Radiation imaging apparatus, comprising: radiation source control means for feedback-controlling the radiation source so as to be irradiated.

  It is preferable that the new imaging condition calculated by the imaging condition calculation unit is an imaging condition for reducing the energy of the radiation.

  It is preferable that the new imaging condition calculated by the imaging condition calculation unit is an imaging condition for reducing the radiation dose.

  The new imaging condition calculated by the imaging condition calculation unit is preferably an imaging condition that shortens the irradiation time of the radiation.

  It is preferable that the new imaging condition calculated by the imaging condition calculation means is an imaging condition that limits the radiation field.

  When reading out the radiation image from the radiation detection means, the signal corresponding to the signal charge is amplified and read out while changing the gain of the amplification means having a variable gain and after the irradiation of the radiation. It is preferable that the image synthesizing unit synthesizes a plurality of the radiographic images to generate a synthetic radiographic image for observing the subject.

  Preferably, the image synthesizing unit synthesizes the plurality of radiographic images read while changing the gain of the amplifying unit so that noise at the time of reading the radiographic image is reduced.

  Preferably, the image synthesizing unit synthesizes the plurality of radiographic images read while changing the gain of the amplifying unit so that the dynamic range is expanded.

  It is preferable that the image synthesizing unit generates the synthesized radiographic image by averaging a plurality of the radiographic images read while changing the gain of the amplifying unit.

  The radiation detection means includes a scintillator that converts the radiation into light, and an image sensor that detects the radiation image by photoelectrically converting light emitted from the scintillator, and the image sensor from the radiation incident side. The scintillators are preferably arranged in this order.

  The radiation imaging system of the present invention is a radiation imaging system including a radiation source and a radiation imaging apparatus that receives radiation transmitted from the radiation source according to the imaging conditions and transmitted through the subject, and captures a radiation image of the subject. The radiation imaging apparatus includes a plurality of pixels that are formed using a single crystal semiconductor and accumulate signal charges according to the amount of incident radiation, and the radiation image based on the amount of signal charges accumulated from the pixels. A radiation image detecting means capable of nondestructively reading out data representing the data, and an element from which the radiation reaches directly without passing through the subject from the radiation image read nondestructively during the radiation irradiation. A non-destructive pixel detecting means for detecting a pixel value of the non-existing pixel, and a pixel value of the non-descriptive pixel, which is read in a non-destructive manner during the radiation irradiation from the radiation irradiation start time. Predicted arrival charge amount calculation means for calculating an expected arrival charge amount of the unaccompanied pixel at the end of radiation irradiation based on the elapsed time until the acquired radiation image and the irradiation end time of the radiation The amount of expected electric charge is compared with a predetermined threshold, and when the amount of expected electric charge exceeds the predetermined threshold, the signal charge accumulated in the missing pixels at the radiation irradiation end time becomes the predetermined threshold. An imaging condition calculation unit that calculates a new imaging condition for controlling the radiation source so that the radiation based on the new imaging condition calculated by the imaging condition calculation unit during the radiation irradiation is Radiation source control means for feedback-controlling the radiation source so as to be irradiated.

  The radiography method of the present invention includes a plurality of pixels that are formed using a single crystal semiconductor, detect radiation images by receiving radiation irradiated to a subject according to imaging conditions, and represent the radiographic images from the pixels. Signal radiation accumulating step for accumulating signal charges in accordance with the amount of radiation incident on the pixels by radiological image detecting means capable of reading data nondestructively, and in the middle of the signal charge accumulating step, the radiographic image A non-destructive radiation image reading step, and a non-destructive pixel detection for detecting a pixel value of a non-destructive pixel that the radiation reaches directly without passing through the subject from the radiographic image obtained in the radiographic image reading step Step, and the pixel value of the unclear pixel and the radiation start time of the radiation are read nondestructively during the radiation irradiation. An expected arrival charge amount calculating step for calculating an expected arrival charge amount of the missing pixel at the end of irradiation of the radiation based on an elapsed time until acquisition of the radiation image and an irradiation end time of the radiation; The amount of expected arrival charge is compared with a predetermined threshold value, and when the amount of expected arrival charge exceeds the predetermined threshold value, the signal charge accumulated in the missing pixel at the radiation irradiation end time becomes the predetermined threshold value. The imaging condition calculating step for calculating the new imaging condition for controlling the radiation source and the radiation based on the new imaging condition calculated in the imaging condition calculating step are irradiated. A feedback control step of performing feedback control of the radiation in the middle of the charge accumulation step.

  According to the present invention, it is possible to prevent deterioration and failure of an FPD capable of nondestructive reading, and to extend the life of a radiographic apparatus using the FPD.

It is explanatory drawing which shows schematically the structure of a X-ray imaging system. It is explanatory drawing which shows the synchronization procedure of an X-ray source and an electronic cassette. It is a disassembled perspective view which shows the structure of an electronic cassette. It is sectional drawing which shows the structure of an electronic cassette. It is explanatory drawing which shows the structure of FPD. It is a block diagram which shows the structure of a console. It is a graph which shows typically frequency distribution of a pixel value. It is a flowchart which shows the effect | action of an X-ray imaging system. It is explanatory drawing which shows the X-ray irradiation of the comparative example which does not perform feedback control of a radiation source, and the change aspect of a blank pixel value. It is explanatory drawing which shows the X-ray irradiation of this invention, and the change aspect of an unclear pixel value. It is explanatory drawing which shows the example which stops irradiation of X-ray | X_line. It is explanatory drawing which shows the example which reads the X-ray image in multiple times, decreasing a gain gradually after completion | finish of X-ray irradiation.

  As shown in FIG. 1, the X-ray imaging system 10 includes an imaging table 11, an electronic cassette (radiation image detection device) 12, an X-ray source 17, a radiation source control device 19, a console 22, a monitor 23, and the like. The X-ray source 17 and the radiation source control device 19 constitute an X-ray generator, and the electronic cassette 12, the console 22, and the monitor 23 constitute an X-ray imaging device.

  An object H is placed on the imaging table 11, and an electronic cassette 12 is set inside the imaging table 11 or between the imaging table 11 and the object H so that X-rays transmitted through the object H are incident. The

  The electronic cassette 12 is detachably attached to the imaging table 11 and is portable even when the imaging table 11 is not used. The electronic cassette 12 controls the operations of the FPD 13 and the FPD 13 in a substantially rectangular housing. Various circuits, a memory 14, a communication unit 16 and the like are provided. The FPD 13 has pixels arranged in a matrix, accumulates signal charges corresponding to the incident X-ray dose in each pixel, converts the accumulated signal charges into voltage signals, and outputs them to detect an X-ray image. . The FPD 13 is a CMOS type capable of nondestructive reading. The memory 14 temporarily stores an X-ray image output from the FPD 13. The communication unit 16 communicates control signals with the console 22 and transmits an X-ray image stored in the memory 14 to the console 22. The electronic cassette 12 is a wireless type that incorporates a battery (not shown) for supplying power to each unit such as the FPD 13, and the communication unit 16 performs wireless communication with the console 22 by using light such as infrared rays or radio waves, for example. .

  The X-ray source 17 includes an X-ray tube that generates X-rays and a collimator that limits an X-ray irradiation field, and irradiates the subject H from the X-ray focal point 17a. The X-ray source 17 is provided so as to be movable with respect to the imaging table 11, and the X-ray irradiation position and angle can be freely adjusted. The X-ray source 17 is connected to a radiation source control device 19. The radiation source control device 19 is provided with a high voltage generator that supplies power to the X-ray source 17, and the tube voltage and tube current corresponding to the imaging conditions input to the radiation source control device 19 are X X-rays are emitted from the X-ray source 17 by being applied to the radiation source 17.

