WO2013027816A1

WO2013027816A1 - Radiographic imaging system, radiographic imaging method and radiographic imaging system error-processing method

Info

Publication number: WO2013027816A1
Application number: PCT/JP2012/071385
Authority: WO
Inventors: 北野浩一; 大田恭義; 岩切直人; 西納直行; 中津川晴康
Original assignee: 富士フイルム株式会社
Priority date: 2011-08-25
Filing date: 2012-08-24
Publication date: 2013-02-28

Abstract

In the radiographic imaging system, radiographic imaging method and radiographic imaging system error-processing method of the invention, a radiation dose control unit performs control so that the radiation dose for the subsequent imaging is increased when the gray value of the radiographic image information is lower than a reference value and the radiation dose for the subsequent imaging is decreased when the gray value is higher than the reference value. The system control unit has an error-processing unit that, when results differ from the results anticipated for said gray value, controls the radiation dose control unit and performs control so that at least one radiographic imaging is performed without increasing the radiation dose for the subsequent imaging.

Description

Radiographic imaging system, radiographic imaging method, and error processing method of radiographic imaging system

The present invention relates to a radiographic image capturing system, a radiographic image capturing method, and a radiographic image capturing system capable of obtaining a moving image of a radiographic image by executing radiographic imaging at a set frame rate using a radiographic image capturing apparatus. It relates to an error handling method.

In recent times, during surgery, during contrast imaging, or during treatment of fractures, etc., radiation image information can be read and displayed immediately from the radiation detector after imaging in order to quickly and accurately treat the patient. is required. As a radiation detector capable of meeting such demands, a solid-state detection element (referred to as a pixel) that converts radiation directly into an electrical signal, or converts radiation into visible light with a scintillator and then converts it into an electrical signal for reading. A radiation detector referred to as a flat panel detector (FPD) using the above has been developed.

In particular, an X-ray diagnostic imaging device is proposed in which a radiographic image is displayed on a monitor by executing radiography at a set frame rate, so that, for example, the catheter entry status with respect to the subject can be grasped in real time. Has been.

Conventionally, in such an X-ray diagnostic imaging apparatus, a method for detecting a structural change due to body movement of a subject with a difference between two frames has been proposed (see JP 2007-82908 A).

Also, conventionally, an X-ray diagnostic imaging apparatus using automatic brightness control (Automatic Brightness Control (ABC)) for controlling X-ray irradiation energy during imaging has been proposed (see JP 2011-98009 A).

Further, conventionally, there has been proposed a technique for detecting an abnormality in communication with an operation unit for X-ray irradiation and applying a safe parameter according to an imaging region (see JP 2009-297304 A). .

However, the method described in Japanese Patent Application Laid-Open No. 2007-82908 is carried out to prevent re-imaging due to improper imaging, but detection of exposure failure due to radiation source / generator malfunction and detection of exposure failure. It does not take into account the correspondence of the case, and does not include the concept of the difference between the gray values between the two frames.

The method described in Japanese Patent Application Laid-Open No. 2011-98009 uses a technique related to automatic brightness control (ABC). From the description in paragraphs [0015] and [0019], etc. Only one frame of computation is included. For this reason, it is not possible to carry out abnormality detection using a difference in gray value between a plurality of frames under automatic luminance control.

The method described in Japanese Patent Application Laid-Open No. 2009-297304 has automatic brightness control as an auxiliary means, and when communication is abnormal, exposure is continued by any means including automatic brightness control. With this technology, it is not possible to cope with poor exposure due to malfunction of the tube, radiation source, and generator, and the exposure continues and excessive exposure increases.

As described above, with the techniques described in Japanese Patent Application Laid-Open Nos. 2007-82908 and 2011-98009, it is impossible to detect an exposure failure using the difference in gray value between a plurality of frames in automatic luminance control. With the techniques disclosed in Japanese Patent Application Laid-Open Nos. 2007-82908, 2011-98009, and 2009-297304, it is impossible to take measures to reduce the risk related to the X-ray irradiation dose in the response after the failure is detected. .

That is, the automatic brightness control obtains the gray value (average of QL value, etc.) of the region of interest for each frame, determines whether it is higher or lower than the gray value expected from the imaging region and imaging conditions, This is a method of raising and lowering the X-ray irradiation energy in the next radiography. If any abnormality occurs and the gray value decreases, control for increasing the X-ray irradiation energy is performed without detecting the abnormality. If such control is left unattended, there is a risk of subjecting a subject (such as a patient) to high energy exposure despite the occurrence of an abnormality in the system and the X-ray irradiation control system.

The present invention has been made in consideration of such problems, and in a system using automatic brightness control, when an abnormality occurs in the system or the radiation irradiation control system, the subject (patient etc.) has high irradiation energy. It is an object of the present invention to provide a radiographic imaging system, a radiographic imaging method, and an error processing method of the radiographic imaging system that can prevent exposure to radiation and reduce the risk after an abnormality has occurred. .

[1] A radiographic imaging system according to a first aspect of the present invention includes a radiation apparatus having a radiation source, an irradiation dose control unit for controlling an irradiation dose irradiated from the radiation source, and the radiation source transmitted through the subject. A radiation image capturing device that converts the radiation of the radiation into radiation image information, and a system control unit that controls the radiation image capture device to perform radiation imaging at a set frame rate. The irradiation dose control unit increases the irradiation dose of the next radiation when the gray value of the radiation image information is lower than the reference value, and increases the next radiation dose when the gray value is higher than the reference value. The system control unit controls the irradiation dose control unit when the gray value is different from an expected result. Without increasing the radiation dose of the next radiation, and having an error processing unit for controlling to perform at least one of radiography.

The irradiation dose control unit increases the next radiation irradiation dose when the gray value of the radiation image information is lower than the reference value, and sets the next radiation irradiation dose when the gray value is higher than the reference value. It is controlled to decrease. Under such control, if the gray value is significantly lower than the reference value due to abnormalities in the system or radiation irradiation control system, etc., normally, control is performed to increase the irradiation dose to the maximum. Therefore, the radiation set to the maximum irradiation dose is irradiated to the subject.

In order to avoid this, the first aspect of the present invention increases the irradiation dose of the next radiation by controlling the irradiation dose control unit when the gray value is different from the expected result. And control to perform at least one radiation imaging.

This can prevent the subject (patient, etc.) from being exposed with high irradiation energy when an abnormality occurs in the system or the radiation irradiation control system, and can reduce the risk after the abnormality has occurred. it can.

[2] In the first aspect of the present invention, the system control unit includes a gray level difference acquisition unit that acquires a gray level difference of radiation image information based on at least two times of radiography, and the gray level difference is near a specified value. In the case where the above change has occurred, an error notification unit may be provided that notifies the error that the gray value is different from the expected result.

[3] In the first aspect of the present invention, when the error notification is made, the error processing unit controls the irradiation dose control unit to decrease without increasing the next radiation irradiation dose. Then, control may be performed so that at least one radiation imaging is performed.

[4] In the first aspect of the present invention, the error notification unit is configured such that the gray value of the radiation image information based on the current radiographing is lower than the gray value of the radiographic image information based on the previous radiography by the vicinity of the specified value or more. In this case, the error notification may be performed.

[5] In the first aspect of the present invention, the irradiation dose control unit uses the upper limit signal value Vmax corresponding to the upper limit value of the irradiation dose and the lower limit signal value Vmin corresponding to the lower limit value of the irradiation dose as a dynamic range, and the reference value By outputting to the radiation apparatus a control signal in which the reference signal value corresponding to is Vo and the signal value for controlling the irradiation dose based on the comparison between the gray value and the reference value is V, the radiation source You may make it control the irradiation dose of the radiation from.

[6] When the error notification is performed in [5], the irradiation dose control unit sets the signal value V to be equal to or lower than a reference signal value Vo based on an instruction from the error processing unit. It may be.

[7] In [5], when the error notification is performed, the irradiation dose control unit sets the signal value V to a reference signal value Vo or less based on an instruction from the error processing unit, And you may make it set to the value according to the difference of the said light and shade value.

[8] In [7], the irradiation dose control unit sets the signal value V to the lower limit signal value Vmin × constant Ka (when the difference between the gray values is not less than the specified value and less than the second specified value). 0 <Ka <1.0) may be set, and the signal value V may be set to the lower limit signal value Vmin when the difference between the gray values is equal to or less than the second specified value.

[9] In the first aspect of the present invention, the one-time radiography may be performed at the set frame rate.

[10] In the first aspect of the present invention, the error processing section performs radiation imaging by gradually increasing the radiation dose each time radiation imaging is performed at the frame rate after performing the one radiation imaging. You may make it control to perform.

