JP2011036393A - Radiation image photographing system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To speedily obtain a main photographic image free from unevenness in pixel signal values occurring between a position corresponding to a line read in pre-photographing and a position corresponding to a line which is not read in pre-photographing. <P>SOLUTION: In a radiation image capturing system 1, a radiation image detector 107, in the pre-photographing, reads an area including a subject H in a detection part 107a at predetermined line intervals to obtain a pre-photographic image and, in the main photographing, reads all the lines in the area including the subject H in the detection part 107a to obtain the main photographic image. A console 200 corrects the unevenness in pixel signal values occurring between the position corresponding to the line read in pre-photographing and the position corresponding to the line which is not read in pre-photographing, and displays thereof on a display part 213. <P>COPYRIGHT: (C)2011,JPO&INPIT

Description

  The present invention relates to a radiographic imaging system.

  2. Description of the Related Art Conventionally, in a radiographic imaging system, a radiographic image is obtained by irradiating a subject with a small amount of X-rays and transmitting through the subject before main imaging for imaging a diagnostic imaging image (referred to as a main imaging image). Pre-imaging that is read by a detector is performed, and irradiation conditions such as an X-ray irradiation dose at the time of actual imaging are determined based on X-ray dose information obtained by pre-imaging.

  Further, in Patent Document 1, reading is performed from a radiation image detector at a predetermined line interval in pre-imaging, and all lines of the radiation image detector are read in main imaging, and the value of each pixel read by pre-imaging. In addition, it is described that the value of each pixel read by the main imaging is added for each pixel to obtain an X-ray image.

  In pre-imaging, reading is performed by thinning out pixels from the radiation image detector at predetermined line intervals, thereby reducing the amount of data for determining imaging conditions such as the X-ray irradiation amount during actual imaging, and starting pre-imaging The time from the actual shooting to the end of the main shooting can be shortened.

Japanese Patent No. 3711160

  Incidentally, when thinning reading is performed at a predetermined line interval during pre-imaging as in Patent Document 1, the residual potential after pre-imaging in the radiation image detector is as shown in FIG. That is, Lread <L, where Lread is the residual potential of the pixel located in the line L1 that was read during the pre-photographing, and L is the residual potential of the pixel located in the line L2 that was not read during the pre-photographing. When the main shooting is performed in such a state that the residual potential is uneven, the pixel signal value (density value) at the position corresponding to the line L1 read in the pre-shooting in the main shooting image is read in the pre-shooting. As a result, the pixel signal value at the position corresponding to the line L2 that has not been reduced becomes smaller.

  In Patent Document 1, an image is obtained by adding the value of each pixel acquired from the radiation image detector by pre-imaging and the value of each pixel acquired after the main imaging for each pixel. Obtainable.

However, for example, when the amount of data is very large, such as a captured image obtained by performing phase contrast magnification imaging on a breast and detecting using a half-cut size FPD of 50 μm pixels, if the addition is performed for each pixel, the actual captured image is not captured. There is a problem that it takes a very long time to acquire.
It is conceivable to perform a so-called reset to remove the charge accumulated in the radiation image detector after the pre-imaging, but as shown in FIG. 11, the charge reading efficiency in the radiation image detector is not 100%. In particular, when the X-ray dose is a small dose, the readout efficiency is low. For this reason, as shown in FIG. 10B, a residual potential remains even after resetting, and an image without the above-described unevenness cannot be obtained even if reset is performed in the same manner as in normal imaging. In other words, it does not become zero by one reset, it is necessary to repeat reset operation several times or lengthen the reset time, and the time until the main imaging becomes longer, and for the patient who is pressing the breast It was not preferable.

  An object of the present invention is to make it possible to quickly acquire an actual captured image having no uneven pixel signal value that occurs between a position corresponding to a line read during pre-imaging and a position corresponding to a line not read. That is.

In order to solve the above-mentioned problem, the invention described in claim 1
A radiation source for irradiating the subject with radiation;
A radiation detection element that outputs charges according to the amount of radiation that has passed through the subject has a detection unit that is arranged in a two-dimensional manner, and obtains a captured image of the subject by reading the charge generated by the detection unit A radiological image detector;
Control means for irradiating a predetermined amount of radiation from the radiation source in advance before the main photographing of the subject and causing the radiation image detector to acquire a pre-photographed image for determining photographing conditions at the time of the main photographing. A radiographic imaging system,
The radiological image detector obtains a pre-captured image by reading an area including the subject in the detection unit at a predetermined line interval during pre-imaging, and all lines of the region including the subject of the detection unit during main imaging. To obtain the actual image,
Correction means for correcting unevenness of pixel signal values generated between a position corresponding to a line read during the pre-photographing in the main captured image and a position corresponding to a line not read is provided.

The invention according to claim 2 is the invention according to claim 1,
The radiation image detector performs a reset process for removing charges remaining in the detection unit after reading in pre-imaging.

The invention according to claim 3 is the invention according to claim 1 or 2,
The radiographic image capturing system is a system that performs magnified photographing of the subject.

  According to the present invention, it is possible to quickly acquire a main captured image without pixel signal value unevenness occurring between a position corresponding to a line read during pre-photographing and a position corresponding to a line not read. It becomes.