  X-ray quality (energy) is determined by tube voltage, and dose is determined by tube current and irradiation time. The radiation source controller 19 is provided with an operation panel 19a, through which imaging conditions such as tube voltage, tube current, and irradiation time of the X-ray tube are input. The imaging conditions are determined according to the imaging site, the physique of the subject, the age, and the like. In addition, an irradiation switch 21 for inputting an irradiation start signal is connected to the radiation source control device 19. The radiation source controller 19 is communicably connected to the console 22 of the X-ray imaging apparatus, and synchronizes the electronic cassette 12 and the X-ray source 17 by communicating with the console 22.

  As shown in FIG. 2, when the irradiation switch 21 is pressed and turned on, the radiation source control device 19 detects a synchronization start signal and starts operations such as filament heating of the X-ray tube and target rotation. As a result, the X-ray source 17 is shifted to an irradiation enabled state (hereinafter referred to as a ready state). When the X-ray source 17 becomes ready, the radiation source control device 19 transmits an accumulation start signal to the console 22. When the console 22 receives the accumulation start signal, the console 22 shifts the electronic cassette 12 from the standby operation to the accumulation operation for accumulating signal charges. When the electronic cassette 12 shifts to the accumulation operation, an irradiation start signal is transmitted to the radiation source controller 19. Upon receiving the irradiation start signal, the radiation source control device 19 supplies power corresponding to the tube voltage and tube current set in the imaging conditions to the X-ray source 17, so that the X-ray source 17 follows the imaging conditions. Irradiation of X-rays with a certain quality and dose toward the subject H is started. After starting X-ray irradiation, the radiation source control device 19 activates a timer to start measuring the irradiation time.

  The radiation source control device 19 monitors the timer, and when the irradiation time set in the imaging conditions has elapsed, sends an end synchronization signal to the X-ray source 17 to end the irradiation. In addition, the radiation source control device 19 transmits an end synchronization signal to the console 22. When the console 22 receives the end synchronization signal, the console 22 shifts the electronic cassette 12 from the accumulation operation to the read operation.

  As will be described later, the radiation source control device 19 receives input of new imaging conditions from the console 22 during X-ray irradiation, and changes the X-ray quality, dose, energy, irradiation time, etc. during irradiation. There are things to do.

  In FIG. 1, a console 22 is a control device that controls the operation of the electronic cassette 12 by transmitting a control signal via a communication unit 24. An input device (not shown) such as a keyboard and a mouse is connected to an irradiation switch 21. Yes. For example, an imaging condition similar to that of the radiation source control device 19 is input to the console 22 in advance using an input device.

  The console 22 is communicably connected to a medical information system such as a HIS (hospital information system) or RIS (radiation information system) (not shown), and can receive an X-ray imaging order from the medical information system. . An operator such as a doctor or a medical radiographer confirms subject information, a region to be imaged, and the like on the received imaging order, determines imaging conditions suitable for the information, and inputs them to the X-ray control unit 19 and the console 22 respectively. In this way, the shooting conditions are set. The photographing conditions input to the console 22 are set in the electronic cassette 12. The set shooting conditions are used for setting the initial value of the gain of the amplifier, as will be described later.

  X-rays irradiated from the X-ray source 17 and transmitted through the subject H are incident on the electronic cassette 12. The electronic cassette 12 detects an X-ray image representing the subject H by accumulating signal charges corresponding to the amount of incident X-rays in the accumulation operation. An X-ray image corresponding to the signal charge accumulated in the read operation is read. The read X-ray image is transmitted to the console 22 via the communication unit 24.

  The console 22 is provided with an image processing unit 26. The image processing unit 26 performs various types of image processing on the received X-ray image and outputs the processed image to the monitor 23. The image processing performed by the image processing unit 26 is, for example, defect correction processing or offset processing. The defect correction process is a process for correcting pixel values of defective pixels for which pixel values corresponding to the X-ray irradiation dose cannot be obtained by interpolation or the like. The offset processing is image processing for removing noise or the like due to dark current by subtracting an X-ray image (offset image) captured in advance without the subject H from an X-ray image obtained by capturing the subject H. . Note that the image processing unit 26 may perform image processing other than the correction processing, such as X-ray image composition processing, as will be described later.

  In addition to displaying an X-ray image, the monitor 23 displays an operation screen for operating the console 22.

  After the image processing is completed, the console 22 adds the input imaging conditions as supplementary information to the processed X-ray image, and converts it into a standard file format of a medical image file such as DICOM format. The converted image file is transmitted to a storage device such as an image server (not shown).

  As shown in FIGS. 3 and 4, the casing 31 of the electronic cassette 12 includes a front cover 31 a that covers the FPD 13 from the incident surface side on which X-rays enter and a back cover 31 b that covers the FPD 13 from the back. Both the front cover 31a and the back cover 31b are made of a metal material having a low X-ray transmittance (for example, stainless steel). However, the front cover 31a is provided with an X-ray transmission window 32 made of carbon or the like having a high X-ray transmittance so that X-rays are incident on the FPD 13. In the housing 31, in order from the front cover 31a side, the FPD 13, the support 33, and various circuit boards 34 to 37 for controlling the operation of the FPD 13 are arranged.

  The FPD 13 is an indirect conversion type composed of a scintillator 41 and an image sensor 42. The scintillator 41 is a phosphor that emits visible light having a light amount corresponding to the amount of incident X-rays, and is made of, for example, CsI: Tl (thallium activated cesium iodide), GOS (gadolinium oxysulfide), or the like. The image sensor 42 is a device that photoelectrically converts visible light emitted from the scintillator 41, and is a unit in which, for example, a plurality of CMOS sensor chips formed from a single crystal silicon substrate are bonded to increase the detection surface 43a. . The detection surface 43 a on which the pixels that perform photoelectric conversion are arranged is disposed so as to face the scintillator 41. The FPD 13 is disposed between the front cover 31a and the support 33 so that the image sensor 42 is on the front cover 31a side and the scintillator 41 is on the support 33 side (back cover 31b side). That is, the electronic cassette 12 is a backside illumination method (ISS), and the FPD 13 is arranged so that the detection surface 43a faces the X-ray incident side surface of the scintillator 41 that emits the largest amount of light. The scintillator 41 is bonded to the support 33 and the image sensor 42, and the image sensor 42 is bonded to the scintillator 41 and the front cover 31a.

  The support 33 is provided on the back side of the FPD 13 and bisects the space in the housing 31. Various circuit boards 34 to 37 are attached to the back side of the support 33. The support 33 is made of stainless steel, for example, and is fixed to the housing 31. Further, at the upper end and the lower end of the support 33, notches 33 a and 33 b are formed in the center portions, respectively, and a flexible cable 44 that connects the FPD 13 and the various circuit boards 34 to 37 is inserted. TCP (tape carrier package) type IC chips 46 and 47 are mounted on the flexible cable 44.

  The circuit board 34 is a driving circuit board on which a driving circuit 51 for driving the FPD 13 is formed. The circuit board 35 is an A / D conversion circuit board on which an output circuit (see FIG. 5) including the A / D conversion circuit 52 is formed. The A / D conversion circuit 52 converts an analog signal output from an IC chip 47 described later into a digital signal. The circuit board 36 is a board on which the memory 14 and the control unit 61 are formed. The control unit 61 controls each unit of the FPD 13 and controls communication with the console 22. The circuit board 37 is a power circuit board on which a power circuit is formed. The power supply circuit is a circuit that supplies power to each unit, and includes circuit elements such as an AC / DC converter that converts alternating current into direct current, and a DC / DC converter that converts direct current voltage into a voltage necessary for the operation of each circuit. .

