JP2007050052A - Radiation imaging apparatus and its control method - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation imaging apparatus capable of correcting the change of an image density caused by a sensor in a moving image photographing, thereby outputting an excellent moving image. <P>SOLUTION: With respect to image data obtained when imaging the radiation images of continuous n frames for a subject 507, at least two or more image regions of different image densities are selected (step S101), an average output value in the respective selected image regions of the respective image data is calculated (step S102), the change amount of the average output value of the respective image regions in the image data of the second to n-th frames is calculated with the average output value of the respective image regions in the image data of the first frame as a reference (step S103), and the image density in the other image data is corrected on the basis of the calculated result. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

  The present invention relates to a radiation imaging apparatus suitable for use in medical diagnosis and industrial nondestructive inspection, a control method thereof, and a program for causing a computer to execute the control method.

  Conventionally, as an X-ray imaging system installed in a hospital or the like, a film imaging system that irradiates a patient with X-rays and exposes X-rays transmitted through the patient onto a film, and an X-ray transmitted through the patient as an electrical signal. There is a digital photographing system in which the electric signal is converted into a digital value using an AD converter and is taken into a memory.

  At present, the mainstream of the latter digital imaging system is, for example, as shown in Patent Document 1 below, in which an X-ray image is formed on a stimulable phosphor called an imaging plate (IP) using BaBr: Eu as a representative material. Once the image is stored. Thereafter, by scanning the IP with a laser beam, the visible light from the IP is converted into an electric signal by a photomultiplier tube or the like and digitized.

In Patent Document 2 below, a wavelength converter such as a phosphor that converts X-rays into visible light, and visible light that is emitted in proportion to the X-ray dose from the wavelength converter irradiated with X-rays are converted into electrical signals. A method of converting an X-ray into an electric signal by a conversion element constituted by an optical sensor made of amorphous silicon or the like to be converted is shown. Typical examples of the wavelength converter include Gd ₂ O ₂ S: Tb and CsI: Tl. This apparatus is called an FPD (Flat Panel Detector). Further, some FPDs are configured by a conversion element using Se, PbI _2, or the like that directly absorbs X-rays and converts them into electric signals without using a wavelength converter.

  In general, a method using IP requires a complicated process in which an image formed into a latent image with X-rays is once scanned with a laser, and thus is not suitable for moving image shooting like a film.

  In addition, X-rays are irradiated to the primary phosphor, photoelectrons from the phosphor screen are accelerated and focused by an electron lens, and the fluorescence image (X-ray image) on the secondary phosphor screen is converted into an electrical signal by an imaging tube or CCD. There is a device to do. This is a method suitable for moving image photographing called image intensifier (II), and is a general method well known as fluoroscopic photographing of the stomach. This I.I. I. In addition, an electric signal can be detected as a digital value, which is one of digital photographing methods.

  As described above, there are a wide variety of apparatuses for digitizing X-ray images, and in recent years, the demand has been increasing. If the image data can be digitized, there is an advantage that photographing data can be easily recorded, displayed, printed, and stored. Therefore, there is an increasing demand for digitization in the medical field.

  In the recent medical field where the film imaging method, that is, the analog imaging method is changing to the digital imaging method described above, simple X-ray imaging is performed as the first step of X-ray imaging. For example, in the case of the chest, this is called chest simple X-ray imaging, and X-ray imaging of the front (or side) of the chest of a human body is performed. In order to cover the entire chest (upper body) of the human body, it is generally said that a photographing area is required to be a half-cut size (35 cm × 43 cm) or more, preferably 43 cm × 43 cm or more. In chest X-ray photography, distortion of surrounding images is regarded as a problem. I. FPD is a more promising digital imaging method in the future.

  Patent Document 3 below discloses an example of an FPD capable of capturing a moving image using a wavelength converter and an MIS photoelectric conversion element made of an amorphous silicon thin film semiconductor as a conversion element. If amorphous silicon is used, a radiation imaging apparatus having a large area, good yield, and low cost can be provided.

  Further, when a shadow is confirmed as a result of a doctor interpreting a simple captured image, CT imaging is generally performed as the second step of X-ray imaging. In CT imaging, a tomographic image is obtained for a part confirmed by simple imaging. In CT imaging, a large number of images are acquired while changing the incident angle of the X-rays irradiated to the patient, and then a reconstruction process is performed by a computer or the like to obtain a tomographic image of the patient. This CT imaging may also be considered as video imaging in the sense that images are taken continuously in time series.

  In Patent Document 4 below, a large-area X-ray detection element in which conversion elements for X-ray detection are arranged in a two-dimensional array is used, and a cone beam X-ray is irradiated to a patient and a CT is performed while performing a helical scan. A method for obtaining an image has also been proposed.

JP-A-5-224322 JP-A-8-116044 JP 2003-78124 A Japanese Patent Laid-Open No. 4-343836

  I. I. Is suitable for moving image shooting and is widely used for gastroscopy, etc., but there is a problem in that the image distortion at the periphery due to the influence of the electronic lens is large and an accurate image cannot be obtained at the same magnification. An FPD using an amorphous semiconductor material such as hydrogenated amorphous silicon as a sensor can solve this problem and obtain an image having no one-to-one distortion.

  However, in the FPD using this amorphous semiconductor material as a sensor, the defect level density increases due to the impurities mixed into the film during the film formation process and the disorder of bonding of the particles in the film (unbonded hands). There is a problem that the output signal fluctuates when shooting as a moving image by causing a change in dark current and a change in sensitivity in the conversion process. This is a phenomenon in which the image density changes with time, particularly when an image is observed as a moving image, and is not preferable as an image. Even in the case of CT imaging, if the output signal changes, that is, if the image density changes with time, there arises a problem that a correct tomographic image cannot be obtained.

  The present invention has been made in view of the above-described problems, and at the time of moving image shooting, a radiation imaging apparatus capable of correcting a change in image density caused by a sensor and capable of outputting a good moving image image, An object is to provide a control method and a program therefor.

  The radiation imaging apparatus of the present invention is a radiation detection circuit unit in which conversion elements that convert radiation including radiation transmitted from a radiation source and transmitted through the subject into electrical signals are arranged in a two-dimensional array. When the radiographic image of continuous n frames (n is a natural number of 2 or more) is taken for the subject, the electrical signal detected by the radiation detection circuit unit is stored as continuous n image data. An image area selection means for selecting at least two or more image areas having different image densities for each of the image data, and an average output value calculation for calculating an average output value of each of the image areas in each of the image data And a change amount of the average output value in other image data on the basis of the average output value of each image area in any one piece of image data A change amount calculating unit that, based on the calculation result of the variation calculation means, is characterized in that it has a correction means for correcting the image density in the other image data.