[11] A radiographic imaging method according to a second aspect of the present invention includes a radiation apparatus having a radiation source, an irradiation dose control unit that controls an irradiation dose irradiated from the radiation source, and the radiation source that has passed through a subject. In a radiographic imaging method of performing radiography at a set frame rate using a radiographic imaging device having a radiation detection device that converts the radiation of the image into radiation image information, the irradiation dose control unit includes the radiation When the gray value of the image information is lower than the reference value, the next radiation dose is increased, and when the gray value is higher than the reference value, the next radiation dose is controlled to be reduced, When the gray value is different from the expected result, the irradiation dose control unit is controlled to increase at least once without increasing the irradiation dose of the next radiation. Characterized in that it has an error processing step of controlling to perform radiography.

[12] An error processing method for a radiographic imaging system according to the third aspect of the present invention includes a radiation apparatus having a radiation source, an irradiation dose control unit that controls an irradiation dose emitted from the radiation source, and transmitted through the subject. In an error processing method for a radiographic imaging system that performs radiography at a set frame rate using a radiographic imaging device having a radiation detection device that converts radiation from the radiation source into radiation image information, The irradiation dose control unit increases the next radiation irradiation dose when the gray value of the radiation image information is lower than the reference value, and sets the next radiation irradiation dose when the gray value is higher than the reference value. If the gray value is different from the expected result, the irradiation dose control unit is controlled and the next radiation is controlled. Without increasing the radiation dose, and having an error processing step of controlling to perform at least one of radiography.

As described above, according to the radiographic image capturing system, the radiographic image capturing method, and the error processing method of the radiographic image capturing system according to the present invention, in the system using the automatic brightness control, the system and the radiation irradiation control system are abnormal. When this occurs, it is possible to prevent the subject (patient or the like) from being exposed with high irradiation energy, and to reduce the risk after the occurrence of an abnormality.

It is a block diagram which shows the radiographic imaging system which concerns on this Embodiment. It is a block diagram which mainly shows the structure of the radiation apparatus and radiation detection apparatus of a radiographic imaging system. It is a circuit diagram which shows the structure of a radiation detection apparatus, and shows the structure of a radiation detector especially. It is a block diagram which mainly shows the structure of the system control part of a radiographic imaging system. It is a flowchart (the 1) which shows the processing operation of a radiographic imaging system. It is a flowchart (the 2) which shows the processing operation of a radiographic imaging system. It is a time chart which shows the processing operation of a radiographic imaging system. It is a figure which shows roughly the structure for 3 pixels of the radiation detector which concerns on a modification. It is a schematic block diagram of TFT shown in FIG. 8 and an electric charge storage part.

Hereinafter, embodiments of the radiographic image capturing system, the radiographic image capturing method, and the error processing method of the radiographic image capturing system according to the present invention will be described with reference to FIGS.

First, as shown in FIG. 1, the radiographic image capturing system 10 according to the present exemplary embodiment includes a radiographic image capturing device 12 and a radiographic image capturing device 12 that are set at a set frame rate (for example, 15 frames / second to 60 frames). And a system control unit 14 that performs control so as to execute radiation imaging at a time of 1 second / second). A console 16 is connected to the system control unit 14 so that data communication with the console 16 is possible. Connected to the console 16 are a monitor 18 for image observation and diagnostic imaging, and an input device 20 (keyboard, mouse, etc.) for operation input. An operator (physician or radiographer) uses the input device 20 to set a radiation exposure dose or a radiographic frame rate suitable for the current situation in an operation or catheter insertion operation while observing a moving image. Data input using the input device 20 and data created and edited by the console 16 are input to the system control unit 14. Further, radiation image information and the like from the system control unit 14 is supplied to the console 16 and displayed on the monitor 18.

The radiographic imaging device 12 includes a radiation device 28 that irradiates radiation 26 toward a subject 24 on an imaging table 22, a radiation detection device 30 that converts radiation 26 transmitted through the subject 24 into radiation image information, and a radiation detection device. A detection device control unit 32 that transmits and receives data such as radiation image information between the system control unit 14 and the system control unit 14, and controls the radiation detection device 30 based on an instruction from the system control unit 14 (including moving drive). Have.

The movement detection of the radiation detection apparatus 30 is performed when a relatively wide range is imaged, for example, a moving image of the spine or a moving image of the catheter entry position. That is, in such imaging, a movement control signal based on an operation input from an operator (doctor or radiographer) is output from the system control unit 14 and input to the detection device control unit 32. Based on the movement control signal from the system control unit 14, the detection device control unit 32 controls the movement drive mechanism (not shown) to move the radiation detection device 30.

As shown in FIG. 2, the radiation device 28 is based on a radiation source 34, a radiation source controller 36 that controls the radiation source 34 based on an instruction from the system controller 14, and an instruction from the system controller 14. And an automatic collimator unit 38 that widens or narrows the irradiation area of the radiation 26.

The radiation detector 30 includes a radiation detector 40, a battery 42 as a power source, a cassette control unit 44 that drives and controls the radiation detector 40, and a signal including radiation image information from the radiation detector 40. A transmitter / receiver 46 for transmitting and receiving data is accommodated. The radiation image information output from the transceiver 46 is input to the system control unit 14 and the console 16 via the detection device control unit 32 and is displayed on the monitor 18. That is, radiation image information based on radiation imaging at a set frame rate is sequentially input to the system control unit 14, and thus a moving image of the radiation image information is displayed on the monitor 18 in real time.

The cassette control unit 44 and the transceiver 46 are provided with lead plates or the like on the irradiation surface side of the cassette control unit 44 and the transceiver 46 in order to avoid damage due to the radiation 26 being irradiated. Is preferred.

As the radiation detector 40, for example, the radiation 26 that has passed through the subject 24 is once converted into visible light by a scintillator, and the converted visible light is a solid-state detection element (hereinafter referred to as “a-Si”). An indirect conversion type radiation detector (including a front side reading method and a back side reading method) that converts to an electric signal can also be used. An ISS (Irradiation Side Sampling) type radiation detector, which is a surface reading method, has a configuration in which a solid detection element and a scintillator are sequentially arranged along the irradiation direction of the radiation 26. A PSS (Penetration Side Sampling) type radiation detector, which is a back side reading method, has a configuration in which a scintillator and a solid state detection element are sequentially arranged along the radiation 26 irradiation direction. Further, as the radiation detector 40, in addition to the above-described indirect conversion type radiation detector, direct conversion in which the dose of the radiation 26 is directly converted into an electric signal by a solid detection element made of a substance such as amorphous selenium (a-Se). A type of radiation detector can be employed.

Next, as an example, the circuit configuration of the radiation detection apparatus 30 when the indirect conversion type radiation detector 40 is employed will be described in detail with reference to FIG.

The radiation detector 40 has a photoelectric conversion layer 52 in which each pixel 50 made of a material such as a-Si that converts visible light into an electrical signal is formed on an array of matrix thin film transistors (hereinafter referred to as TFTs 54). It has the structure arranged in. In this case, in each pixel 50, the charge generated by converting visible light into an electrical signal (analog signal) is accumulated, and the charge can be read out as an image signal by sequentially turning on the TFT 54 for each row. .

A gate line 56 extending in parallel with the row direction and a signal line 58 extending in parallel with the column direction are connected to the TFT 54 connected to each pixel 50. Each gate line 56 is connected to a line scan driver 60, and each signal line 58 is connected to a multiplexer 62. Control signals Von and Voff for controlling on / off of the TFTs 54 arranged in the row direction are supplied from the line scan driving unit 60 to the gate line 56. In this case, the line scan driving unit 60 includes a plurality of switches SW1 for switching the gate lines 56, and a first address decoder 64 for outputting a selection signal for selecting the switches SW1. An address signal is supplied from the cassette control unit 44 to the first address decoder 64.

Further, the charge held in each pixel 50 flows out to the signal line 58 via the TFTs 54 arranged in the column direction. This charge is amplified by the charge amplifier 66. A multiplexer 62 is connected to the charge amplifier 66 through a sample and hold circuit 68.

That is, the read charge of each column is input to the charge amplifier 66 of each column via each signal line 58. Each charge amplifier 66 includes an operational amplifier 70, a capacitor 72, and a switch 74. When the switch 74 is off, the charge amplifier 66 converts the charge signal input to one input terminal of the operational amplifier 70 into a voltage signal and outputs the voltage signal. The charge amplifier 66 amplifies and outputs the electrical signal with the gain set by the cassette control unit 44. Information relating to the gain of the charge amplifier 66 (gain setting information) is supplied from the system control unit 14 to the cassette control unit 44 via the detection device control unit 32. The cassette control unit 44 sets the gain of the charge amplifier 66 based on the supplied gain setting information.