  The IC chip 46 is an amplifier or a shift register that constitutes the drive circuit 51 together with circuit elements formed on the circuit board 41, and includes a vertical operation circuit 63 (see FIG. 5). The IC chip 47 is an ASIC that constitutes a read circuit including a CDS circuit 68 (see FIG. 5) described later.

  As shown in FIG. 5, the pixel PX of the FPD 13 includes a photodiode PD and circuit elements such as transistors M1 to M3, and performs an accumulation operation, a read operation, and a reset operation according to the driving state of the transistors M1 to M3. . The accumulation operation is an operation for accumulating signal charges generated by photoelectric conversion, and the read operation is an operation for outputting a voltage signal corresponding to the amount of accumulated signal charges. The reset operation is an operation for discarding the accumulated signal charge.

  The photodiode PD is an element that generates a signal charge corresponding to the amount of incident light from the scintillator 41 by photoelectric conversion, and is connected to the gate electrode of the amplification transistor M1 and the source electrode of the reset transistor M3. When accumulating signal charges and reading a voltage signal from the pixel PX, the reset transistor M3 is turned off, and the gate electrode of the amplifying transistor M1 corresponds to the amount of signal charge accumulated in the photodiode PD. A voltage is applied.

  A power supply voltage is applied to the source electrode of the amplifying transistor M1, and a pixel selecting transistor M2 is connected to the drain electrode, and a voltage corresponding to the amount of signal charge applied to the gate electrode is amplified at a predetermined amplification factor. Then, it is applied to the source electrode of the pixel selecting transistor M2. The gate electrode of the pixel selection transistor M2 is connected to the row selection line GL, the drain electrode is connected to the signal line SL, and when a gate signal is input from the row selection line GL, the voltage of the source electrode is changed to the signal line SL. Output to. Thereby, the voltage signal of the pixel PX is read out through the signal line SL. Note that since the FPD 13 is a CMOS type capable of nondestructive reading, each pixel PX holds a signal charge until resetting even if a voltage signal is read from the pixel PX.

  When discarding the signal charge accumulated in the photodiode PD, the resetting transistor M3 is turned on. The gate electrode of the reset transistor M3 is connected to the reset line RST, and a reset signal is input through the reset line RST. When the reset transistor M3 is turned on due to the input of the reset signal, the pixel PX in this row discards the signal charge accumulated in the photodiode PD to the drain electrode side of the reset transistor M3.

  The FPD 13 includes a control unit 61, a timing generator (TG) 62, a vertical scanning circuit 63, a horizontal scanning circuit 64, and the like.

  The control unit 61 comprehensively controls each unit of the FPD 13. The TG 62 generates a timing signal based on an instruction from the control unit 61. In addition, the control unit 61 sets the gain of the amplifier 72 of the output circuit 71 described later to a value instructed from the console 22 during the read operation. The vertical scanning circuit 63 and the horizontal scanning circuit 64 operate based on the clock signal input from the TG 62.

  The vertical scanning circuit 63 is a drive circuit for the pixel PX, selects a row of the pixel PX to be driven, and inputs a gate signal and a reset signal to the row selection line GL and the reset line RST of the selected row, thereby causing the pixel PX Causes the accumulation operation, the read operation, and the reset operation to be performed. The horizontal scanning circuit 64 is a circuit that selects the column of the pixels PX from which the voltage signal is read, and the column from which the reading is performed by turning on one of the column selection transistors 67 provided on each signal line SL. select.

  The row selection line GL is a wiring for inputting a gate signal for controlling the operation of the pixel PX from the vertical scanning circuit 63, and is provided for each row of the pixel PX. Based on the gate signal input through the row selection line GL, the pixel PX performs an accumulation operation or a read operation. The reset line RST is a wiring for inputting a reset signal to the pixels PX, and is provided for each row of the pixels PX. The pixel charge PX on the reset line RST to which the reset signal is input is reset after the signal charge is discarded.

  The signal line SL is a wiring for reading out a voltage signal (imaging signal) corresponding to the signal charge amount from each pixel PX, and is provided for each column of the pixels PX. A correlated double sampling (CDS) circuit 68 and a column selection transistor 67 are connected to the end of the signal line SL. The CDS circuit 68 operates based on the clock signal input from the TG 62, and samples the voltage signal so that noise due to reading is removed from the pixel PX on the row selection line GL selected by the vertical scanning circuit 63. And hold. The voltage signal held by the CDS circuit 68 is input to the output circuit 71 via the output bus line 69 when the column selection transistor 67 is turned on by the horizontal scanning circuit 64.

The output circuit 71 includes an amplifier 72 and an A / D conversion circuit 52. The amplifier 72 amplifies the voltage signal input from the CDS circuit 68 and inputs it to the A / D conversion circuit 52. The A / D conversion circuit 52 converts the voltage signal amplified by the amplifier 72 into digital data. Digital data output from the A / D conversion circuit 52 is temporarily stored in the memory 14 as an X-ray image and transmitted to the console 22 via the communication unit 16. The amplifier 72 is a variable gain amplifier whose gain can be freely adjusted, and is set to a predetermined value (initial value γ ₁ ) set based on imaging conditions in the first X-ray image reading in the reading operation. When the X-ray image is read again, the new value set by the control unit 61 is set. The data output from the output circuit 71 is stored in the memory 14 as described above, and is transmitted to the console 22 via the communication unit 16.

  As shown in FIG. 6, the console 22 includes a frame memory 81 along with the communication unit 24 and the image processing unit 26 described above. The frame memory 81 is a storage device capable of temporarily storing a plurality of X-ray images, and sequentially stores a plurality of X-ray images read by changing the gain of the amplifier 72 as will be described later. The

  In addition, the image processing unit 26 includes a missing pixel detection unit 82, an expected arrival charge amount calculation unit 83, an imaging condition calculation unit 84, an image synthesis unit 85, an offset processing unit 86, a defect correction processing unit 87, and the like.

  The element missing pixel detection unit 82 acquires an X-ray image (hereinafter referred to as “element missing detection image”) that has been read nondestructively during X-ray irradiation from the frame memory 81, and does not pass through the subject H directly. Detect the pixel where the line entered. Specifically, as shown in FIG. 7, the simple pixel detection unit 82 obtains a distribution of pixel values of all pixels. The distribution mode of the pixel values has a substantially constant distribution shape according to the imaging region (head, chest, abdomen, etc.) regardless of the gain of the amplifier 72. FIG. 7 shows, as an example, a frequency distribution of pixel values in the case of an adult chest X-ray image, and distribution A having two peaks represents an area of breast tissue or the like (lung, heart, rib, etc.). , Distribution B mainly represents pixels (so-called missing pixels) in which the X-rays directly reach the detection surface 43a without passing through the subject H.

  The simple pixel detection unit 82 detects the distribution B from the frequency distribution shape of the pixel values. Then, the pixel value D of the peak position value of the distribution B is obtained, and this is input as the representative value D of the pixel value of the missing pixel to the expected arrival charge amount calculation unit 83.

  The distribution shape of the pixel value itself is almost unchanged according to the imaging region. However, when the gain of the amplifier 72 is small, the frequency distribution of the pixel value as shown in FIG. When the gain of the amplifier 72 is large, the whole shifts to the right. Therefore, when X-ray images having different gains at the time of reading are compared, the specific value of the representative value D that is the peak value of the distribution B is different. .