  In the method for controlling a radiation imaging apparatus according to the present invention, radiation detection is performed in which conversion elements that convert radiation including radiation transmitted from a radiation source and transmitted through the subject into electrical signals are arranged in a two-dimensional array. In the circuit unit, when the radiographic image of the continuous n frames (n is a natural number of 2 or more) is taken with respect to the subject, the n images in which the electrical signals detected by the radiation detection circuit unit are continuously acquired. A method for controlling a radiation imaging apparatus including a memory for storing data, wherein for each of the image data, an image region selection step of selecting at least two or more image regions having different image densities, and each of the image data An average output value calculating step for calculating an average output value of the image area, and an average output value of each image area in any one piece of image data as a reference A change amount calculating step for calculating the change amount of the average output value in the other image data, and a correction step for correcting the image density in the other image data based on the calculation result of the change amount calculating step. It is characterized by having.

  The program of the present invention is a radiation detection circuit unit in which conversion elements that convert radiation including radiation transmitted from a radiation source and transmitted through the subject into electrical signals are arranged in a two-dimensional array. A memory that stores electrical signals detected by the radiation detection circuit unit as continuous n image data when radiographic images of continuous n frames (n is a natural number of 2 or more) are captured on a subject. A program for causing a computer to execute processing in a radiation imaging apparatus comprising: processing for selecting at least two or more image regions having different image densities for each of the image data; and each of the image data in the image data The process of calculating the average output value of the image area and the average output value of each image area in any one piece of image data as a reference A process for causing a computer to execute a process of calculating a change amount of the average output value in other image data and a process of correcting an image density in the other image data based on the calculated change amount. is there.

  According to the present invention, at least two or more image regions having different image densities are selected for each image data obtained when a continuous n-frame radiation image is captured on a subject, and each selected image data is selected. Calculate the average output value in each image area of the image data, and calculate the amount of change in the average output value of each image area in the other image data based on the average output value of each image area in any one piece of image data Since the image density in the other image data is corrected based on the calculation result, a change in image density caused by the sensor can be corrected during moving image shooting, and a good moving image can be obtained. Can be output. This makes it possible to provide a radiation imaging apparatus with high diagnostic efficiency and promote the realization of a higher quality medical environment for the future.

  Preferred embodiments of the present invention will be described below with reference to the drawings. In the embodiment of the present invention, an example using X-rays as radiation is shown. However, the radiation of the present invention is not limited to X-rays, and electromagnetic waves such as α-rays, β-rays, γ-rays, etc. Contains.

(First embodiment)
FIG. 1 is a schematic configuration diagram of an X-ray imaging apparatus according to the first embodiment of the present invention.
X-rays emitted in a pulse form from an X-ray tube 501 serving as an X-ray source are applied to a subject 507. Here, the subject 507 is mainly a human (patient). The X-rays transmitted through the subject 507 are converted from X-rays to visible light by a wavelength converter 502 such as a phosphor, and the visible light from the wavelength converter 502 is converted into an electric signal by the photoelectric conversion element 508. As a result, an X-ray fluoroscopic image of the subject 507 (patient) is obtained as an electrical signal.

As a material of the photoelectric conversion element 508, for example, a non-single crystal semiconductor such as amorphous silicon is used and formed as a pixel on the glass substrate 506. The wavelength converter 502 and the photoelectric conversion element 508 have a structure that is substantially adhered by bonding or the like, and an X-ray detection circuit unit 503 is configured in which conversion elements including both are arranged in a plurality of matrices. . The wavelength converter 502 is made of, for example, a material containing at least one of Gd ₂ O ₂ S, Gd ₂ O _3, and CsI as a main component. The X-ray power source 504 is a power source that supplies a voltage to the X-ray tube 501 and supplies a high voltage for accelerating electrons in the X-ray tube 501.

  In this embodiment, X-rays incident on the wavelength converter 502 are converted into visible light. However, the X-rays incident on the photoelectric conversion element 508 are absorbed without using the wavelength converter 502, and The X-ray may be directly converted into an electric signal. In this case, the material of the photoelectric conversion element 508 includes, for example, at least one of lead iodide, mercury iodide, selenium, cadmium telluride, gallium arsenide, gallium phosphorus, zinc sulfide, and silicon as a main component. Things. Thus, even if the conversion element in this invention is comprised by the wavelength converter 502 and the photoelectric conversion element 508, it comprises the conversion element by the photoelectric conversion element 508 using the material which converts an X-ray directly into an electrical signal. But you can.

  The memory 505 continuously captures the image signal detected by the X-ray detection circuit unit 503 when the subject 507 is captured with an X-ray transmission image of consecutive n frames (n is a natural number of 2 or more). A memory for storing digital image data, and has an area for storing each image data. Image data stored in the memory 505 is appropriately subjected to image processing in the image processing unit 520. The image processing unit 520 is special hardware such as a general-purpose personal computer or microcomputer, or a dedicated image processor, and can process a large amount of image data stored in the memory 505 at high speed. . The display unit 530 is a display unit for displaying an image based on the image data processed by the image processing unit 520, and is a CRT, a liquid crystal monitor, or the like.

  The image processing unit 520 includes, for each image data stored in the memory 505, an image area selection unit 521 that selects at least two image areas having different image densities, and an average output value of each image area in each image data. And an average output value calculating means 522 for calculating. Furthermore, it has a change amount calculation means 523 for calculating a change amount of the average output value of each image area in other image data with reference to the average output value of each image area in any one piece of image data. Furthermore, based on the calculation result by the change amount calculation unit 523, a correction unit 524 that corrects image density in other image data, and an image based on the image data that has undergone image density correction processing are displayed on the display unit 530. The display control means 525 for performing the control at the time is provided.

  In the X-ray imaging apparatus of the present embodiment, image data of an X-ray transmission image of the subject 507 stored in the memory 505 is subjected to image processing by the image processing unit 520, so that original image data (processed image data) is obtained. One object is to correct the amount of change in image density included. For this description, it is assumed that the memory 505 can store image data for n frames and stores image data for n frames.