The other input terminal of the operational amplifier 70 is connected to GND (ground potential) (ground). When all the TFTs 54 are turned on and the switch 74 is turned on, the charge accumulated in the capacitor 72 is discharged by the closed circuit of the capacitor 72 and the switch 74, and the charge accumulated in the pixel 50 is closed. It is swept out to GND (ground potential) via the switch 74 and the operational amplifier 70. The operation of turning on the switch 74 of the charge amplifier 66 to discharge the charge accumulated in the capacitor 72 and sweeping out the charge accumulated in the pixel 50 to GND (ground potential) is a reset operation (empty reading operation). Call it. That is, in the reset operation, the voltage signal corresponding to the charge signal stored in the pixel 50 is discarded without being output to the multiplexer 62.

The multiplexer 62 includes a plurality of switches SW2 for switching the signal line 58 and a second address decoder 76 for outputting a selection signal for selecting the switch SW2. An address signal is supplied from the cassette control unit 44 to the second address decoder 76. An A / D converter 78 is connected to the multiplexer 62, and radiation image information converted into a digital signal by the A / D converter 78 is supplied to the cassette control unit 44.

The TFT 54 functioning as a switching element may be realized in combination with another imaging element such as a CMOS (Complementary Metal-Oxide Semiconductor) image sensor. Furthermore, it can be replaced with a CCD (Charge-Coupled Device) image sensor that transfers charges while shifting the charges with a shift pulse corresponding to a gate signal referred to as a TFT.

The cassette control unit 44 of the radiation detection apparatus 30 includes an address signal generation unit 80, an image memory 82, and a cassette ID memory 84, as shown in FIG.

For example, the address signal generator 80 sends an address signal to the first address decoder 64 of the line scan driver 60 and the second address decoder 76 of the multiplexer 62 shown in FIG. 3 based on the read control information from the system controller 14. Supply. The read control information includes, for example, progressive mode, interlace mode (odd row read mode, even row read mode, second row read mode, third row read mode, etc.), binning mode (1 pixel / 4 pixel read mode, 1 pixel / 6-pixel readout mode, 1-pixel / 9-pixel readout mode, etc.) are included. For example, in the 1-pixel / 4-pixel readout mode, two adjacent gate lines are simultaneously activated (set to Von), and two adjacent signal lines are selected at the same time. In this mode, charges for pixels are mixed and read as one pixel. The address signal generator 80 generates an address signal corresponding to the mode indicated by the read control information, and outputs the address signal to the first address decoder 64 of the line scan driver 60 and the second address decoder 76 of the multiplexer 62. The read control information is created by the system control unit 14 based on an operation input from an operator, for example, and is input to the cassette control unit 44 of the radiation detection apparatus 30.

The image memory 82 stores radiation image information detected by the radiation detector 40. The cassette ID memory 84 stores cassette ID information for specifying the radiation detection apparatus 30. The transceiver 46 transmits the cassette ID information stored in the cassette ID memory 84 and the radiation image information stored in the image memory 82 to the system control unit 14 via the detection device control unit 32 by wired communication or wireless communication.

Then, as shown in FIG. 4, the system controller 14 of the radiographic image capturing system 10 calculates the gray value Da of the region of interest of the radiographic image information and stores it in the gray value storage unit 100. And an irradiation dose control unit 104 that performs the same control as the automatic brightness control.

The irradiation dose control unit 104 includes a reference value generation unit 106 that generates a reference value Db corresponding to an imaging region, and an irradiation dose control signal (hereinafter referred to as a first control) according to a difference between the current gray value Da and the reference value Db. A control signal generator 108 that generates a signal Sa1 and outputs it to the radiation device 28, and a control signal (first control signal Sa1 and a later-described first control signal Sa1) and a normal operation period Ta and an error processing period Tb (see FIG. 7). And a switching unit 110 that switches between two control signals Sa2).

The error processing period Tb is a period (for example, 5 to 10 seconds) set in advance from the time ta when the error notification Sb described later is performed. A period other than the error processing period Tb is a normal operation period Ta.

As the gray value Da, an average value of the pixel values (QL values) of all the pixels included in the region of interest is used. The gray value Da is the same concept as the luminance value.

The first control signal Sa1 output from the control signal generator 108 uses the upper limit signal value Vmax corresponding to the upper limit value of the irradiation dose and the lower limit signal value Vmin corresponding to the lower limit value of the irradiation dose as the dynamic range, and is set to the reference value Db. The corresponding reference signal value is Vo, and the signal value corresponding to the difference between the gray value Da and the reference value Db is V. The signal form may be an analog signal or a digital signal. 0 may be used as the reference signal value Vo.

In the normal operation period Ta, the control signal generation unit 108 generates and outputs a signal value (Vo + V) for increasing the irradiation dose of the next radiation imaging when the current gray value Da is lower than the reference value Db. When the current gray value Da is higher than the reference value Db, a signal value (Vo-V) for reducing the irradiation dose of the next radiography is generated and output.

Therefore, in the normal operation period Ta, the irradiation dose control unit 104 increases the next irradiation dose when the gray value Da of the region of interest is lower than the reference value Db, and the gray value Da of the region of interest is higher than the reference value Db. If it is high, control to reduce the next irradiation dose.

In addition, the system control unit 14 includes a shading difference acquisition unit 112, an error notification unit 114, and an error processing unit 116.

The density difference acquisition unit 112 acquires the difference ΔD of the density value Da of the region of interest based on at least two radiation irradiations.

The error notification unit 114 performs error notification Sb, assuming that an error has occurred when the obtained gray value is different from the expected result. The result different from the expected result is that normally, in the irradiation dose control, the radiation device 28 is feedback-controlled so that the difference d between the obtained gray value Da and the reference value Db is reduced. Da is normally a value approximated to the reference value Db. For example, when the maximum value of the gray value Da is set to +128, the result is expected to be within ± 10 with respect to the reference value Db. . In such a case, when the gray value Da suddenly becomes 0 or almost 0, or becomes the maximum value or almost the maximum value, the gray value Da is different from the expected result. This is considered to be based on an abnormality in the radiation device 28, an exposure failure due to a malfunction of the generator, or an abnormality in the radiation detection device 30, the detection device control unit 32, or the like.

Therefore, the error notification unit 114 sets the specified value Dc in advance to determine whether or not the obtained gray value Da is a result different from the expected result, and the gray value difference ΔD is determined. An error notification (Sb) is made when an error has occurred when the value changes near the prescribed value Dc. The vicinity of the specified value Dc refers to, for example, a range from the specified value Dc × (1.0−coefficient Kc) to the specified value Dc × (1.0 + coefficient Kc). Here, the coefficient Kc is 0 <Kc <1.0, and is set in advance by simulation or experiment. In the present embodiment, for example, a range of 0.1 to 0.2 can be used as the coefficient Kc.

In the present embodiment, the difference ΔD between the gray values without using the difference d between the gray value Da and the reference value Db is based on the fact that the radiographic image information between frames has high correlation. . For example, in the catheter insertion operation, even when the region of interest moves from, for example, a region having a generally low gray value to a region having a high gray value, the gray value difference ΔD is substantially constant. By comparing the value difference ΔD with the vicinity of the specified value, an error can be reliably detected.

When the error notification Sb is performed, the error processing unit 116 controls the irradiation dose control unit 104 so as to reduce at least one radiation imaging without increasing the next radiation irradiation dose. To control. Of course, the radiation imaging may be performed with the irradiation dose immediately before the error notification without decreasing.

The error notification unit 114 performs the error notification Sb particularly when the gray value Da of the region of interest based on the current radiographing is lower than the gray value Da of the region of interest based on the previous radiography by the vicinity of the specified value Dc. . As the prescribed value Dc, a value of 25% or more, preferably 30% or more of 1/2 of the maximum density value (maximum density value) can be used. Further, regarding the relationship between the radiation exposure amount to the radiation detector 40 and the gray value Da of the radiation image, the gray value Da decreases as the exposure amount to the radiation detector 40 decreases. Therefore, the gray value Da of the region of interest based on the current radiography is lower than the gray value Da of the region of interest based on the previous radiography by at least the vicinity of the specified value Dc. Indicates a white image. This is considered to be based on the radiation source 34 and the error of poor exposure due to the malfunction of the generator.