  Further, as described later, from the viewpoint of monitoring the dose of the missing pixels, it is considered that the maximum pixel value of the entire missing detection image is set as the representative value D regardless of the frequency distribution of the pixel values. If the pixel having the maximum pixel value is simply selected as the determination pixel, the pixel value of the defective pixel that outputs a value that does not depend on the incident X-ray dose may be obtained. For this reason, the missing pixel detection unit 82 obtains the pixel value corresponding to the peak of the distribution B and sets it as the representative value D in order to surely obtain the pixel value of the missing pixel having no defect.

The expected arrival charge amount calculation unit 83 is based on the representative value D input from the missing pixel detection unit 82, and the expected amount of signal charge accumulated in the missing pixels at the end of X-ray irradiation (hereinafter, the expected arrival amount). Q) (referred to as charge amount) is calculated. For example, the irradiation start time of the X-ray Ts, the irradiation end time of X-ray Te, when the time read the background-detection image and _{t 1, Q = D (t} 1 -Ts) / (Te-Ts) Thus, the expected charge amount Q is estimated. The X-ray irradiation start time Ts is the input time of the start synchronization signal, and the X-ray irradiation end time Te is determined from the imaging conditions. The predicted arrival charge amount calculation unit 83 inputs the calculated expected arrival charge amount Q to the imaging condition calculation unit 84.

Here, during the X-ray irradiation period, it is assumed that the X-ray dose is constant and the expected charge amount Q increases in proportion to the time, but before and after the start of X-ray irradiation and the end of irradiation Before and after the hour, the amount of X-ray irradiation per unit time varies. Therefore, the expected arrival charge amount Q may be calculated more accurately in consideration of these. In any case, the expected arrival charge amount Q can be calculated from the representative value D, the read-out time t _{1 of the} image for detecting missing elements, the X-ray irradiation start time Ts, and the irradiation end time Te.

  When the expected arrival charge amount Q of the missing pixel is input from the expected arrival charge amount calculation unit 83, the imaging condition calculation unit 84 first compares the expected arrival charge amount Q with a predetermined threshold Th. The predetermined threshold Th is a value that determines an X-ray dose that may enter the image sensor 42 in one X-ray imaging, and is determined in advance by an experiment. If the X-ray dose received by the missing pixels (and all the pixels) in one X-ray imaging is less than or equal to the threshold Th, damage to the image sensor 42 due to the X-ray irradiation can be suppressed, so a certain number of imaging is guaranteed. Is done.

  When the expected arrival charge amount Q is equal to or less than the predetermined threshold Th (Q ≦ Th), the imaging condition calculation unit 84 sends a control signal indicating that X-ray irradiation under the current imaging conditions may be continued to the radiation source control device 19. To enter. In this case, the radiation source control device 19 continues the X-ray irradiation according to the imaging conditions input from the operation panel 19a before the imaging is started.

On the other hand, when the expected arrival charge amount Q exceeds the predetermined threshold Th (Q> Th), the imaging condition calculation unit 84 reads the threshold value Th, the expected arrival charge amount Q, the read-out time t ₁ of the element for detecting missing images, Based on the irradiation end time Te, a new imaging condition is calculated so that the signal charge amount accumulated in the unclear pixel at the X-ray irradiation end Te becomes a threshold Th (or a predetermined value equal to or less than the threshold Th). Input to the X-ray source controller 19. The new imaging condition calculated by the imaging condition calculation unit 84 is, for example, the tube voltage of the X-ray tube. When a new imaging condition input from the imaging condition calculation unit 84 is input, the X-ray control unit 19 immediately irradiates the X-ray according to the new imaging condition from the X-ray source 17.

  Here, an example is given in which the imaging condition calculation unit 84 calculates the tube voltage as a new imaging condition when the expected arrival charge amount Q exceeds the predetermined threshold Th, but the imaging condition calculation unit 84 calculates. The new imaging condition may be the tube current of the X-ray tube. Further, an imaging condition for changing both the tube current and the tube voltage of the X-ray tube may be calculated. By reducing the tube voltage, the energy of the irradiated X-rays is reduced. Further, when the tube current is lowered, the X-ray dose reaching the image sensor 42 is reduced. By reducing the dose and energy of the X-rays to be irradiated, the signal charge amount accumulated in the unclear pixels at the end Te of the X-ray irradiation is suppressed to the threshold Th.

  When the X-ray tube current of the X-ray tube is changed to reduce the dose of X-rays to be irradiated and the X-ray tube voltage is lowered to reduce the energy of the X-rays to be irradiated, It remains the same that the amount of signal charge accumulated in the unexposed pixels at the end of irradiation is reduced to the threshold value Th and damage to the image sensor 42 is reduced. However, the lower the energy at the same dose, the less damage the image sensor 42 has. For this reason, it is particularly preferable that the imaging condition calculation unit 84 reduces at least the energy of X-rays when calculating a new imaging condition. For this reason, in this example, the energy of X-rays to be irradiated is reduced by lowering the tube voltage.

  As described above, the missing pixel detection unit 82, the expected arrival charge amount calculation unit 83, and the imaging condition calculation unit 84 function during X-ray imaging. On the other hand, the image composition unit 85, the offset processing unit 86, and the defect correction processing unit 87 function after reading out the X-ray image.

  Further, as described above, the X-ray imaging apparatus 10 controls the X-ray dose or the like based on the non-destructive readout image that is read nondestructively during the X-ray irradiation, after the X-ray irradiation ends. The X-ray imaging apparatus 10 reads out X-ray images a plurality of times while changing the gain of the amplifier 72. The number of readings after the end of X-ray irradiation is determined in advance by the imaging region or the like.

  The image composition unit 85 synthesizes the plurality of X-ray images to generate a composite X-ray image, and the offset processing unit 86 performs an offset process on the composite X-ray image. In addition, the defect correction processing unit 84 performs defect correction processing on the combined X-ray image after the offset processing. The composite X-ray image after the defect correction process is displayed on the monitor 23 or transmitted to an image server (not shown).

  The image composition unit 85 acquires a plurality of X-ray images obtained after the end of X-ray irradiation from the frame memory 81 and combines them to generate a composite X-ray image. When switching to X-ray irradiation based on a new imaging condition calculated so that the pixel value of the missing pixel is equal to or less than the threshold value Th during X-ray irradiation based on the imaging condition input first, the pixel PX is accumulated. The amount of signal charge to be reduced is smaller than that in the case of continuing X-ray irradiation based on the initial imaging conditions. For this reason, when the X-ray image is detected by changing the imaging conditions during the X-ray irradiation, when the X-ray image is read from the FPD 13, the noise component of the read circuit 71 and the like becomes relatively large, and the read X-ray The image as it is may not be suitable for observation of the subject H. For this reason, the X-ray imaging apparatus 10 reads out X-ray images a plurality of times while changing the gain of the amplifier 72, and combines them with the image synthesis unit 85, thereby reducing noise caused by the readout circuit 71 and the like. The dynamic range is expanded, and a composite X-ray image more suitable for observation of the subject H is generated.

For example, the image composition unit 85 generates a composite X-ray image by averaging a plurality of X-ray images. Here, the image composition unit 85 generates a composite X-ray image by simple averaging, but the composite X-ray image is obtained by weighting and averaging in accordance with the gain value at the time of reading each X-ray image. May be generated. Further, when a high-sensitivity image obtained by reading out an X-ray image with two types of gains and low and a low-sensitivity image are combined, the high-sensitivity image X _h , the low-sensitivity image X _l , weights _{w h} image _{X h,} weights _{w l} of the low-sensitivity image, a threshold value t which shows the synthesis starting level, using a sensitivity ratio S of the high-sensitivity image _{X h} and the low-sensitivity image _{X l,} a synthetic X-ray image _{X c , X c = {w h ·} X h + w l · (X l -t / S + t)} / (w h + w l) may be calculated according to. In this case, w _h + w _l = 1 may be set.