FIG. 2 shows processing of the X-ray imaging apparatus according to the first embodiment of the present invention, and an example in which five image areas are selected for each of n pieces of image data stored in the memory 505. It is the shown schematic.
Here, in the present invention, at least two image areas are selected by the image area selection unit 521. In the example of FIG. 2, the image area A, the image area B, the image area C, the image area D, and the image area E are selected. Five image areas are selected. These five image areas are selected for the image data of all n frames, but the shape, area, and coordinates of the selected image areas do not change between the frames. These five image areas are not selected at random, but follow a certain selection guideline.

  As one selection guideline in selection of each image area by the image area selection means 521, it is desirable that the image density is not the same between the image areas, but there is a slight density difference. This is because the accuracy of correction data described later is improved. In addition, as another selection guideline in selecting each image area by the image area selection unit 521, an area where the subject 507 does not move is desirable. This is because an object of the present invention is to correct an output change due to the characteristics of the photoelectric conversion element 508. If an output change due to an image region in which the subject 507 moves (a change in the subject 507) is used, the present invention is used. This is because there is a possibility of inducing a correction error for the correction of the density change. However, there is no problem if the amount of movement of the subject 507 is small.

  Further, as another selection guideline in selecting each image area by the image area selection unit 521, an image area as wide as possible with respect to the imaging area of the X-ray imaging apparatus is selected. This is because the variation in the output fluctuation (density fluctuation) in the surface is corrected. Further, as another selection guideline in selecting each image area by the image area selection unit 521, it is better that the same density (the same output) is as much as possible in one image area. This is because there is a possibility that the correction accuracy may be lowered if a mixture of extremely different densities in one image area.

  In FIG. 2, the shape of the image area A to the image area E is a square shape. However, the shape is not necessarily a square shape, and may be, for example, a round shape or a triangular shape. The image area selected by the image area selection unit 521 may be one pixel, but a plurality of pixels are desirable in order to improve correction accuracy. However, it should be noted that if the image area becomes too large, one image area may have various densities. The selection area may be selected at an appropriate time while the observer sees the image. However, since it takes time and is not efficient, in the present embodiment, the image processing unit 520 performs the selection based on the selection guideline described above. The image area selection means 521 automatically selects it.

  FIG. 3 is a characteristic diagram showing an average output value of each image area in the image data of each frame. FIG. 3 shows the average output value of each image area shown in FIG. 2 calculated for each frame. The horizontal axis represents the frame number and the vertical axis represents the average output value of each image area. .

  In FIG. 3, the change of the average output value of each image area can be seen. That is, when displayed as a moving image, the image density changes with time. In the example of FIG. 3, the average output value of each image area is image area C> image area E> image area B> image area D> image area A. Each image region selected by the image region selection means 521 is selected from a wide range of image regions based on the selection guide described above so that the density between the image regions is different. The value of n indicating the number of frames depends on the use of the X-ray imaging apparatus. For example, in the case of performing fluoroscopic imaging of the stomach, imaging is performed for 10 seconds at a speed of 30 frames per second (30 fps). , N = 300.

  FIG. 4 is a characteristic diagram showing the amount of change in the average output value of each image region in the image data of other frames with reference to the average output value of each image region in the image data of the first frame. In FIG. 4, the horizontal axis is the average output value of each image region, the output average value of the first frame is plotted on the X axis, and the vertical axis is the average output value of each image region in the image data of the first frame. As a reference, the amount of change in the average output value of each image region in the image data of other frames (second to nth frames) is used.

  This amount of change means the correction amount of the image density in the image data of other frames (second to nth frames). That is, the pixel having the same output value as that of the image area A in the second frame is corrected to the same output value as that of the image area A of the first frame. Also, pixels having the same output value as the image area B in the second frame are corrected to the same output value as the image area B in the first frame. The same applies to the other image areas C to E. The same applies to the image areas of the third frame, the fourth frame,..., The nth frame. However, the correction amount of the pixel having an intermediate output value between the image regions may be calculated by approximation based on the change amounts of the image regions A to E. The approximation at this time may be obtained by performing function approximation such as the least square method. In FIG. 4, an arrow (↓) shown in the drawing indicates a correction amount of a pixel having an intermediate output value between the average output value of the image region E and the average output value of the image region C in the image data of the fifth frame. This is an example.

  In FIG. 4, the change amount (correction amount) is calculated based on the average output value of each image area in the image data of the first frame, but the reference frame is not limited to the first frame. Any one frame may be sufficient. That is, in the embodiment of the present invention, the image density of another frame is corrected with respect to the frame selected as the reference.

  FIG. 5 is a flowchart showing a control process of the X-ray imaging apparatus according to the first embodiment of the present invention. The series of processing shown in FIG. 5 is performed by the image processing unit 520 in FIG.

  First, in step S101, at least two or more image regions having different image densities for each of the image data of the first to nth frames stored in the memory 505 in the image region selection unit 521 based on the selection guide described above. Select.

  Subsequently, in step S102, the average output value calculation unit 522 calculates an average output value in each image region of each image data selected in step S101.

  Subsequently, in step S103, the change amount calculation unit 523 selects each of the image data of the second to nth frames based on the average output value in each of the selected image regions of the image data of the first frame. A change amount of the average output value in the image area is calculated.

  Subsequently, in step S104, the correction unit 524 generates correction data for each pixel value of each image data in the second to nth frames based on the calculation result in step S103.

  Subsequently, in step S105, the correction means 524 corrects the image density in each image data of the second to nth frames based on the generated correction data. Thereafter, an image based on the image data of the first to nth frames subjected to the image density correction processing is displayed on the display unit 530 under the control of the display control unit 525.

FIG. 6 is a two-dimensional circuit diagram of the X-ray detection circuit unit 503 of the X-ray imaging apparatus according to the first embodiment of the present invention.
The X-ray detection circuit unit 503 includes at least a photoelectric conversion circuit unit 701 and a reading circuit unit 707. In the photoelectric conversion circuit portion 701 in FIG. 6, 3 × 3 = 9 pixels are described for simplification of description.

In the photoelectric conversion circuit unit 701, S1-1 to S3-3 are MIS type photoelectric conversion elements, T1-1 to T3-3 are switch elements (TFTs), and G1 to G3 are gate wirings for turning on / off the TFTs. , M1 to M3 are signal wirings, and the Vs line is a wiring for applying an accumulation bias to the photoelectric conversion elements S1-1 to S3-3. The electrodes on the sides of the photoelectric conversion elements S1-1 to S3-3 that are painted black are G electrodes, and the opposite sides are D electrodes. Although the D electrode is shared with a part of the Vs line, an impurity semiconductor layer such as a thin N ⁺ layer is used as the D electrode for convenience of light incidence. The Vs line is biased by the power supply Vs. SR1 is a first shift register that applies a driving pulse voltage to the gate wirings G1 to G3. The voltage Vg (on) that turns on the switch elements (T1-1 to T3-3) and the switch element (T1-1). The voltage Vg (off) for turning off (˜T3-3) is supplied from the outside to the first shift register SR1.