The error processing unit 116 includes a switching signal output unit 118 and a control signal calculation unit 120.

The switching signal output unit 118 outputs the switching signal Sc to the switching unit 110 when the error notification Sb is received and when the error processing period Tb has elapsed.

When the system is activated, the switching unit 110 sets the signal path of the control signal to the control signal generation unit 108 side of the irradiation dose control unit 104, and the first control signal Sa1 from the control signal generation unit 108 is sent to the radiation apparatus 28. The signal path of the control signal is switched to the control signal calculation unit 120 side of the error processing unit 116 based on the input of the switching signal Sc from the switching signal output unit 118 based on the error notification Sb. The second control signal Sa2 from the control signal calculation unit 120 is supplied to the radiation device 28. Further, based on the input of the switching signal Sc from the switching signal output unit 118 based on the passage of the error processing period Tb, the signal path of the control signal is switched again to the control signal generation unit 108 side of the irradiation dose control unit 104.

Meanwhile, the control signal calculation unit 120 calculates the second control signal Sa2 for each frame in the error processing period Tb and outputs the second control signal Sa2 for each frame. The second control signal Sa2 has the same signal form as the first control signal Sa1 described above.

Here, as an example of a calculation method in the control signal calculation unit 120, the signal value V1 of the second control signal Sa2 in the first radiography after the error notification Sb is given is, for example, the following four types. It is done.

(1) The signal value V1 is set to be equal to or less than the reference signal value Vo.

(2) Set the signal value V1 to the lower limit signal value Vmin.

(3) The signal value V1 is set to a value that is equal to or less than the reference signal value Vo and that corresponds to the difference ΔD of the gray value.

(4) When the difference ΔD between the gray levels is greater than or equal to the specified value Dc and less than the second specified value Dd, the signal value V1 is set to the lower limit signal value Vmin × constant Ka (0 <Ka <1.0). When the value difference ΔD is equal to or smaller than the second specified value Dd, the signal value V1 is set to the lower limit signal value Vmin.

Further, the signal value Vj of the second control signal Sa2 in the second and subsequent radiation imaging in the error processing period Tb can be obtained as follows.

That is, Va is the signal value generated for the radiography immediately before the error notification Sb, F is the number of frames in the error processing period Tb, and j (j = 2, 3,..., The signal value Vj can be obtained by the following arithmetic expressions (A) and (B).
Signal value Vj = V1 + ΔV × (j−1) (A)
ΔV = (Va−V1) / F (B)

As a result, the error processing unit 116 receives the error notification Sb, performs the first radiography, and then gradually increases the irradiation dose every time radiography is performed at the currently set frame rate. It will be controlled to perform. When the error processing period Tb elapses, the signal value Va almost returns immediately before the error notification Sb, and the normal operation can be smoothly performed.

Here, the processing operation of the radiographic imaging system 10 will be described with reference to the flowcharts of FIGS. 5 and 6 and the time chart of FIG.

First, in step S1 of FIG. 5, the system control unit 14 stores an initial value (= 1) in the counter k of the number of times of shooting at normal times.

In step S2, the control signal generation unit 108 of the irradiation dose control unit 104 sets the signal value V to the reference signal value Vo. For example, 0 is used as the reference signal value Vo.

In step S3, the system control unit 14 determines whether a time corresponding to the latest frame rate has elapsed since the start of the previous radiation imaging. When the value of the counter k is an initial value or when the time corresponding to the latest frame rate has elapsed since the start of the previous radiation imaging, the process proceeds to the next step S4, and the control signal generator 108 determines the current signal value. For example, V is stored in a register as the latest signal value Va.

In step S5, the control signal generation unit 108 outputs the current signal value V to the radiation device 28 via the switching unit 110. The radiation source control unit 36 of the radiation apparatus 28 changes the tube voltage, tube current, imaging time, etc. of the radiation source 34 based on the signal value V from the irradiation dose control unit 104, and outputs the radiation source 34. Set the new exposure dose.

In step S6, the system control unit 14 outputs an exposure start signal Sd to the radiation device 28 at the start of the k-th radiation imaging. The radiation source control unit 36 of the radiation apparatus 28 controls the radiation source 34 based on the input of the exposure start signal Sd from the system control unit 14, and irradiates the radiation with the irradiation dose set from the radiation source 34. Let

In step S <b> 7, the system control unit 14 outputs an exposure notification Se indicating that the radiation device 28 has started exposure to the radiation detection device 30 via the detection device control unit 32.

In step S8, the radiation detection apparatus 30 performs charge accumulation and charge reading based on the input of the exposure notification Se. That is, the radiation 26 that has passed through the subject 24 is once converted into visible light by the scintillator, and the visible light is photoelectrically converted in each pixel 50 to accumulate an amount of electric charge corresponding to the amount of light. Then, a synchronization signal (for example, a vertical synchronization signal) is output at the start of the reading period and input to the detection device control unit 32. The detection device controller 32 synchronizes the reception timing of the radiation image information with the output timing of the radiation image information from the radiation detection device 30 based on the input of the synchronization signal.

In the subsequent readout period, the radiation detection apparatus 30 reads out charges in accordance with currently set readout control information (information indicating progressive mode, interlace mode, and binning mode), and uses the image memory 82, for example, in a FIFO manner. Outputs radiation image information. Radiation image information from the radiation detection device 30 is supplied to the system control unit 14 via the detection device control unit 32.

In step S <b> 9, the gray value acquisition unit 102 calculates the gray value Da of the region of interest (for example, the average value of the pixel values (QL values) of all the pixels included in the region of interest) in the radiographic image information. Store in the storage unit 100.

In step S10, the system control unit 14 transfers the supplied radiation image information to the console 16. The console 16 stores the transferred radiation image information in the frame memory, and displays it on the monitor 18 as a radiation image obtained by the k-th radiation imaging, that is, a radiation image of the k-th frame in the normal operation period.

In step S11, the light / dark difference obtaining unit 112 obtains the light / dark value Da difference ΔD of the region of interest based on the two times of radiation irradiation. For example, when the latest density value (the density value based on the current radiography) stored in the density value storage unit 100 is Da1, and the latest previous density value (the density value based on the previous radiography) is Da2. The difference ΔD = Da1−Da2 is calculated.

In step S12, it is determined whether or not the gray value Da1 is lower than the gray value Da2 by the vicinity of the specified value Dc. This determination is made based on whether or not the difference ΔD is negative and the absolute value | ΔD | ≧ Dc.

When the difference ΔD is positive or the absolute value | ΔD | <Dc, the process proceeds to step S13, and the control signal generation unit 108 of the irradiation dose control unit 104 sets the difference d between the current gray value Da and the reference value Db. Calculate. That is, the difference d = Da−Db is calculated.

In step S14, the control signal generation unit 108 determines whether or not the difference d <0. If the difference d is negative, the process proceeds to step S15, and the signal value V is increased according to the difference d. On the contrary, if the difference d is positive, the process proceeds to step S16, and the signal value V is decreased according to the difference d.

In step S17, the value of the counter k is updated by +1.

In step S18, the system control unit 14 determines whether or not there is a system termination request. If there is no system termination request, the process returns to step S3, and the processes after step S3 are repeated. Until it is determined in step S12 that the gray value Da1 is lower than the gray value Da2 by the vicinity of the specified value Dc, the operations in steps S3 to S18 are repeated, and the radiation image at the set frame rate is displayed on the monitor 18. Will be displayed.

When it is determined in step S12 that the gray value Da1 is lower than the gray value Da2 by the vicinity of the specified value Dc, the process proceeds to step S19 in FIG. 6 and the switching signal output unit 118 switches to the switching unit 110. The signal Sc is output. As a result, the signal path of the control signal is switched to the control signal calculation unit 120 side of the error processing unit 116.

In step S20, the system control unit 14 stores an initial value (= 1) in the counter j of the number of times of photographing at the time of error processing.

In step S21, the system control unit 14 determines whether or not a time corresponding to the latest frame rate has elapsed since the start of the previous radiation imaging. When the time has elapsed since the start of the previous radiation imaging, the process proceeds to the next step S22, and whether or not the first radiation imaging is entered in the error processing period Tb, that is, the value of the counter j is an initial value. It is determined whether or not.

If the value of the counter j is the initial value, the process proceeds to the next step S23, and the control signal calculation unit 120 calculates the signal value V1 of the second control signal Sa2 in the first radiation imaging. As the calculation method, any one of (1) to (4) described above can be selected.