  The operation of the X-ray imaging system 10 configured as described above will be described below with reference to the flowchart shown in FIG.

  As shown in FIG. 8, when X-ray imaging is performed by the X-ray imaging system 10, first, imaging conditions are input to the radiation source controller 19 and the console 22 (step S11). The X-ray tube voltage, tube current, and irradiation time are input to the radiation source controller 19 as imaging conditions. The imaging conditions input to the radiation source control device 19 are determined based on the subject information such as the imaging part, the physique of the subject H, the age, and the like confirmed on the imaging order received by the console 22.

  The imaging conditions input to the console 22 are transmitted to the electronic cassette 12 via the communication unit 24. When the imaging condition is input from the console 22, the electronic cassette 12 starts a standby operation (step S12). The standby operation is an operation of waiting for the start of X-ray irradiation by repeatedly performing a reset operation at a predetermined timing.

  When the irradiation switch 21 is pressed, a synchronization start signal is transmitted to the X-ray source 17 via the radiation source control device 19, and the X-ray source 17 is shifted to the ready state. When the X-ray source 17 becomes ready, the radiation source control device 19 transmits an accumulation start signal to the console 22 and the electronic cassette 12 starts an accumulation operation (step S13).

  When the electronic cassette 12 shifts to the accumulation operation, the radiation source control unit 19 receives an irradiation start signal from the console 22 and starts X-ray irradiation from the X-ray source 17 with a tube voltage and a tube current according to the imaging conditions. (Step S14).

  When the electronic cassette 12 starts the accumulation operation, the electronic cassette 12 performs nondestructive reading at a predetermined timing during the X-ray irradiation (step S15). The X-ray image obtained here is transmitted to the console 22 as an image for detecting missing elements and stored in the frame memory 81.

  The background missing detection image is read from the frame memory 81 by the background missing pixel detection unit 82, and the background missing pixel is detected based on the background missing detection image (step S16). More specifically, the pixel value D at the peak position of the distribution B (see FIG. 7) to which the missing pixels belong is analyzed as a representative value D of the missing pixels by analyzing the histogram of the missing images. Detected.

  Then, the expected arrival charge amount calculation unit 83 calculates the expected arrival charge amount Q of the missing pixel based on the representative value D (step S17). As described above, the expected arrival charge amount Q is a signal charge amount accumulated in the missing pixels at the end of the X-ray irradiation when the X-ray irradiation is continued according to the imaging condition input first.

  Thereafter, in the imaging condition calculation unit 84, the expected arrival charge amount Q is compared with a threshold Th (step S18). When the expected arrival charge amount Q exceeds the threshold value Th, the imaging condition calculation unit 84 sets the signal charge amount accumulated in the missing pixels to the threshold value Th (or a predetermined value equal to or less than the threshold value Th) at the end of X-ray irradiation. Then, a new photographing condition is calculated (step S19). The imaging conditions calculated here are obtained by reducing the tube voltage with respect to the initial imaging conditions.

  When the new imaging condition calculated by the imaging condition calculation unit 84 is calculated in this way, the X-rays to be irradiated are feedback-controlled based on the new imaging condition (step S20). Specifically, when receiving a new imaging condition input from the imaging condition calculation unit 84, the radiation source control device 19 immediately irradiates the X-ray according to the imaging condition from the X-ray source 17. As a result, X-rays with reduced energy are emitted from the middle of X-ray irradiation compared to X-rays based on the imaging condition input first.

  On the other hand, when the expected charge amount Q is equal to or less than the threshold Th (S18: No), the calculation of new imaging conditions (S19) and the feedback control (S20) for reducing the X-ray energy are not performed, and the X-ray controller 19 Is based on the input of the control signal from the imaging condition calculation unit 84 and continues the X-ray irradiation according to the imaging condition input first.

When the irradiation time set in the imaging conditions has elapsed after monitoring the timer, the radiation source control device 19 sends an end synchronization signal to the X-ray source 17 to end the X-ray irradiation (step S21). At the same time, the radiation source control device 19 transmits an end synchronization signal to the console 22, and the console 22 shifts the electronic cassette 12 from the accumulation operation to the read operation (step S22). In the read operation performed after the X-ray irradiation is completed, X-ray images are read a predetermined number of times while changing the gain of the amplifier 72. For example, in the first (first) reading in the reading operation, the gain of the amplifier 72 of the output circuit 71 is set to a predetermined value (initial value) γ ₁ according to the shooting conditions, and in that state all the pixels are set. Read nondestructively. In the next (second) reading, the gain of the amplifier 72 is set to a value γ ₂ (γ ₁ ≠ γ ₂ ) different from the initial value γ ₁ . Thereafter, the same is true, and the value of the gain is different for each predetermined number of readings.

  The read X-ray images are transmitted to the console 22 and stored in the frame memory 81. Then, the X-ray image is synthesized (averaged) by the image synthesis unit 85 to generate a synthesized X-ray image (step S23). The generated composite X-ray image is subjected to offset processing (step S24) and defect correction processing (step S25), and is displayed on the monitor 23 or transmitted to an image server (not shown). At the same time, the electronic cassette 12 performs a reset operation and discards the signal charge accumulated in the pixel PX (step S26).

  As described above, the X-ray imaging system 10 feedback-controls the X-rays to be irradiated based on the pixel values (representative values D) of the non-destructive pixels read nondestructively during the X-ray irradiation, and the X war victim It operates so that the pixel value of the missing pixel at the end does not exceed the threshold Th.

  As shown in FIG. 9, when the X-ray source 17 is not subjected to feedback control and X-ray irradiation is performed in accordance with the first inputted imaging conditions, the irradiation control 21 is turned on to start irradiation to the source control device 19. X-ray irradiation starts at time Ts when the signal is input, and X-ray irradiation ends at time Te, which is the time when the irradiation time determined by the imaging conditions has elapsed. The electronic cassette 12 performs a standby operation before starting X-ray irradiation (to Ts), performs an accumulation operation during X-ray irradiation (Ts to Te), and performs a reading operation after X-ray irradiation ends (Te to). .

  The pixel value of the missing pixels during this period (hereinafter referred to as the missing pixel value) changes as follows. Before starting the X-ray irradiation, a slight signal charge may be generated due to noise caused by dark current. However, since the electronic cassette 12 is in a standby operation and repeatedly performs a reset operation at a predetermined interval, the pixel without pixels The value is almost zero. When X-ray irradiation is started, the blank pixel value increases almost in proportion to the X-ray irradiation time. Then, at the end of X-ray irradiation, for example, it is assumed that the missing pixel value exceeds the threshold Th.

  Since the image sensor 42 uses a single crystal silicon substrate, the larger the accumulated value of the X-ray irradiation amount, the more the deterioration or failure occurs. Therefore, if the unclear pixel value exceeds the threshold Th in one X-ray imaging. Deterioration or failure occurs, and the number of shootings that can be performed before it can no longer be used is reduced. Naturally, as the amount exceeding the threshold Th increases, the probability of occurrence of the image sensor 42 and the failure increases, and the electronic cassette 12 becomes short-lived.

On the other hand, as shown in FIG. 10, in the X-ray imaging system 10, after starting X-ray irradiation (Ts˜), an X-ray image is read nondestructively as an element for detecting missing elements at a predetermined timing t _1. The missing pixel value D is obtained. Then, the X-ray source 17 is feedback-controlled based on the unclear pixel value D to reduce the X-ray irradiation energy. The irradiation energy of X-rays after t ₁ is determined so that the missing pixel value of Te at the end of X-ray irradiation becomes the threshold value Th. The pixel value does not exceed the threshold value Th.