  The reading circuit unit 707 reads the parallel output signals from the photoelectric conversion circuit unit 701, converts them in series, and outputs them. A1 to A3 are operational amplifiers in which signal wirings M1 to M3 and an inverting terminal (−) are respectively connected, and capacitive elements Cf1 to Cf3 are connected between the inverting terminal (−) and the output terminal, respectively. . When the switching elements (T1-1 to T3-3) are turned on, the capacitive elements Cf1 to Cf3 integrate currents flowing from the photoelectric conversion elements S1-1 to S3-3 to the capacitive elements and convert them into voltage amounts. . RES1 to RES3 are switches that reset the capacitive elements Cf1 to Cf3 to a reset bias V (reset), and are connected in parallel with the capacitive elements Cf1 to Cf3. In FIG. 6, the reset bias V (reset) is represented by 0 V, that is, GND.

  CL1 to CL3 are sample and hold capacitors for temporarily storing signals accumulated by the operational amplifiers A1 to A3 and the capacitive elements Cf1 to Cf3, Sn1 to Sn3 are switches for sample and hold, B1 to B3 are buffer amplifiers, Sr1 to Sr3 Is a switch for serially converting parallel signals, SR2 is a second shift register that provides pulses for serial conversion to the switches Sr1 to Sr3, and Ab is a buffer amplifier that outputs the serially converted signal. SW-res is a switch for resetting the non-inverting terminals of the operational amplifiers A1 to A3 to the reset bias V (reset) (reset to 0 V in FIG. 4), and SW-ref is for refreshing the non-inverting terminals of the operational amplifiers A1 to A3. These are switches for refreshing to a bias V (refresh), and they are controlled by a “REFRESH” signal. Specifically, the switch SW-ref is turned on when the “REFRESH” signal is “Hi”, while the switch SW-res is turned on when the signal is “Lo”, and the switches are not turned on at the same time. It has become.

FIG. 7 is a time chart showing the operation of the X-ray detection circuit unit shown in FIG.
FIG. 7 shows the operation of the X-ray detection circuit unit 503 for two frames. The operation of the X-ray detection circuit unit shown in FIG. 6 will be described using the time chart of FIG.

First, the photoelectric conversion period will be described.
In this photoelectric conversion period, the D electrodes of all the photoelectric conversion elements S1-1 to S3-3 are in a state of being biased to the reading power source Vs (positive potential). At this time, all the signals of the first shift register SR1 are “Lo”, and all the switch elements (T1-1 to T3-3) are turned off. When an X-ray pulse is emitted from the X-ray tube 501 in this state, visible light is irradiated to the D electrode (N ⁺ electrode) of each photoelectric conversion element through the wavelength converter 502, and the i layer of the photoelectric conversion element Inside, electron and hole carriers are generated. The electrons generated at this time move to the D electrode by the power source Vs, but holes are stored at the interface between the i layer and the insulating layer in the photoelectric conversion elements S1-1 to S3-3, and the X-ray from the X-ray tube 501 is stored. Retained after line is off.

Next, the reading period will be described.
The operation in this readout period is the photoelectric conversion elements S1-1 to S1-3 in the first row, the photoelectric conversion elements S2-1 to S2-3 in the second row, and the photoelectric conversion elements S3-1 to S3 in the third row. -3.

  First, in order to read out the charges (image signals) of the photoelectric conversion elements S1-1 to S1-3 in the first row, the first wiring is connected to the gate wiring G1 of the switch elements (T1-1 to T1-3) in the first row. A gate pulse is applied from the shift register SR1. At this time, the high level of the gate pulse is a voltage of Vg (on) supplied from the outside. As a result, the switch elements (T1-1 to T1-3) in the first row are turned on, and the charges accumulated in the photoelectric conversion elements S1-1 to S1-3 in the first row are changed to the first row. The current flows through the switch elements (T1-1 to T1-3), and flows into the capacitive elements Cf1 to Cf3 connected to the operational amplifiers A1 to A3 to be integrated.

  Although not particularly shown in FIG. 6, a read capacitor is added to the signal wirings M1 to M3, and the charges of the photoelectric conversion elements S1-1 to S1-3 in the first row are switched to the switching elements ( The data is transferred to the read capacity side via T1-1 to T1-3). However, since the signal wirings M1 to M3 are virtually grounded by the reset bias (GND) of the non-inverting terminals (+) of the operational amplifiers A1 to A3, there is no potential variation due to the transfer operation and the signal wirings M1 to M3 are held in GND. is there. That is, the electric charges of the photoelectric conversion elements S1-1 to S1-3 in the first row are transferred to the capacitive elements Cf1 to Cf3.

  The output terminals of the operational amplifiers A1 to A3 change as shown in FIG. 7 according to the charge amount of the photoelectric conversion elements S1-1 to S1-3 in the first row. Since the switch elements (T1-1 to T1-3) in the first row are turned on simultaneously, the outputs of the operational amplifiers A1 to A3 change simultaneously. That is, they are output in parallel. In this state, by turning on the “SMPL” signal, the output signals from the operational amplifiers A1 to A3 are transferred to the sample hold capacitors CL1 to CL3, and the SMPL signal is turned off and temporarily held.

  Next, by applying pulses from the second shift register SR2 in the order of the switches Sr1, Sr2, and Sr3, the charges held in the sample-and-hold capacitors CL1 to CL3 are changed to the amplifier Ab in the order of CL1, CL2, and CL3. Is output from. As a result, the charges (image signals) of the photoelectric conversion elements S1-1 to S1-3 in the first row are serially converted and output. The charge (image signal) read operation of the photoelectric conversion elements S2-1 to S2-3 in the second row is the same as the charge (image signal) read operation of the photoelectric conversion elements S3-1 to S3-3 in the third row. Done.