In step S24, the control signal calculation unit 120 outputs the signal value V1 to the radiation device 28 via the switching unit 110. The radiation source control unit 36 of the radiation apparatus 28 changes the tube voltage, tube current, imaging time, etc. of the radiation source 34 based on the signal value V1 from the error processing unit 116 and outputs the radiation source 34 from the radiation source 34. Set the exposure dose to a new exposure dose.

On the other hand, if the value of the counter j is other than the initial value, the process proceeds to step S25, and the control signal calculation unit 120 calculates the signal value Vj of the second control signal Sa2 in the second and subsequent radiographs from the above-described calculation formula (A). And (B).

In step S26, the control signal calculation unit 120 outputs the signal value Vj to the radiation device 28 via the switching unit 110. The radiation source control unit 36 of the radiation apparatus 28 sets the irradiation dose based on the signal value Vj from the error processing unit 116.

When the process in step S24 or the process in step S26 is completed, the process proceeds to the next step S27, and the system control unit 14 sends an exposure start signal Sd to the radiation apparatus 28 at the start of the j-th radiation imaging. Is output. The radiation source control unit 36 of the radiation apparatus 28 controls the radiation source 34 based on the input of the exposure start signal Sd from the system control unit 14, and irradiates the radiation with the irradiation dose set from the radiation source 34. Let

In step S28, the system control unit 14 outputs an exposure notification Se indicating that the radiation device 28 has started exposure to the radiation detection device 30 via the detection device control unit 32.

In step S29, the radiation detection apparatus 30 performs charge accumulation and charge readout based on the input of the exposure notification Se. Since this operation is the same as the operation in step S8 described above, a duplicate description thereof is omitted here.

In step S30, the system control unit 14 transfers the supplied radiation image information to the console 16. The console 16 stores the transferred radiation image information in the frame memory and displays it on the monitor 18 as a radiation image obtained by the j-th radiation imaging, that is, a radiation image of the jth frame in the error processing period.

In step S31, the value of the counter j is updated by +1.

In step S32, the switching signal output unit 118 determines whether or not the error processing period Tb has elapsed. If the error processing period Tb has not elapsed, the process returns to step S21, and the processes after step S21 are repeated.

When the error processing period Tb has elapsed, the process proceeds to the next step S33, and the switching signal output unit 118 outputs the switching signal Sc to the switching unit 110. As a result, the signal path of the control signal is switched again to the control signal generation unit 108 side of the irradiation dose control unit 104.

Thereafter, the process returns to step S3 in FIG. 5 and the processes after step S3 are repeated.

Then, in the above-described step S18, when it is determined that there is a system termination request, the processing in the radiation image capturing system 10 is terminated.

In the example of FIG. 7, for example, radiation imaging is performed at the start time tn + 1 of the (N + 1) th radiation imaging in the normal operation period Ta, so that the radiation image information D (N + 1) obtained by the N + 1th radiation imaging is transmitted to the system control unit 14. ) Is supplied. The system control unit 14 transfers the supplied radiation image information D (N + 1) to the console 16 and displays it on the monitor 18 as a radiation image of the Nth frame. In addition, the system control unit 14 calculates a difference ΔD between the previous (Nth) gray value and the current (N + 1) gray value. Since the difference ΔD (indicated by “−2” for convenience) is not a value that has decreased by more than a predetermined value (for example, “−50”), the normal operation is continued and the system control unit 14 changes the density of the region of interest. The signal value V of the first control signal Sa1 corresponding to the difference between the value Da and the reference value Db is calculated and output to the radiation device 28. The radiation device 28 sets the irradiation dose of the radiation source 34 to an amount corresponding to the signal value V (for convenience, indicated by “+2”) for the next N + 2th radiography.

Similarly, N + 2th radiography and N + 3th radiography are sequentially performed, and the radiation apparatus 28 determines the radiation dose of the radiation source 34 as a signal value V (for convenience) for the next N + 4th radiography. The amount is set according to “−1”.

Then, the gray value Da of the region of interest obtained by the N + 4th radiography is approximately 0, for example, and the signal value V of the first control signal Sa1 corresponding to the difference between the gray value Da and the reference value Db is, for example, the maximum value ( For convenience, the maximum value Vmax is output to the radiation device 28 and set to the maximum irradiation dose. As a result, in the next N + 5th radiography, the subject 24 is irradiated with the radiation 26 set to the maximum irradiation dose.

In order to avoid this, in the present embodiment, the system control unit 14 calculates a difference ΔD between the previous (N + 3) gray value and the current (N + 4) gray value. Since the difference ΔD (denoted by “−128” for convenience) is a value that has decreased by more than a specified value (for example, “−50”), the operation of the error processing unit 116 is started, and the system control unit 14 Then, the signal value V1 of the second control signal Sa2 of the first radiation imaging in the error processing period Tb is calculated and output to the radiation device 28. The radiation device 28 sets the irradiation dose of the radiation source 34 to an amount corresponding to the signal value V1 (for example, the lower limit signal value “Vmin”) for the first radiation imaging in the error processing period Tb. As a result, in the first radiography in the error processing period Tb, the radiation 26 set to the minimum irradiation dose is irradiated to the subject 24, and the burden of exposure on the subject 24 can be reduced.

In the second and subsequent radiation imaging in the error processing period Tb, the irradiation dose gradually increases for each radiation imaging by the processing in step S25 and step S26, and error processing is performed when the error processing period Tb elapses. Radiation imaging is performed with the irradiation dose immediately before the start of the period Tb, and the normal operation is smoothly performed.

As described above, in the radiographic imaging system 10 according to the present exemplary embodiment, in the system using automatic luminance control, for example, when an abnormality occurs in the system or the radiation irradiation control system, the subject 24 (patient or the like) Exposure to high irradiation energy can be prevented, and the risk after an abnormality has occurred can be reduced.

In the radiographic imaging system 10, another preferable mode is shown. At the start of the error processing period Tb, an instruction to narrow the irradiation area is output from the system control unit 14 to the automatic collimator unit 38 of the radiation apparatus 28. May be. For example, it is narrowed within a range of ¼ to 1/10 of the irradiation region immediately before the start of the error processing period Tb. This ratio is set in advance by simulation or experiment according to the imaging region or the like. Thereby, the burden concerning the exposure to the subject 24 can be further reduced.

Also, at the start of the error processing period Tb, an instruction to increase the gain of the charge amplifier 66 may be output from the system control unit 14 to the radiation detection device 30 via the detection device control unit 32. Thereby, even if sensitivity improves and irradiation energy is set low, the radiographic image information which has the same width of light and shade as usual can be obtained. It can be expected that this can be effectively used for observation and diagnosis even for a moving image of low irradiation energy displayed in the error processing period Tb.

Further, at the start of the error processing period Tb, the reading mode in the radiation detection apparatus 30 may be set to, for example, an interlace mode. The burden on the signal processing system related to charge readout in the radiation detection apparatus 30 can be reduced, and the risk that an error will occur again can be reduced.

Of course, the radiographic imaging system according to the present invention is not limited to the above-described embodiment, and various configurations can be adopted without departing from the gist of the present invention.

For example, the radiation detector 40 may be the radiation detector 600 according to the modification shown in FIGS. FIG. 8 is a schematic cross-sectional view schematically showing the configuration of three pixel portions of the radiation detector 600 according to the modification.

As shown in FIG. 8, the radiation detector 600 includes a signal output unit 604, a sensor unit 606 (photoelectric conversion unit), and a scintillator 608 sequentially stacked on an insulating substrate 602. A pixel unit is configured by the sensor unit 606. A plurality of pixel portions are arranged in a matrix on the substrate 602, and the signal output portion 604 and the sensor portion 606 in each pixel portion are configured to overlap each other.

The scintillator 608 is formed on the sensor unit 606 with a transparent insulating film 610 interposed therebetween. The scintillator 608 converts the radiation 26 incident from above (the side opposite to the side where the substrate 602 is located) into light and emits light. The body is formed into a film. The wavelength range of light emitted by the scintillator 608 is preferably the visible light range (wavelength 360 nm to 830 nm), and in order to enable monochrome imaging by the radiation detector 600, the wavelength range of green is included. Is more preferable.

Specifically, the phosphor used in the scintillator 608 preferably contains cesium iodide (CsI) when imaging using X-rays as the radiation 26, and the emission spectrum upon X-ray irradiation is 420 nm to 700 nm. It is particularly preferred to use some CsI (Tl) (cesium iodide with thallium added). Note that the emission peak wavelength of CsI (Tl) in the visible light region is 565 nm.