  For this reason, in the X-ray imaging system 10, as described above, the image sensor 42 is deteriorated as compared with the comparative example in which the feedback control of the X-ray to be irradiated is not performed on the basis of the missing pixel value during the X-ray irradiation. And the occurrence of failure can be suppressed, and the life of the electronic cassette 12 can be extended. Specifically, assuming that the minimum number of times of imaging by the electronic cassette 12 is determined, in the case of the above-described comparative example, the minimum number of times of imaging is obtained when X-ray imaging is performed such that the blank pixel value exceeds the threshold Th. Although the image sensor 42 is deteriorated / failed before imaging, the X-ray imaging system 10 can reliably perform imaging for the minimum number of imaging times.

  Compared with the case where the so-called region-of-interest pixel is monitored and the X-ray feedback control is performed, the region of interest is usually the subject portion, and therefore the X-ray is fed back so that the pixel value of the region of interest is not saturated. However, the pixel value exceeds the threshold value Th in the simple pixel without the subject H. The position of the missing pixel changes according to the subject H, and if the subject H changes, the pixel that was a missing pixel in the previous imaging may become a pixel of interest in the next imaging. For this reason, even if the pixels of the region of interest are monitored, the lifetime of the electronic cassette 12 cannot be extended. On the other hand, since the X-ray imaging apparatus 10 monitors the missing pixels, even if the subject H changes and the missing pixels become the pixels of the region of interest, the life of the electronic cassette 12 is not affected.

  In the above-described embodiment, the example in which the tube voltage of the X-ray tube is lowered to reduce the energy of the X-ray when performing feedback control of the X-ray to be irradiated has been described. He also mentioned an example of reducing the tube current and reducing the X-ray dose. However, as shown in FIG. 11, the X-ray irradiation time is shortened from Te to Te ′, and the X-ray irradiation is stopped at the time when the blank pixel value at the end of the X-ray irradiation becomes the threshold Th. , It may be possible to prevent the missing pixel value from exceeding the threshold Th. In the case of shortening only the X-ray irradiation time, the X-ray dose and energy are based on the imaging conditions input first. As described above, in order to prevent the deterioration and failure of the single crystal silicon substrate due to the incident amount of X-rays, it is most effective to reduce the energy of X-rays. At the same time, it is more preferable to lower the tube voltage of the X-ray tube to lower the X-ray energy.

  In the above-described embodiment, an example in which the X-ray energy is reduced has been described as an example of X-ray feedback control. However, the X-ray source 17 is controlled so as to narrow the X-ray irradiation range. The signal charge amount of the missing pixels may not be increased. In this way, it is possible to prevent the image sensor 42 from being damaged while continuing the X-ray irradiation according to the imaging condition input first.

  In the above-described embodiment, the example in which the X-ray energy is feedback controlled has been described. However, the collimator of the X-ray source 17 may be controlled to limit the irradiation field. For example, pixels belonging to the distribution A representing the subject H are detected by histogram analysis in the same manner as calculating the representative value D from the image for detecting missing images. Then, the collimator is controlled so that a predetermined range (for example, a rectangle) including all the pixels belonging to the distribution A becomes a new irradiation field, and the irradiation field is limited to the portion of the subject H. In this way, since the missing pixels whose predicted arrival dose Q is likely to exceed the threshold Th are not irradiated with X-rays, the image sensor 42 is prevented from being damaged by excessive X-ray irradiation exceeding the threshold Th. In this case, since it is only necessary to reduce the dose of X-rays to the unexposed pixels, it is not necessary that the entire subject H is included in the new irradiation field, and a part of the subject H (region of interest) ) May be limited to the irradiation field.

  In the above-described embodiment, an example in which non-destructive readout is performed only once during X-ray irradiation to obtain an image for detecting missing elements is described. However, non-destructive readout is performed a plurality of times during X-ray irradiation. May be. When multiple non-destructive readouts are performed during X-ray irradiation to obtain a plurality of undetected images, a representative value D is calculated for each of these, and an expected charge reached based on a change in the representative value D The quantity Q is calculated. In this way, it is possible to calculate the expected arrival charge amount Q with higher accuracy than when performing non-destructive reading only once as in the above-described embodiment.

  In the above-described embodiment, even when an image for detecting a missing element is obtained during X-ray irradiation, reading is performed in the same manner as after the end of X-ray irradiation. Since the amount of irradiation of the line is still small and the pixel value is small, it is preferable to read out with increased sensitivity by binning a plurality of pixels to reduce noise due to readout. In this way, when the pixels are binned and read out, the reading speed is also improved, so the time until the X-ray is feedback-controlled is shortened, and the control accuracy is improved. If only X-ray feedback control is to be performed at a high speed, so-called thinning readout may be performed in which data of some pixels is not read out when obtaining an image for detecting missing elements. However, the reading by binning or thinning has a sufficient accuracy since the resolution is lowered, but the resolution is not so necessary for feedback control.

  In addition, as can be seen from the fact that binning and thinning-out reading may be performed in this way, in the above-described embodiments, the element missing pixel detection unit 82 is represented by histogram analysis using all pixels for element missing detection. Although the value D is determined, it is not necessary to use all pixels. Specifically, it is sufficient to obtain pixel values uniformly from the entire region of the X-ray image so that the above-described histogram analysis can be performed. The degree to which the histogram analysis can be performed is such that the resolution is not extremely different from the X-ray image finally used for observation.

  In the above-described embodiment, the X-ray image is read out a predetermined number of times while changing the gain of the amplifier 72 after the X-ray irradiation is finished. It may be determined based on the image for use. For example, the imaging condition calculation unit 84 determines the number of readouts after the end of X-ray irradiation so as to be proportional to the expected arrival charge amount Q. When the expected arrival charge amount Q is large, if the unclear pixel value is suppressed to the threshold value Th by X-ray feedback control, the difference between the expected arrival charge amount Q and the actual signal charge amount is large, and the signal charge amount actually obtained And the pixel value of an X-ray image becomes small. For this reason, in order to obtain an X-ray image having the same image quality as that obtained when imaging is performed under the first input imaging conditions, the larger the expected charge amount Q, the greater the number of readouts after the end of X-ray irradiation. This is because it is necessary to reduce noise in a combined X-ray image generated from a plurality of X-ray images.

  Further, the number of readings may be determined based on an X-ray image obtained first after the end of X-ray irradiation. For example, when the pixel value of the first X-ray image obtained after the end of X-ray irradiation is small and the amount of signal charge accumulated in the pixel PX is small, the number of readings is increased. In this way, as described above, the noise of the composite X-ray image is reduced, and the deterioration of the image quality can be minimized as compared with the case where the X-ray is not feedback-controlled.

  In the above-described embodiment, the example in which the gain of the amplifier 72 is changed every time the X-ray image is read has been described. However, the X-ray image to be read is divided into a plurality of areas, and the gain of the amplifier 72 is changed for each area. An X-ray image may be read out.

  In the above-described embodiment, an example in which X-ray images are read out a plurality of times after the end of X-ray irradiation has been described. For example, a region of interest is defined as an imaging condition, pixels in the region of interest are not saturated, and The gain of the amplifier 72 may be determined so as to obtain the lowest noise, and the X-ray image may be read out only once. The gain of the amplifier 72 may be determined based on, for example, the pixel value of the pixel of the region of interest in the background missing detection image so that the pixel of the region of interest does not saturate and has the lowest noise. In this case, destructive readout may be performed.