  The charges of the photoelectric conversion elements S1-1 to S1-3 in the first row can be obtained by sampling and holding the output signals from the operational amplifiers A1 to A3 in the sample hold capacitors CL1 to CL3 by the SMPL signal for the photoelectric conversion elements. This is output from the unit 701. Therefore, during the serial conversion and output by the switches Sr1 to Sr3 in the readout circuit unit 707, the refresh operation of the photoelectric conversion elements S1-1 to S1-3 in the first row in the photoelectric conversion circuit unit 701; Capacitance elements Cf1 to Cf3 can be reset.

  In the refresh operation of the photoelectric conversion elements S1-1 to S1-3 in the first row, the switch SW-ref is turned on by setting the “REFRESH” signal to “Hi”, and the switches RES1 to RES3 are turned on by the “RC” signal. This is achieved by applying a voltage Vg (on) to the gate wiring G1 of the TFTs (T1-1 to T1-3) in the first row. That is, the G electrodes of the photoelectric conversion elements S1-1 to S1-3 in the first row are refreshed to the refresh bias V (refresh) by the refresh operation. Thereafter, the photoelectric conversion element transitions from a refresh operation state to a reset operation for bringing the photoelectric conversion state into a state where photoelectric conversion is possible.

  The reset operation is performed in a state where the voltage Vg (on) is applied to the gate wiring G1 of the switch elements (T1-1 to T1-3) in the first row and the switches RES1 to RES3 are in the conductive state, and the “REFRESH” signal To “Lo”. By this operation, the G electrodes of the photoelectric conversion elements S1-1 to S1-3 in the first row are reset to the reset bias V (reset) = GND, and simultaneously reset the charges accumulated in the capacitive elements Cf1 to Cf3. . After the reset operation is completed, a gate pulse of the gate line G2 can be applied subsequently. That is, during the serial conversion operation by the second shift register SR2 with respect to the charges of the photoelectric conversion elements S1-1 to S1-3 in the first row, the photoelectric conversion elements S1-1 to S1- in the first row are simultaneously performed. 3 is reset, the capacitive elements Cf1 to Cf3 are reset, and the charges of the photoelectric conversion elements S2-1 to S2-3 in the second row are transferred to the signal wirings M1 to M3 by the first shift register SR1. It becomes possible.

  With the above operation, charges (image signals) of all the photoelectric conversion elements S1-1 to S3-3 from the first row to the third row can be output. A continuous image can be acquired by repeating the operation for one frame a plurality of times.

FIG. 8 is a block diagram of each component that processes an analog signal output from the readout circuit unit 707 of FIG. 6 in the X-ray imaging apparatus according to the first embodiment of the present invention.
In FIG. 8, an AD converter (ADC) 61, a CPU 62, a shift register 63, and memory units 64 (1) to 64 (n) are configured as the respective components. For example, in this embodiment, the AD converter (ADC) 61 is configured in the X-ray detection circuit unit 503, the memory units 64 (1) to 64 (n) are configured in the memory 505, and the CPU 62 and the shift register 63. Is configured in the image processing unit 520.

  The AD converter (ADC) 61 converts the analog signal output from the reading circuit unit 707 into a digital signal. The memory units 64 (1) to 64 (n) are for storing image signals from the first frame (F1) to the nth frame (Fn) as image data.

  The analog signal output from the reading circuit unit 707 is input to the AD converter (ADC) 61. The resolution of the AD converter (ADC) 61 varies depending on the purpose of diagnosis, but in the case of chest radiography, 12 to 14 bits or more is appropriate. The digital signal from the AD converter (ADC) 61 is stored as image data in the memory units 64 (1) to 64 (n) for each frame. In FIG. 8, n memory units are arranged, and image data corresponding to shooting from the first frame (F1) to the nth frame (Fn) is stored. Signals from the respective memory units are arithmetically processed in a CPU (Central Processing Unit) 62.

  FIG. 9 is a timing chart showing the operation of the CPU 62 shown in FIG. 8, and includes the X-ray generation timing for each frame (F1,..., Fn-1, Fn).

FIG. 10 is a top view illustrating a schematic configuration of the photoelectric conversion circuit unit 701 illustrated in FIG. 6.
Here, the MIS type photoelectric conversion element 101 corresponds to the photoelectric conversion elements S1-1 to S3-3 in FIG. 6, and the switch element 102 is a switch element (T1-1 to T3) which is a TFT in FIG. -3). In addition, the photoelectric conversion element 101 and the switch element 102 are configured using a non-single-crystal semiconductor thin film such as an amorphous silicon semiconductor thin film, and FIG. 11 is a schematic cross-sectional view taken along the line AB in the photoelectric conversion circuit portion 701 illustrated in FIG. In the following description, for the sake of simplicity, the MIS photoelectric conversion element will be described simply as a photoelectric conversion element.

  The photoelectric conversion element 101 and the switch element 102 (here, an amorphous silicon TFT, which will be simply referred to as a TFT hereinafter) are formed on the same insulating substrate 103. The lower electrode of the photoelectric conversion element 101 is shared by the same first metal thin film layer 104 as the lower electrode (gate electrode) of the TFT 102, and the upper electrode of the photoelectric conversion element 101 is the upper electrode (source electrode, The second metal thin film layer 105 that is the same as the drain electrode) is shared.

  The first metal thin film layer 104 and the second metal thin film layer 105 also share the gate drive wiring 106 and the matrix signal wiring 107 in the photoelectric conversion circuit portion 701 shown in FIG. In FIG. 10, a total of 4 pixels of 2 × 2 is described as the number of pixels. The hatched portion in FIG. 10 is a light receiving surface of the photoelectric conversion element 101. Reference numeral 109 denotes a power supply line for applying a bias to the photoelectric conversion element. Reference numeral 110 denotes a contact hole for connecting the photoelectric conversion element and the TFT.

  As shown in FIGS. 10 and 11, when a configuration using an amorphous silicon semiconductor as a main material is used, the photoelectric conversion element 101, the switch element 102, the gate driving wiring 106, and the matrix signal wiring 107 are placed on the same insulating substrate 103. The photoelectric conversion circuit portion 701 having a large area can be easily manufactured at a low cost.

Next, a single device operation of the photoelectric conversion element 101 will be described.
FIG. 12 is an energy band diagram for explaining the device operation of the photoelectric conversion element 101.

  12A and 12B show the operation in the refresh mode and the photoelectric conversion mode, respectively, and FIG. 12C shows the operation in the saturated state. Further, M1 to M2 shown in the horizontal direction of FIGS. 12A to 12C represent the state of the thickness direction of each layer shown in FIG.