The scintillator 608 may be formed, for example, by vapor-depositing CsI (Tl) having a columnar crystal structure on a vapor deposition base. When the scintillator 608 is formed by vapor deposition as described above, Al is often used as the vapor deposition substrate from the viewpoint of X-ray transmittance and cost, but is not limited thereto. Note that in the case where GOS is used as the scintillator 608, the scintillator 608 may be formed by applying GOS to the surface of the TFT active matrix substrate without using a vapor deposition substrate. Alternatively, after the GOS is applied to the resin base to form the scintillator 608, the scintillator 608 may be bonded to the TFT active matrix substrate. As a result, the TFT active matrix substrate can be preserved even if GOS application fails.

The sensor unit 606 includes an upper electrode 612, a lower electrode 614, and a photoelectric conversion film 616 disposed between the upper electrode 612 and the lower electrode 614.

Since the upper electrode 612 needs to make the light generated by the scintillator 608 incident on the photoelectric conversion film 616, it is preferable that the upper electrode 612 is made of a conductive material that is transparent at least with respect to the emission wavelength of the scintillator 608. It is preferable to use a transparent conductive oxide (TCO) having a high transmittance for visible light and a low resistance value. Note that although a metal thin film such as Au can be used as the upper electrode 612, a resistance value tends to increase when the transmittance of 90% or more is obtained, so that the TCO is preferable. For example, ITO, IZO, AZO, FTO, SnO ₂ , TiO ₂ , ZnO _{2 and the} like can be preferably used, and ITO is most preferable from the viewpoint of process simplicity, low resistance, and transparency. Note that the upper electrode 612 may have a single configuration common to all the pixel portions, or may be divided for each pixel portion.

The photoelectric conversion film 616 includes an organic photoconductor (OPC: Organic Photo Conductors), absorbs light emitted from the scintillator 608, and generates a charge corresponding to the absorbed light. If the photoelectric conversion film 616 includes an organic photoconductor (organic photoelectric conversion material), the photoelectric conversion film 616 has a sharp absorption spectrum in the visible light region, and electromagnetic waves other than light emitted by the scintillator 608 are almost absorbed by the photoelectric conversion film 616. In addition, noise generated when the radiation 26 is absorbed by the photoelectric conversion film 616 can be effectively suppressed. Note that the photoelectric conversion film 616 may be configured to include amorphous silicon instead of the organic photoconductor. In this case, it has a wide absorption spectrum and can efficiently absorb light emitted by the scintillator 608.

The organic photoconductor constituting the photoelectric conversion film 616 preferably has a peak wavelength closer to the emission peak wavelength of the scintillator 608 in order to absorb light emitted by the scintillator 608 most efficiently. Ideally, the absorption peak wavelength of the organic photoconductor coincides with the emission peak wavelength of the scintillator 608. However, if the difference between the two is small, the light emitted from the scintillator 608 can be sufficiently absorbed. . Specifically, the difference between the absorption peak wavelength of the organic photoconductor and the emission peak wavelength of the scintillator 608 with respect to the radiation 26 is preferably within 10 nm, and more preferably within 5 nm.

Examples of organic photoconductors that can satisfy such conditions include quinacridone organic compounds and phthalocyanine organic compounds. For example, since the absorption peak wavelength in the visible region of quinacridone is 560 nm, if quinacridone is used as the organic photoconductor and CsI (Tl) is used as the material of the scintillator 608, the difference between the peak wavelengths can be within 5 nm. Thus, the amount of charge generated in the photoelectric conversion film 616 can be substantially maximized.

The sensor unit 606 is a stack of a part that absorbs electromagnetic waves, a photoelectric conversion part, an electron transport part, a hole transport part, an electron blocking part, a hole blocking part, a crystallization prevention part, an electrode, an interlayer contact improvement part, or the like. An organic layer formed by mixing is included. The organic layer preferably contains an organic p-type compound (organic p-type semiconductor) or an organic n-type compound (organic n-type semiconductor).

An organic p-type semiconductor is a donor organic semiconductor (compound) typified by a hole-transporting organic compound and refers to an organic compound having a property of easily donating electrons. More specifically, an organic compound having a smaller ionization potential when two organic materials are used in contact with each other. Therefore, any organic compound can be used as the donor organic compound as long as it is an electron-donating organic compound.

Organic n-type semiconductors are acceptor organic semiconductors (compounds) typified mainly by electron-transporting organic compounds and refer to organic compounds that have the property of easily accepting electrons. More specifically, the organic compound having the higher electron affinity when two organic compounds are used in contact with each other. Therefore, any organic compound can be used as the acceptor organic compound as long as it is an electron-accepting organic compound.

Since the materials applicable as the organic p-type semiconductor and organic n-type semiconductor and the configuration of the photoelectric conversion film 616 are described in detail in Japanese Patent Application Laid-Open No. 2009-32854, description thereof is omitted. Note that the photoelectric conversion film 616 may be formed by further containing fullerenes or carbon nanotubes.

The thickness of the photoelectric conversion film 616 is preferably as large as possible in terms of absorbing light from the scintillator 608. However, when the thickness is larger than a certain level, the photoelectric conversion film 616 is generated in the photoelectric conversion film 616 by a bias voltage applied from both ends of the photoelectric conversion film 616. Since electric field strength is reduced and charges cannot be collected, the thickness is preferably 30 nm to 300 nm, more preferably 50 nm to 250 nm, and particularly preferably 80 nm to 200 nm.

The photoelectric conversion film 616 has a single configuration common to all pixel portions, but may be divided for each pixel portion. The lower electrode 614 is a thin film divided for each pixel portion. However, the lower electrode 614 may have a single configuration common to all the pixel portions. The lower electrode 614 can be made of a transparent or opaque conductive material, and aluminum, silver, or the like can be preferably used. The thickness of the lower electrode 614 can be, for example, 30 nm or more and 300 nm or less.

In the sensor unit 606, by applying a predetermined bias voltage between the upper electrode 612 and the lower electrode 614, one of charges (holes, electrons) generated in the photoelectric conversion film 616 is moved to the upper electrode 612. The other can be moved to the lower electrode 614. In the radiation detector 600 according to this modification, a wiring is connected to the upper electrode 612, and a bias voltage is applied to the upper electrode 612 via the wiring. In addition, the polarity of the bias voltage is determined so that electrons generated in the photoelectric conversion film 616 move to the upper electrode 612 and holes move to the lower electrode 614, but this polarity is opposite. May be.

The sensor unit 606 constituting each pixel unit only needs to include at least the lower electrode 614, the photoelectric conversion film 616, and the upper electrode 612. In order to suppress an increase in dark current, the electron blocking film 618 and the hole blocking are included. It is preferable to provide at least one of the films 620, and it is more preferable to provide both.

The electron blocking film 618 can be provided between the lower electrode 614 and the photoelectric conversion film 616. When a bias voltage is applied between the lower electrode 614 and the upper electrode 612, electrons are transferred from the lower electrode 614 to the photoelectric conversion film 616. It is possible to suppress the dark current from increasing due to the injection of.

An electron donating organic material can be used for the electron blocking film 618. The material actually used for the electron blocking film 618 may be selected according to the material of the adjacent electrode, the material of the adjacent photoelectric conversion film 616, and the like, and 1.3 eV or more from the work function (Wf) of the material of the adjacent electrode. Those having a large electron affinity (Ea) and an Ip equivalent to or smaller than the ionization potential (Ip) of the material of the adjacent photoelectric conversion film 616 are preferable. Since the material applicable as the electron donating organic material is described in detail in Japanese Patent Application Laid-Open No. 2009-32854, description thereof is omitted.

The thickness of the electron blocking film 618 is preferably 10 nm or more and 200 nm or less, more preferably 30 nm or more and 150 nm or less, and particularly preferably, in order to surely exhibit the dark current suppressing effect and prevent a decrease in photoelectric conversion efficiency of the sensor unit 606. It is good to set it to 50 nm or more and 100 nm or less.

The hole blocking film 620 can be provided between the photoelectric conversion film 616 and the upper electrode 612. When a bias voltage is applied between the lower electrode 614 and the upper electrode 612, the hole blocking film 620 is applied from the upper electrode 612 to the photoelectric conversion film 616. It is possible to suppress the increase in dark current due to the injection of holes.

An electron-accepting organic material can be used for the hole blocking film 620. The thickness of the hole blocking film 620 is preferably 10 nm or more and 200 nm or less, more preferably 30 nm or more and 150 nm or less, and particularly preferably, in order to reliably exhibit the dark current suppressing effect and prevent a decrease in photoelectric conversion efficiency of the sensor unit 606. Is preferably 50 nm to 100 nm.