  In the above-described embodiment, the example in which the X-ray image is read a plurality of times while the signal charge is held in the pixel PX after the end of the X-ray irradiation has been described. However, the CDS circuit 68 samples the voltage signal, In the held state, a plurality of data may be acquired for the same pixel PX while changing the gain of the amplifier 72. This corresponds to acquiring each X-ray image obtained by performing reading a plurality of times for each row. Thus, the signal charge of the pixel PX can be discarded simultaneously with the reading.

In the above-described embodiment, the gain value of the amplifier 72 is arbitrary, but the gain of the amplifier 72 is preferably 2 ⁿ (n is an integer other than 0) times the initial value γ ₁ , for example. This is because the calculation of the pixel value for generating the composite X-ray image can be performed by bit shift, and it is easy and fast. Although explanation is omitted in the above embodiment, it is preferable to add header information indicating a gain value at the time of reading to the X-ray image.

In the above-described embodiment, the value of the gain of the amplifier 72 is arbitrary, but it is preferable that the gain of the amplifier 72 is gradually decreased every time reading is performed a plurality of times. For example, as shown in FIG. 12, when X-ray images are read five times in total, the gain γ ₂ set at the second reading is smaller than the gain γ ₁ at the first (first) reading. Set a value within the range. The gain γ ₃ set at the time of the third reading is set to a value within a range smaller than the gain γ ₂ at the time of the second reading. The same applies to the fourth and subsequent readings, and γ ₁ > γ ₂ > γ ₃ > γ ₄ > γ ₅ >. This is because, even in the absence of X-ray irradiation, the noise component due to dark current increases with the passage of time. Therefore, if high sensitivity (high gain) reading is performed later, the noise component will increase. Because. As described above, when the gain is gradually lowered at each re-reading, the gain γ ₁ at the first (first) reading is a value determined in advance according to the photographing condition as described above, but the gain γ at the first reading is set. ₁ is intentionally set higher than the appropriate value. In this way, the above conditions can be easily satisfied.

  In the above-described embodiment, the example in which the offset process and the defect correction process are performed after generating a composite X-ray image from a plurality of X-ray images having different gains at the time of reading has been described. A composite X-ray image may be generated after performing an offset process and a defect correction process on the X-ray image. A composite X-ray image may be generated after the offset process, and then a defect correction process may be performed, or a composite X-ray image may be generated after the defect correction process, and then the offset process may be performed. The order in which the offset process and the defect correction process are performed is also arbitrary.

  In the above-described embodiment, an example in which re-reading is performed while changing the gain of the amplifier 72 of the output circuit 71 has been described. However, the gain of the transistor M1 of each pixel PX is variable and a voltage signal is output from each pixel PX. The gain may be adjusted at the stage of performing. Since the pixel PX of the FPD 13 has a larger pixel size than a small solid-state image sensor used in a digital camera or the like, a variable gain amplifier can be used for each pixel PX of the FPD 13.

  In the above-described embodiment, an example in which an entire image of the detection surface 43a is obtained when obtaining an image for detecting a missing image is described. However, the image for detecting a missing image is read from a part of the pixels on the detecting surface 43a. It may be an image. For example, when the imaging conditions are determined, a region where there is no subject H and is likely to be a missing pixel (hereinafter referred to as a non-interest region) is determined. Since the background missing detection image is for detecting the pixel value of the background missing pixel, the representative value D and the like are determined to obtain only the image of the non-interesting region, and Thus, the X-ray can be feedback controlled. Thus, when an image of a non-interest area is obtained as an image for detecting missing elements, the number of pixels to be acquired is smaller than when the entire image of the detection surface 43a is read out, so X-ray feedback control is performed more quickly. It can be carried out. However, when an image of a non-interest region is obtained, the histogram analysis cannot be performed as described above. For example, the average value of the pixel values of the image of the non-interest region may be set as the representative value D.

  In the above-described embodiment, the threshold value Th is a predetermined value that is determined in advance regardless of the shooting conditions, but the threshold value Th is preferably determined according to the shooting conditions. That is, it is preferable that the threshold value Th is a different value for each photographing condition. For example, since the damage to the image sensor 42 increases as the energy and dose of X-rays increase, the threshold Th is preferably set low when the irradiated X-rays are high energy (high dose), and low energy (low dose) This is because the threshold Th can be set high to allow more signal charges to be accumulated.

  In the above-described embodiment, the example in which the electronic cassette 12 and the X-ray source 17 are synchronized by the radiation source control device 19 communicating with the console 22 has been described. However, the radiation source control unit 19 and the console 22 communicate with each other. In addition, an X-ray imaging system that detects the start and end of X-ray irradiation, the incident X-ray dose, and the like with an electronic cassette and performs automatic exposure control (AEC) is also known. In such an X-ray imaging system, for example, in addition to the pixel PX for detecting an X-ray image, a phototimer for measuring an incident X-ray dose or the like is provided in front of the image sensor 42 (surface on the X-ray incident side). May be provided. However, since the X-ray imaging system 10 calculates the amount of incident X-rays using the pixel PX for X-ray image detection, a phototimer is unnecessary, and there is no loss of incident X-ray energy or the like due to the phototimer. . For this reason, in the X-ray imaging system 10, it is easy to reduce the exposure dose of the subject H by reducing the energy of X-rays.

  In the above-described embodiment, the example in which the X-ray feedback control is performed based on the predicted arrival charge amount Q has been described. However, the amount calculated for performing the X-ray feedback control is not limited to the predicted arrival charge amount Q. For example, based on the representative value D, an arrival dose (expected arrival dose) to the unclear pixel at the end of X-ray irradiation may be calculated. Also in this case, the control mode and the like are the same as those in the above-described embodiment.

  In the above-described embodiment, the missing pixels and the pixel values (representative values D) of the missing pixels are detected by performing histogram analysis on the missing images. However, the image shown in the missing images is displayed. May be detected by image analysis that divides the subject H and the subject H. In this case, for example, the average value of the detected element missing pixels may be set as the representative value D in the above-described embodiment.

  In the above-described embodiment, the example in which the image sensor 42 is formed using a single crystal silicon substrate has been described. However, the present invention also applies to the case where a crystal semiconductor other than silicon, a compound semiconductor, or an organic semiconductor is used for the image sensor 42. Is preferred.

  As mentioned above, although embodiment of this invention was described, this invention is not limited to said each embodiment, A various change is possible in the range which does not deviate from the meaning of this invention.

  For example, in each of the above embodiments, the back-illuminated electronic cassette 12 has been described as an example. However, a front-incident (PSS) type electronic cassette in which the scintillator and the FPD are arranged in this order from the X-ray incident side is also preferably used. it can. In the above-described embodiment, an example in which the FPD 13 is a CMOS type has been described. However, an image sensor of another aspect may be used as long as non-destructive reading can be performed. Furthermore, although the indirect conversion type FPD 13 using the scintillator 41 has been described as an example, a direct conversion type FPD that converts X-rays directly into charges by an X-ray conversion layer such as amorphous selenium without using a scintillator can also be used. .

  In addition, synchronization control related to the start and end of X-ray irradiation between an X-ray generation device including an X-ray source and a radiation source control device and an X-ray imaging device including an electronic cassette and a console is performed by communication of synchronization signals. As described in the example, the X-ray imaging apparatus may be provided with a function of self-detecting the start and end of X-ray irradiation by providing an irradiation detection sensor or the like. This eliminates the need for synchronization signal communication between the X-ray generator and the X-ray imaging apparatus. In this case, an FPD pixel may be used as an irradiation detection sensor.