Specifically, M1 is a lower electrode (G electrode) formed of the first metal thin film layer 104 (here, for example, Cr) in FIG. The amorphous silicon nitride (a-SiNx) layer corresponds to the a-SiN insulating thin film layer 111 in FIG. 11, and is a layer that blocks passage of both electrons and holes. The a-SiNx layer needs to have a thickness that does not cause a tunnel effect, and is usually set to 500 angstroms or more. The hydrogenated amorphous silicon (a-Si: H) layer corresponds to the a-Si semiconductor thin film layer 112 in FIG. 11, and is formed of an intrinsic semiconductor layer (i layer) that is not intentionally doped with a dopant. A photoelectric conversion semiconductor layer. The N ⁺ layer corresponds to the N ⁺ layer 119 in FIG. 11, and is a non-single layer such as an N-type a-Si: H layer formed to prevent hole injection into the a-Si: H layer. It is a single conductivity type carrier injection blocking layer made of a crystalline semiconductor. M2 is an upper electrode (D electrode) formed of the second metal thin film layer 105 (here, Al, for example) in FIG.

In Figure 11, since the second metal thin film layer 105 (D electrode) but does not completely cover the N ⁺ layer 113, between the D electrode and the N ⁺ layer 113 in which electrons move is freely performed, D The electrode and the N ⁺ layer 113 are always at the same potential, and this is assumed in the following description.

  The photoelectric conversion element 101 has two types of operation modes, a refresh mode and a photoelectric conversion mode, depending on how the voltage is applied to the D electrode and the G electrode.

In FIG. 12A showing the refresh mode, the D electrode is given a negative potential with respect to the G electrode, and holes indicated by black circles in the i layer are guided to the D electrode by an electric field. At the same time, electrons indicated by white circles are injected into the i layer. At this time, some holes and electrons recombine and disappear in the N ⁺ layer and the i layer. If this state continues for a sufficiently long time, the holes in the i layer are swept out of the i layer.

In order to change the refresh mode to the photoelectric conversion mode shown in FIG. 12B, a positive potential is applied to the D electrode with respect to the G electrode. Thereby, electrons in the i layer are instantaneously guided to the D electrode. However, since the N ⁺ layer functions as an injection blocking layer, holes are not introduced into the i layer. When light enters the i layer in this state, the light is absorbed and an electron / hole pair is generated. The electrons generated here are guided to the D electrode by the electric field, and the holes move in the i layer and reach the interface between the i layer and the a-SiNx layer, but cannot move in the a-SiNx layer. Will stay. At this time, electrons move to the D electrode, and holes move to the interface with the a-SiNx layer in the i layer, so that a current flows from the G electrode in order to maintain electrical neutrality in the photoelectric conversion element 101. This current is proportional to the incident light because it corresponds to the electron-hole pair generated by the light.

  After maintaining the state of the photoelectric conversion mode in FIG. 12B for a certain period and then entering the state of the refresh mode in FIG. 12A again, the holes remaining in the i layer are introduced to the D electrode as described above. At the same time, a current corresponding to this hole flows. The amount of holes corresponds to the total amount of light incident during the photoelectric conversion mode period. At this time, a current corresponding to the amount of electrons injected into the i layer also flows, but since this amount is approximately constant, it may be detected by subtracting. That is, the photoelectric conversion element 101 outputs the amount of light incident in real time, and at the same time can detect the total amount of light incident during a certain period.

  However, in the photoelectric conversion element 101, when the period of the photoelectric conversion mode is increased for some reason or when the illuminance of incident light is strong, current may not flow even though light is incident. This is because, as in the saturated state shown in FIG. 12 (c), a large number of holes remain in the i layer, and the electric field in the i layer is reduced by the holes, and the generated electrons are not guided. This is because they recombine with holes. If the incident state of light changes in this saturation state, the current may flow in an unstable manner. However, if the refresh mode of FIG. 12A is set again, the holes in the i layer are swept out to the next. In this photoelectric conversion mode, a current proportional to light flows again.

In the above description, when holes in the i layer are swept out in the refresh mode of FIG. 12A, it is ideal to sweep out all the holes. However, sweeping out some holes is also effective. There is no problem in obtaining the same current as described above. That is, at the detection opportunity in the next photoelectric conversion mode, it does not have to be in the saturation state of FIG. 12C, the potential of the D electrode with respect to the G electrode in the refresh mode, the period of the refresh mode, and the implantation of the N ⁺ layer The characteristics of the blocking layer can be determined. Furthermore, in the refresh mode of FIG. 12A, the injection of electrons into the i layer is not a necessary condition, and the potential of the D electrode with respect to the G electrode is not limited to negative. This is because when many holes remain in the i layer, even if the potential of the D electrode with respect to the G electrode is a positive potential, the electric field in the i layer is applied in the direction in which the holes are guided to the D electrode. Similarly, the characteristics of the N ⁺ layer that is the injection blocking layer are not necessarily required to be able to inject electrons into the i layer.

(Second Embodiment)
FIG. 13 shows the processing of the X-ray imaging apparatus according to the second embodiment of the present invention, and five image areas are selected for each of n chest image data stored in the memory 505. It is the schematic which showed the example.

  In the second embodiment, for example, when diagnosing by observing the respiratory dynamics of the subject (patient) 507, the image region selection unit 521 selects an image region other than the lung field where the subject 507 moves. Here, the image area B and the image area E are areas where the subject 507 does not exist (elementary areas). The shoulder of the image area C and the neck of the image area D are areas where there is relatively no movement in breathing. The mediastinal portion of the image area A may move slightly due to respiration, but the image area is selected because it moves less than the lung field. As an image area to be selected, there is an index that the subject (patient) 507 does not move (the movement is small), but it is not absolute. If the image area to be selected is appropriate, the correction accuracy increases as the number of selections increases. However, the time required for calculation by the image processing unit 520 is longer, which is case by case.

(Third embodiment)
FIG. 14 is a two-dimensional circuit diagram of the X-ray detection circuit unit of the X-ray imaging apparatus according to the third embodiment of the present invention.
The difference from FIG. 6 is that the photoelectric conversion elements S1-1 to S3-3 are not MIS type sensors but PIN type sensors. Unlike the MIS type sensor, this PIN type sensor can continuously shoot without performing a refresh operation, and therefore, it can generally have a higher frame rate than the MIS type sensor. Further, since the photoelectric conversion elements S1-1 to S3-3 are configured by PIN sensors, the readout circuit unit 702 has a configuration different from that of the readout circuit unit 707 in FIG.