The material actually used for the hole blocking film 620 may be selected according to the material of the adjacent electrode, the material of the adjacent photoelectric conversion film 616, and the like, and 1.3 eV from the work function (Wf) of the material of the adjacent electrode. As described above, it is preferable that the ionization potential (Ip) is large and the Ea is equal to or larger than the electron affinity (Ea) of the material of the adjacent photoelectric conversion film 616. Since the material applicable as the electron-accepting organic material is described in detail in Japanese Patent Application Laid-Open No. 2009-32854, description thereof is omitted.

Note that, among the charges generated in the photoelectric conversion film 616, when a bias voltage is set so that holes move to the upper electrode 612 and electrons move to the lower electrode 614, the electron blocking film 618 and the hole blocking are set. The position of the film 620 may be reversed. Further, it is not necessary to provide both the electron blocking film 618 and the hole blocking film 620. If either one is provided, a certain dark current suppressing effect can be obtained.

As shown in FIG. 9, the signal output unit 604 is provided on the surface of the substrate 602 corresponding to the lower electrode 614 of each pixel unit, and the storage capacitor 622 that accumulates the electric charge moved to the lower electrode 614, The TFT 624 converts the electric charge accumulated in the accumulation capacitor 622 into an electric signal and outputs the electric signal. The region where the storage capacitor 622 and the TFT 624 are formed has a portion that overlaps with the lower electrode 614 in plan view. With such a structure, the signal output unit 604 and the sensor unit 606 in each pixel unit are connected to each other. There will be overlap in the thickness direction. If the signal output unit 604 is formed so as to completely cover the storage capacitor 622 and the TFT 624 with the lower electrode 614, the plane area of the radiation detector 600 (pixel unit) can be minimized.

The storage capacitor 622 is electrically connected to the corresponding lower electrode 614 through a wiring made of a conductive material that penetrates an insulating film 626 provided between the substrate 602 and the lower electrode 614. Thereby, the charge collected by the lower electrode 614 can be moved to the storage capacitor 622.

In the TFT 624, a gate electrode 628, a gate insulating film 630, and an active layer (channel layer) 632 are stacked, and a source electrode 634 and a drain electrode 636 are formed on the active layer 632 with a predetermined interval. The active layer 632 can be formed of, for example, amorphous silicon, amorphous oxide, organic semiconductor material, carbon nanotube, or the like. Note that the material forming the active layer 632 is not limited thereto.

The amorphous oxide that can form the active layer 632 is preferably an oxide containing at least one of In, Ga, and Zn (for example, In—O-based), and at least two of In, Ga, and Zn. Oxides containing one (eg, In—Zn—O, In—Ga—O, and Ga—Zn—O) are more preferred, and oxides containing In, Ga, and Zn are particularly preferred. As the In—Ga—Zn—O-based amorphous oxide, an amorphous oxide whose composition in a crystalline state is represented by InGaO ₃ (ZnO) _m (m is a natural number less than 6) is preferable, and InGaZnO is particularly preferable. ₄ is more preferable. Note that the amorphous oxide that can form the active layer 632 is not limited thereto.

Examples of the organic semiconductor material that can form the active layer 632 include, but are not limited to, phthalocyanine compounds, pentacene, vanadyl phthalocyanine, and the like. The configuration of the phthalocyanine compound is described in detail in Japanese Patent Application Laid-Open No. 2009-212389, so that the description thereof is omitted.

If the active layer 632 of the TFT 624 is formed of an amorphous oxide, an organic semiconductor material, or a carbon nanotube, the radiation 26 such as X-rays is not absorbed, or even if it is absorbed, a very small amount remains. Generation of noise in the unit 604 can be effectively suppressed.

Further, when the active layer 632 is formed of carbon nanotubes, the switching speed of the TFT 624 can be increased, and a TFT 624 having a low light absorption in the visible light region can be formed. In addition, when the active layer 632 is formed of carbon nanotubes, the performance of the TFT 624 is remarkably deteriorated only by mixing a very small amount of metallic impurities into the active layer 632, so that extremely high purity carbon nanotubes are separated by centrifugation or the like.・ It needs to be extracted and formed.

Here, any of the above-described amorphous oxide, organic semiconductor material, carbon nanotube, and organic photoconductor can be formed at a low temperature. Therefore, the substrate 602 is not limited to a substrate having high heat resistance such as a semiconductor substrate, a quartz substrate, and a glass substrate, and a flexible substrate such as plastic, aramid, or bionanofiber can also be used. Specifically, flexible substrates such as polyesters such as polyethylene terephthalate, polybutylene phthalate, polyethylene naphthalate, polystyrene, polycarbonate, polyethersulfone, polyarylate, polyimide, polycycloolefin, norbornene resin, polychlorotrifluoroethylene, etc. Can be used. If such a plastic flexible substrate is used, it is possible to reduce the weight, which is advantageous for carrying around, for example.

In addition, the photoelectric conversion film 616 is formed from an organic photoconductor, and the TFT 624 is formed from an organic semiconductor material, whereby the photoelectric conversion film 616 and the TFT 624 are formed at a low temperature on a plastic flexible substrate (substrate 602). It is possible to reduce the thickness and weight of the radiation detector 600 as a whole. Thereby, the radiation detection apparatus 30 that accommodates the radiation detector 600 can be made thinner and lighter, and convenience in use outside the hospital is improved. In addition, since the base material of the photoelectric conversion portion is made of a material having flexibility different from that of general glass, it is possible to improve damage resistance when the device is carried or used.

In addition, the substrate 602 is provided with an insulating layer for ensuring insulation, a gas barrier layer for preventing permeation of moisture and oxygen, an undercoat layer for improving flatness or adhesion to electrodes, and the like. May be.

Since aramid can be applied to a high-temperature process of 200 ° C. or higher, the transparent electrode material can be cured at a high temperature to lower its resistance, and can also be used for automatic mounting of a driver IC including a solder reflow process. Moreover, since aramid has a thermal expansion coefficient close to that of ITO (Indium Tin Oxide) or a glass substrate, warping after manufacturing is small and it is difficult to crack. In addition, aramid can form a substrate thinner than a glass substrate or the like. Note that the substrate 602 may be formed by stacking an ultrathin glass substrate and an aramid.

Bionanofiber is a composite of a cellulose microfibril bundle (bacterial cellulose) produced by bacteria (acetobacterium Xylinum) and a transparent resin. The cellulose microfibril bundle has a width of 50 nm and a size of 1/10 of the visible light wavelength, and has high strength, high elasticity, and low thermal expansion. By impregnating and curing a transparent resin such as acrylic resin or epoxy resin into bacterial cellulose, a bio-nanofiber having a light transmittance of about 90% at a wavelength of 500 nm can be obtained while containing 60-70% of the fiber. Bionanofiber has a low coefficient of thermal expansion (3-7ppm) comparable to silicon crystals, and is as strong as steel (460MPa), highly elastic (30GPa), and flexible. Compared to glass substrates, etc. Thus, a thin substrate 602 can be formed.

In this modification, a signal output unit 604, a sensor unit 606, and a transparent insulating film 610 are sequentially formed on a substrate 602, and a scintillator 608 is attached to the substrate 602 using an adhesive resin having low light absorption. Thus, the radiation detector 600 is formed.

In the radiation detector 600 according to the above-described modification, the photoelectric conversion film 616 is made of an organic photoconductor, and the active layer 632 of the TFT 624 is made of an organic semiconductor material. Therefore, the photoelectric conversion film 616 and the signal output unit 604 are used. Therefore, the radiation 26 is hardly absorbed. Thereby, the fall of the sensitivity with respect to the radiation 26 can be suppressed.

Both the organic semiconductor material constituting the active layer 632 of the TFT 624 and the organic photoconductor constituting the photoelectric conversion film 616 can be formed at a low temperature. Therefore, the substrate 602 can be formed of a plastic resin, aramid, or bionanofiber that absorbs less radiation 26. Thereby, the fall of the sensitivity with respect to the radiation 26 can be suppressed further.

Further, for example, when the radiation detector 600 is attached to a portion of the irradiation surface in the housing and the substrate 602 is formed of a highly rigid plastic resin, aramid, or bionanofiber, the rigidity of the radiation detector 600 itself may be increased. Therefore, the irradiation surface portion of the housing can be formed thin. In addition, when the substrate 602 is formed of a highly rigid plastic resin, aramid, or bionanofiber, the radiation detector 600 itself has flexibility, so that even when an impact is applied to the irradiated surface, the radiation detector 600 is not easily damaged. .

The radiation detector 600 described above may be configured as follows.