  Moreover, although the X-ray imaging apparatus was shown in the example which comprises an electronic cassette and a console, the function for controlling the electronic cassette among the console functions is configured as an imaging control apparatus separate from the console, and the electronic cassette and imaging control are performed. You may comprise an X-ray imaging apparatus by three apparatuses, an apparatus and a console. Further, the function of the imaging control device may be built in the electronic cassette, or instead of or in addition, the function of the image processing unit 26 may be built in the electronic cassette. Further, in addition to the imaging control device, the electronic cassette and the imaging control device or console may be integrated, for example, the function of the console is built in the electronic cassette.

  The X-ray imaging apparatus of the present invention is not limited to an electronic cassette, and may be a stationary X-ray imaging apparatus in which an FPD is built in an imaging table.

  Further, the circuit element 41 and the drive circuit 51 formed on the circuit board 41 are used as the IC chip 46, and the ASIC that constitutes the readout circuit including the CDS circuit 68 and the like is used as the IC chip 47. Although examples have been described, these may be formed on the substrate of the image sensor 42. That is, the various circuits and the like may be formed on the single crystal silicon substrate integrally with the pixel PX and the like. In this way, if various circuits such as the drive circuit 51 and the CDS circuit 68 are provided on the same substrate integrally with the pixel PX or the like, IC chips and flexible cables can be reduced, and assembly is facilitated, thereby reducing costs. Can be achieved.

  The present invention can be applied not only to X-rays but also to imaging systems that use other radiation such as gamma rays.

10 X-ray imaging system 12 Electronic cassette 13 FPD
17 X-ray source 22 Console 26 Image processor 31 Case 41 Scintillator 42 Image sensor 52 A / D conversion circuit 71 Output circuit 72 Amplifier

Claims (12)

A radiation image detecting means for detecting a radiation image of a subject by receiving radiation irradiated from a radiation source according to imaging conditions and passing through the subject, and formed using a single crystal semiconductor, and depending on an incident amount of the radiation A plurality of pixels for accumulating signal charges, and radiation image detection means capable of nondestructively reading out data representing the radiation image based on the amount of accumulation of signal charges from the pixels;
From the radiation image read nondestructively during the irradiation of the radiation, a missing pixel detection means for detecting a pixel value of a missing pixel that the radiation reaches directly without passing through the subject;
Based on the pixel value of the unclear pixel, the elapsed time from the irradiation start time of the radiation to the acquisition time of the radiation image read nondestructively during the irradiation of the radiation, and the irradiation end time of the radiation , An expected arrival charge amount calculation means for calculating the signal charge amount accumulated in the unexposed pixels at the end of irradiation of the radiation as an expected arrival charge amount;
The amount of expected arrival charge is compared with a predetermined threshold value, and when the amount of expected arrival charge exceeds the predetermined threshold value, the signal charge accumulated in the missing pixel at the radiation irradiation end time becomes the predetermined threshold value. Imaging condition calculation means for calculating a new imaging condition for controlling the radiation source;
A radiation source that feedback-controls the radiation source so that the radiation based on the new imaging condition calculated by the imaging condition calculation unit is irradiated during the irradiation of the radiation based on the first imaging condition. Control means;
A radiation imaging apparatus comprising:   The radiation imaging apparatus according to claim 1, wherein the new imaging condition calculated by the imaging condition calculation unit is an imaging condition for reducing energy of the radiation.   The radiation imaging apparatus according to claim 1 or 2, wherein the new imaging condition calculated by the imaging condition calculation means is an imaging condition for reducing the radiation dose.   The radiation imaging apparatus according to claim 1, wherein the new imaging condition calculated by the imaging condition calculation unit is an imaging condition for shortening the irradiation time of the radiation.   The radiographic apparatus according to claim 1, wherein the new imaging condition calculated by the imaging condition calculation unit is an imaging condition that limits an irradiation field of the radiation. Amplifying a signal corresponding to the signal charge when reading the radiation image from the radiation detection means, and an amplification means having a variable gain;
An image synthesizing unit that synthesizes a plurality of the radiographic images read while changing the gain of the amplification unit after the irradiation of the radiation, and generates a synthetic radiographic image for observing the subject;
The radiation imaging apparatus according to claim 1, comprising:   7. The image synthesizing unit synthesizes the plurality of radiographic images read while changing the gain of the amplifying unit so that noise at the time of reading the radiographic image is reduced. Radiography equipment.   8. The radiographic apparatus according to claim 6, wherein the image synthesizing unit synthesizes the plurality of radiographic images read while changing the gain of the amplifying unit so that a dynamic range is expanded. .   The said image synthetic | combination means produces | generates the said synthetic | combination radiographic image by averaging the several said radiographic image read, changing the gain of the said amplification means. The radiation imaging apparatus described in 1.   The radiation detection means includes a scintillator that converts the radiation into light, and an image sensor that detects the radiation image by photoelectrically converting light emitted from the scintillator, and the image sensor from the radiation incident side. The radiation imaging apparatus according to claim 1, wherein the scintillators are arranged in the order of the scintillators. In a radiation imaging system comprising a radiation source and a radiation imaging apparatus that receives radiation transmitted from the radiation source according to imaging conditions and transmitted through the subject, and captures a radiation image of the subject,
The radiation imaging apparatus includes:
A plurality of pixels that are formed using a single crystal semiconductor and accumulate signal charges according to the amount of incident radiation, and non-destructive data representing the radiation image based on the amount of signal charges accumulated from the pixels. A radiation image detecting means capable of reading out
From the radiation image read nondestructively during the irradiation of the radiation, a missing pixel detection means for detecting a pixel value of a missing pixel that the radiation reaches directly without passing through the subject;
Based on the pixel value of the unclear pixel, the elapsed time from the irradiation start time of the radiation to the acquisition time of the radiation image read nondestructively during the irradiation of the radiation, and the irradiation end time of the radiation , An expected arrival charge amount calculating means for calculating an expected arrival charge amount of the missing pixel at the end of irradiation of the radiation;
The amount of expected arrival charge is compared with a predetermined threshold value, and when the amount of expected arrival charge exceeds the predetermined threshold value, the signal charge accumulated in the missing pixel at the radiation irradiation end time becomes the predetermined threshold value. An imaging condition calculation means for calculating a new imaging condition for controlling the radiation source;
A radiation source control unit that feedback-controls the radiation source so that the radiation based on the new imaging condition calculated by the imaging condition calculation unit is irradiated during the irradiation of the radiation,
A radiation imaging system comprising: It is formed using a single crystal semiconductor, has a plurality of pixels that receive radiation irradiated to a subject according to imaging conditions and detects a radiation image, and can read data representing the radiation image from the pixels in a non-destructive manner. A signal charge accumulating step for accumulating signal charges corresponding to the amount of radiation incident on the pixels by a possible radiation image detection means;
In the middle of the signal charge accumulation step, a radiation image reading step for reading the radiation image nondestructively,
From the radiation image obtained in the radiation image reading step, a missing pixel detection step for detecting a pixel value of a missing pixel that the radiation reaches directly without passing through the subject;
Based on the pixel value of the unclear pixel, the elapsed time from the irradiation start time of the radiation to the acquisition time of the radiation image read nondestructively during the irradiation of the radiation, and the irradiation end time of the radiation An expected arrival charge amount calculating step for calculating an expected arrival charge amount of the unclear pixel at the end of irradiation with the radiation;
The amount of expected arrival charge is compared with a predetermined threshold value, and when the amount of expected arrival charge exceeds the predetermined threshold value, the signal charge accumulated in the missing pixel at the radiation irradiation end time becomes the predetermined threshold value. An imaging condition calculating step for calculating a new imaging condition for controlling the radiation source;
A feedback control step of performing feedback control of the radiation during the signal charge accumulation step so that the radiation based on the new imaging condition calculated in the imaging condition calculation step is irradiated;
A radiation imaging method comprising:

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