(Fourth embodiment)
FIG. 15 is a schematic diagram showing an example in which the X-ray imaging apparatus is applied to an X-ray diagnostic system according to the fourth embodiment of the present invention.
X-rays 6060 generated by an X-ray tube 6050 serving as an X-ray source pass through a chest 5071 of a subject (patient) 507 and enter an image sensor 6040. X-rays incident on the image sensor 6040 include information inside the body of the subject 507. In the image sensor 6040, in response to the incidence of X-rays, the phosphor converts the light into visible light, which is further photoelectrically converted to obtain an electrical signal. This electric signal is digitally converted and subjected to image processing by an image processor 6070, and is displayed as an image on a display 6080 in a control room (control room) and observed.

  The X-ray tube 6050 of the present embodiment corresponds to, for example, the X-ray tube 501 in FIG. 1, and the image sensor 6040 corresponds to, for example, the X-ray detection circuit unit 503 in FIG. 1 corresponds to, for example, the X-ray power source 504, the memory 505, and the image processing unit 520 in FIG. 1, and the display 6080 corresponds to, for example, the display unit 530.

  In addition, image data generated by image processing by the image processor 6070 can be transferred to a remote place by transmission means such as a telephone line 6090 and displayed on a display 6081 in another place such as a doctor room or stored on an optical disk or the like. It can be stored in the means and can be diagnosed by a remote doctor. The image data can also be recorded as a film 6110 by the film processor 6100.

  According to the embodiment of the present invention, at least two or more image regions having different image densities are selected for each image data obtained when a continuous n-frame radiation image is captured on the subject 507. (Step S101), an average output value in each image area of each selected image data is calculated (Step S102), and the second to nth frames are based on the average output value of each image area in the image data of the first frame. The amount of change in the average output value of each image area in the image data of the image is calculated (step S103), and the image density in the other image data is corrected based on the calculation result (step S104 and step S104). S105) At the time of moving image shooting, the change in image density caused by the sensor can be corrected, and a good moving image can be output. It is possible. This makes it possible to provide a radiation imaging apparatus with high diagnostic efficiency and promote the realization of a higher quality medical environment for the future.

  Further, according to the embodiment of the present invention, since hydrogenated amorphous silicon is used as a sensor (photoelectric conversion element 508), it is possible to provide an FPD that is large in area and inexpensive. That is, I.I. I. Instead, the surrounding image is not distorted at all, and an accurate image with the same magnification can be obtained. In addition, application to an X-ray CT apparatus that continuously captures images at high speed becomes realistic.

  In the X-ray imaging apparatus of the present embodiment, both still images and moving images can be captured, and the imaging data obtained by constructing an X-ray diagnostic system having the X-ray imaging apparatus as shown in FIG. Can be recorded, displayed, printed, and stored easily, and a completely new X-ray diagnostic system can be provided that meets the recent demands for digitization in place of the conventional film imaging system.

  Each unit of FIG. 1 constituting the X-ray imaging apparatus according to the embodiment of the present invention described above and each step of FIG. 5 showing a control method of the X-ray imaging apparatus are stored in a RAM or a ROM of a computer. This can be realized by operating the program. This program and a computer-readable storage medium storing the program are included in the present invention.

  Specifically, the program is recorded in a storage medium such as a CD-ROM, or provided to a computer via various transmission media. As a storage medium for recording the program, a flexible disk, a hard disk, a magnetic tape, a magneto-optical disk, a nonvolatile memory card, and the like can be used in addition to the CD-ROM. On the other hand, as the transmission medium of the program, a communication medium (wired line such as an optical fiber, etc.) in a computer network (LAN, WAN such as the Internet, wireless communication network, etc.) system for propagating and supplying program information as a carrier wave A wireless line or the like.

  In addition, the functions of the X-ray imaging apparatus according to the embodiment of the present invention are realized by executing a program supplied by the computer, and an OS (operating system) or other program running on the computer. When the functions of the X-ray imaging apparatus according to the embodiment of the present invention are realized in cooperation with the application software or the like, all or part of the processing of the supplied program is a function expansion board or function expansion unit of a computer Such a program is also included in the present invention even when the functions of the X-ray imaging apparatus according to the embodiment of the present invention are realized.

1 is a schematic configuration diagram of an X-ray imaging apparatus according to a first embodiment of the present invention. FIG. 5 is a schematic diagram showing processing of the X-ray imaging apparatus according to the first embodiment of the present invention, and showing an example in which five image areas are selected for n pieces of image data stored in a memory. is there. It is the characteristic view which showed the average output value of each image area | region in the image data of each frame. FIG. 6 is a characteristic diagram showing the amount of change in the average output value of each image region in the image data of other frames, with the average output value of each image region in the image data of the first frame as a reference. It is the flowchart which showed the control processing of the X-ray imaging device which concerns on the 1st Embodiment of this invention. FIG. 2 is a two-dimensional circuit diagram of an X-ray detection circuit unit of the X-ray imaging apparatus according to the first embodiment of the present invention. 7 is a time chart showing the operation of the X-ray detection circuit unit shown in FIG. 6. FIG. 7 is a block diagram of each component that processes an analog signal output from the readout circuit unit in FIG. 6 in the X-ray imaging apparatus according to the first embodiment of the present invention. It is a timing chart which shows operation | movement of CPU shown in FIG. FIG. 7 is a top view illustrating a schematic configuration of the photoelectric conversion circuit unit illustrated in FIG. 6. FIG. 11 is a schematic cross-sectional view taken along line AB in the photoelectric conversion circuit unit illustrated in FIG. 10. It is an energy band figure for demonstrating the device operation | movement of a photoelectric conversion element. FIG. 7 is a diagram illustrating processing of the X-ray imaging apparatus according to the second embodiment of the present invention, showing an example in which five image regions are selected for n chest radiographed image data stored in a memory. FIG. It is a two-dimensional circuit diagram of the X-ray detection circuit part of the X-ray imaging device which concerns on the 3rd Embodiment of this invention. It is the schematic which showed the 4th Embodiment of this invention and showed the example which applied the X-ray imaging device to the X-ray diagnostic system.