(1) The photoelectric conversion film 616 may be formed of an organic photoelectric conversion material, and the TFT layer 638 using a CMOS sensor may be formed. In this case, since only the photoelectric conversion film 616 is made of an organic material, the TFT layer 638 including the CMOS sensor may not have flexibility.

(2) The photoelectric conversion film 616 may be formed of an organic photoelectric conversion material, and the flexible TFT layer 638 may be realized by a CMOS circuit including a TFT 624 made of an organic material. In this case, pentacene may be adopted as the material of the p-type organic semiconductor used in the CMOS circuit, and copper fluoride phthalocyanine (F ₁₆ CuPc) may be adopted as the material of the n-type organic semiconductor. Thus, a flexible TFT layer 638 that can have a smaller bending radius can be realized. In addition, by configuring the TFT layer 638 in this way, the gate insulating film can be significantly reduced, and the driving voltage can be lowered. Furthermore, the gate insulating film, the semiconductor, and each electrode can be manufactured at room temperature or 100 ° C. or lower. Furthermore, a CMOS circuit can be directly formed over the flexible substrate 602. Moreover, the TFT 624 made of an organic material can be miniaturized by a manufacturing process in accordance with a scaling law. Note that when the polyimide precursor is applied to a thin polyimide substrate by a spin coat method and heated, the polyimide precursor is changed to polyimide, so that a flat substrate without unevenness can be realized.

(3) Applying a self-alignment placement technique (Fluidic Self-Assembly method) that places a plurality of micron-order device blocks at specified positions on a substrate 602, a photoelectric conversion film 616 and a TFT 624 made of crystalline Si are formed on a resin substrate You may arrange | position on the board | substrate 602 which consists of. In this case, the photoelectric conversion film 616 and TFT 624 as micro device blocks of micron order are fabricated in advance on another substrate and then separated from the substrate, and the photoelectric conversion film 616 and TFT 624 in the liquid are placed on the substrate 602 as the target substrate. Sprinkle on and place statistically. The substrate 602 is processed in advance to be adapted to the device block, and the device block can be selectively placed on the substrate 602. Therefore, an optimal device block (photoelectric conversion film 616 and TFT 624) made of an optimal material can be integrated on an optimal substrate (semiconductor substrate, quartz substrate, glass substrate, etc.), and is not a crystal. It is also possible to integrate device blocks (photoelectric conversion film 616 and TFT 624) optimum for a substrate (flexible substrate such as plastic).

In the radiation detector 600 according to the above-described modification, the light emitted from the scintillator 608 is converted into charges by the sensor unit 606 (photoelectric conversion film 616) located on the side opposite to the side where the radiation source 34 is located. Although it is configured as a so-called back side scanning method (PSS (Penetration Side Sampling) method) for reading an image, it is not limited to this configuration.

For example, the radiation detector may be configured as a so-called surface reading method (ISS (Irradiation Side Sampling) method). In this case, the substrate 602, the signal output unit 604, the sensor unit 606, and the scintillator 608 are laminated in this order along the irradiation direction of the radiation 26, and the light emitted from the scintillator 608 is sensor unit on the side where the radiation source 34 is located. At 606, the radiation image is read after being converted into electric charges. In general, the scintillator 608 emits light more strongly on the irradiation surface side of the radiation 26 than on the back surface side. Therefore, in the radiation detector configured by the front surface reading method, the scintillator is compared with the radiation detector configured by the back surface reading method. The distance until the light emitted in 608 reaches the photoelectric conversion film 616 can be shortened. Thereby, since the diffusion / attenuation of the light can be suppressed, the resolution of the radiation image can be increased.

Claims

A radiation device (28) having a radiation source (34), an irradiation dose control unit (104) for controlling an irradiation dose irradiated from the radiation source (34), and the radiation source (34) transmitted through the subject (24) A radiation detection device (30) for converting radiation (26) from
A system control unit (14) for controlling the radiographic imaging device (12) to perform radiographic imaging at a set frame rate;
The irradiation dose control unit (104)
When the gray value of the radiation image information is lower than a reference value, the next radiation dose is increased, and when the gray value is higher than the reference value, the next radiation dose is controlled to be decreased. ,
The system control unit (14)
When the gray value is different from an expected result, the irradiation dose control unit (104) is controlled to perform at least one radiography without increasing the next radiation irradiation dose. A radiographic imaging system comprising an error processing unit (116) that is controlled to perform.
In the radiographic imaging system of Claim 1,
The system control unit (14)
A density difference acquisition unit (112) that acquires a difference in density values of radiation image information based on at least two times of radiography;
And an error notification unit (114) for notifying an error that the gray value is different from an expected result when a difference in gray value changes by more than a predetermined value. Radiation imaging system.
The radiographic imaging system according to claim 2,
The error processing unit (116)
When the error notification is performed, the irradiation dose control unit (104) is controlled so as to decrease at least one radiation imaging without increasing the next radiation irradiation dose. A radiographic imaging system characterized by that.
In the radiographic imaging system of Claim 2 or 3,
The error notification unit (114)
Radiation image capturing, wherein the error notification is performed when the gray value of the radiation image information based on the current radiography is lower than the specified value by more than the gray value of the radiation image information based on the previous radiography system.
The radiographic imaging system according to any one of claims 2 to 4,
The irradiation dose control unit (104)
The upper limit signal value Vmax corresponding to the upper limit value of the irradiation dose and the lower limit signal value Vmin corresponding to the lower limit value of the irradiation dose are set as the dynamic range, the reference signal value corresponding to the reference value is set as Vo, and the gray value and the reference value are set. The radiation dose of the radiation (26) from the radiation source (34) is controlled by outputting to the radiation device (28) a control signal in which the signal value for controlling the radiation dose based on the comparison with V is V. A radiographic imaging system characterized by:
In the radiographic imaging system of Claim 5,
When the error notification is performed, the irradiation dose control unit (104) sets the signal value V to be equal to or less than the reference signal value Vo based on an instruction from the error processing unit (116). A radiographic imaging system characterized by
In the radiographic imaging system of Claim 5,
When the error notification is performed, the irradiation dose control unit (104), based on an instruction from the error processing unit (116), the signal value V is less than the reference signal value Vo, And the radiographic imaging system characterized by setting to the value according to the difference of the said light and shade value.
In the radiographic imaging system of Claim 7,
The irradiation dose control unit (104)
When the difference between the gray values is not less than the specified value and less than the second specified value, the signal value V is set to the lower limit signal value Vmin × constant Ka (0 <Ka <1.0),
The radiographic imaging system, wherein the signal value V is set to the lower limit signal value Vmin when the difference between the gray values is equal to or less than the second specified value.
The radiographic imaging system according to any one of claims 1 to 8,
The radiographic imaging system according to claim 1, wherein the one radiographic imaging is performed at the set frame rate.
The radiographic imaging system according to any one of claims 1 to 9,
The error processing unit (116)
A radiographic imaging system, wherein after performing the radiographic imaging once, the radiographic imaging system performs control so that the radiographic imaging is performed by gradually increasing the radiation dose every time radiographic imaging is performed at the frame rate.
A radiation device (28) having a radiation source (34), an irradiation dose control unit (104) for controlling an irradiation dose irradiated from the radiation source (34), and the radiation source (34) transmitted through the subject (24) Radiation imaging method for performing radiation imaging at a set frame rate using a radiation image capturing device (12) having a radiation detection device (30) that converts radiation (26) from In
The irradiation dose control unit (104) increases the irradiation dose of the next radiation when the grayscale value of the radiation image information is lower than a reference value, and the next radiation when the grayscale value is higher than the reference value. Control to reduce the irradiation dose of
When the gray value is different from an expected result, the irradiation dose control unit (104) is controlled to perform at least one radiography without increasing the next radiation irradiation dose. The radiographic imaging method characterized by having the step controlled to perform.
A radiation device (28) having a radiation source (34), an irradiation dose control unit (104) for controlling an irradiation dose irradiated from the radiation source (34), and the radiation source (34) transmitted through the subject (24) Radiation imaging system that performs radiation imaging at a set frame rate using a radiation image capturing device (12) having a radiation detection device (30) that converts radiation (26) from In the error handling method of
The irradiation dose control unit (104) increases the irradiation dose of the next radiation when the grayscale value of the radiation image information is lower than a reference value, and the next radiation when the grayscale value is higher than the reference value. Control to reduce the irradiation dose of
When the gray value is different from an expected result, the irradiation dose control unit (104) is controlled to perform at least one radiography without increasing the next radiation irradiation dose. An error processing method of a radiographic imaging system, characterized by performing control so as to be performed.