Explanation of symbols

501 X-ray tube 502 Wavelength converter 503 X-ray detection circuit unit 504 X-ray power source 505 Memory 506 Glass substrate 507 Subject 508 Photoelectric conversion element 520 Image processing unit 521 Image region selection unit 522 Average output value calculation unit 523 Change amount calculation unit 524 Correction unit 525 Display control unit 530 Display unit 701 Photoelectric conversion circuit units 702 and 707 Read circuit units A1 to A3, B1 to B3, Ab Operational amplifier Cf1 to Cf3 Integration capacitance SW-res A reset bias is applied to the (+) terminal of the operational amplifier. Switch SW-ref switches G1 to G3 that apply a refresh bias to the (+) terminal of the operational amplifier gate drive wirings M1 to M3 matrix signal wirings RES1 to RES3 switches S1-1 to S3 that reset load capacitances formed in Cf1 to Cf3 -3 Photoelectric conversion Elements T1-1 to T3-3 Switch elements Sn1 to Sn3 Transfer switches Sr1 to Sr3 for transferring signals to the read capacitors Read switches SR1 for sequentially reading the signals of the read capacitors First shift register SR2 Second shift Register Vs Bias power supply V (reset) Reset power supply V (refresh) Refresh power supply Vg (on) Power supply Vg (off) for turning on TFT Power supply 61 for turning off TFT AD converter (ADC)
62 CPU
63 Shift register 64 (1) to 64 (n) Memory unit 101 Photoelectric conversion element 102 Switch element (TFT)
103 Insulating substrate 104 First metal thin film layer 105 Second metal thin film layer 106 Gate drive wiring 107 Matrix signal wiring 110 Contact hole portion 111 a-SiN insulating thin film layer 112 a-Si semiconductor thin film layer 113 N ⁺ layer 114 Wiring Cross part 115 Protective film

Claims (12)

In a radiation detection circuit unit in which a conversion element that converts radiation including radiation transmitted from a radiation source and transmitted through the subject into an electrical signal is arranged in a two-dimensional array, the radiation detection circuit unit is continuous with the subject. a memory that stores electrical signals detected by the radiation detection circuit unit as continuous n pieces of image data when capturing radiographic images of n frames (n is a natural number of 2 or more);
Image area selection means for selecting at least two or more image areas having different image densities for each of the image data;
Average output value calculating means for calculating an average output value of each of the image regions in each of the image data;
A change amount calculating means for calculating a change amount of the average output value in the other image data on the basis of an average output value of each of the image regions in any one piece of image data;
A radiation imaging apparatus comprising: a correcting unit that corrects an image density in the other image data based on a calculation result by the change amount calculating unit.   The correction unit generates correction data for each pixel value of the other image data based on a calculation result by the change amount calculation unit, and corrects an image density in the other image data based on the correction data. The radiation imaging apparatus according to claim 1, wherein:   The radiation imaging apparatus according to claim 1, wherein the image area selection unit selects an area where the subject does not move as the image area.   The radiation imaging apparatus according to claim 1, wherein one of the image areas is an area where the subject does not exist.   The said conversion element is comprised by the wavelength converter which converts a radiation into visible light, and the photoelectric conversion element which converts the said visible light into an electrical signal, The any one of Claims 1-4 characterized by the above-mentioned. The radiation imaging apparatus described in 1. The radiation imaging apparatus according to claim 5, wherein the wavelength converter includes at least one of Gd ₂ O ₂ S, Gd ₂ O _3, and CsI as a main component.   The radiation imaging apparatus according to claim 5, wherein the photoelectric conversion element is a PIN sensor using an amorphous silicon semiconductor.   The radiation imaging apparatus according to claim 5, wherein the photoelectric conversion element is a MIS type sensor using an amorphous silicon semiconductor. The MIS type sensor is
A lower electrode comprising a first metal thin film layer on a substrate;
An insulating layer formed on the lower electrode and made of amorphous silicon nitride that blocks passage of electrons and holes;
A photoelectric conversion layer made of hydrogenated amorphous silicon formed on the insulating layer;
An N-type injection blocking layer formed on the photoelectric conversion layer and blocking hole injection;
An upper electrode formed of at least a part of the injection blocking layer and made of a transparent conductive layer or a second metal thin film layer;
In refresh mode,
An electric field is applied to the MIS type sensor in a direction that leads holes from the photoelectric conversion layer to the upper electrode,
In photoelectric conversion mode,
An electric field is applied to the MIS type sensor in such a direction that electrons and holes generated by radiation incident on the photoelectric conversion layer remain in the photoelectric conversion layer and guide the electrons to the upper electrode. The radiation imaging apparatus according to claim 8, wherein a hole accumulated in an electron or an electron guided to the upper electrode is detected as an optical signal.   The conversion element absorbs radiation emitted from the radiation source and converts the radiation directly into an electrical signal. Lead iodide, mercury iodide, selenium, cadmium telluride, gallium arsenide, gallium phosphide The radiation imaging apparatus according to claim 1, wherein at least one of zinc sulfide and silicon is a main component. In a radiation detection circuit unit in which a conversion element that converts radiation including radiation transmitted from a radiation source and transmitted through the subject into an electrical signal is arranged in a two-dimensional array, the radiation detection circuit unit is continuous with the subject. Radiation imaging apparatus comprising a memory for storing electrical signals detected by the radiation detection circuit unit as continuous n pieces of image data when imaging of radiation images of n frames (n is a natural number of 2 or more) is performed. Control method,
An image area selecting step for selecting at least two or more image areas having different image densities for each of the image data;
An average output value calculating step for calculating an average output value of each of the image regions in each of the image data;
A change amount calculating step of calculating a change amount of the average output value in other image data with reference to an average output value of each of the image regions in any one piece of image data;
A control method for a radiation imaging apparatus, comprising: a correction step of correcting an image density in the other image data based on a calculation result in the change amount calculation step. In a radiation detection circuit unit in which a conversion element that converts radiation including radiation transmitted from a radiation source and transmitted through the subject into an electrical signal is arranged in a two-dimensional array, the radiation detection circuit unit is continuous with the subject. Radiation imaging apparatus provided with a memory for storing electrical signals detected by the radiation detection circuit unit as continuous n pieces of image data when imaging of radiation images of n frames (n is a natural number of 2 or more) is performed A program for causing a computer to execute the process in
For each of the image data, a process of selecting at least two or more image regions having different image densities;
A process of calculating an average output value of each of the image regions in each of the image data;
A process of calculating a change amount of the average output value in other image data on the basis of an average output value of each of the image regions in any one piece of image data;
A program for causing a computer to execute a process of correcting an image density in the other image data based on the calculated change amount.

Images