It is a figure which shows the example of whole structure of the radiographic imaging system in this Embodiment. It is a figure for demonstrating the focus of the radiation source of FIG. It is a block diagram which shows the functional structure of the mammography apparatus of FIG. It is a figure for demonstrating the detection part of the radiographic image detector of FIG. It is a circuit block diagram of the radiographic image detector of FIG. FIG. 2 is a circuit block diagram for one pixel in the radiation image detector of FIG. 1. It is a block diagram which shows the functional structure of the console of FIG. It is a figure which shows the sequence of the expansion imaging process performed with a mammography apparatus and a console. It is a figure which shows the sequence of the expansion imaging process performed with a mammography apparatus and a console. It is a figure for demonstrating the to-be-photographed area | region and mammary gland area | region detected by S16 of FIG. 8A. (A) is a figure which shows typically the residual potential in the radiographic image detector after pre imaging | photography, (b) is a figure which shows the residual potential in the radiographic image detector after pre imaging | photography and 2nd reset typically. FIG. It is a figure which shows the relationship between the reading efficiency of a radiographic image detector, and X-ray dose. It is a figure which shows an example of the area | region extraction of the predetermined width from a real picked-up image. It is a figure which shows an example of the removal of the low frequency component from the projection corresponding to extraction area | region R1, R2, R3, (a) is a figure which shows the projection corresponding to extraction area | region R1, R2, R3, (b) (A) is a figure which shows an example of a high-pass filter, (c) is a figure which shows the projection after a low frequency component removal. It is a figure which shows an example of the production | generation of the correction data from the projection after the low frequency component removal corresponding to extraction area | region R1, R2, R3, (a) is a figure after the low frequency component removal corresponding to area | region R1, R2, R3 It is a figure which shows an example of the value removal of the edge part of a projection, (b) is a figure which shows an example of a low-pass filter, (c) is a figure which shows an example of correction data. It is a figure which shows the sequence to step S24 of the expansion imaging process B performed with a mammography apparatus and a console. It is a figure which shows the timing sequence of each step after X-ray irradiation switch depression (after step S8 of FIG. 8A) in an expansion imaging process (with 2nd reset). It is a figure which shows the timing sequence of each step after X-ray irradiation switch pressing-down (after step S8 of FIG. 15A) in the expansion imaging process B (without 2nd reset).

  Hereinafter, specific embodiments of the present invention will be described with reference to the drawings. However, the scope of the invention is not limited to the illustrated examples. In the following description, “X-ray” and “radiation” are treated synonymously.

<Embodiment 1>
(Configuration of radiation imaging system 1)
FIG. 1 is a diagram showing a system configuration of a radiographic imaging system 1 of the present invention.
As shown in FIG. 1, the radiographic image capturing system 1 includes a breast image capturing device 100 and a console 200. Data can be transmitted and received between the mammography apparatus 100 and the console 200.

(Configuration of breast image capturing apparatus 100)
First, the configuration of the mammography apparatus 100 will be described.
The breast image capturing apparatus 100 is an image capturing apparatus that performs radiation imaging using a breast as a subject H.
The mammography apparatus 100 is provided with a main body 103 that supports the radiation source 102. The main body 103 is provided with a subject table 104 that is movable in the irradiation direction of the radiation source 102 and supports the subject H, a support shaft 105 that hangs vertically with respect to the main body 103, and the like.

  Above the subject table 104, a compression plate 109 that compresses the subject H supported by the subject table 104 is supported by the main body 103 so as to be movable up and down. A pressure sensor 108 is provided on the surface of the compression plate 109 that contacts the subject H. The pressure sensor 108 may be provided on the subject table 104.

  A radiographic image detector 107 is provided below the subject table 104 so as to be movable up and down along the support shaft 105 based on the control of the control unit 111 (see FIG. 3).

  As the radiation source 102, an X-ray tube that emits X-rays having a radiation wavelength of about 0.1 to 1 mm is used. This X-ray tube emits X-rays by accelerating electrons generated by thermal excitation with a high voltage to collide with a cathode and converting the kinetic energy into radiant energy. When an X-ray image is taken, the acceleration voltage is set as a tube voltage, the amount of electrons generated as a tube current, and the X-ray emission time is set as an X-ray irradiation time. The anode (counter cathode) with which the electrons collide can change the radiated X-ray energy spectrum by changing the kind of copper, molybdenum, rhodium, tungsten or the like. When copper, molybdenum, rhodium, etc. are used as the anode, a relatively low energy line spectrum with a narrow X-ray energy distribution can be obtained, and mammography that interprets X-ray diffraction crystal analysis and fine structures using these characteristics. Used for. When tungsten is used as an anode, it is a broad spectrum, relatively high energy X-ray, which is used for human chest, abdomen, head, and general non-destructive inspection in industry. The medical or industrial use is characterized by a large X-ray dose. That is, it is common to use a rotating anode. Since the mammography apparatus 100 of the present embodiment is an apparatus used for medical purposes, an X-ray tube having a rotating anode of molybdenum, rhodium, or tungsten is preferable, and since it is an apparatus used for mammography purposes, Preferably it is molybdenum or rhodium.

As shown in FIG. 2, the X-ray tube includes a target T that generates X-rays when an electron beam generated by a cathode inside the X-ray tube collides with the X-ray tube. The size at which the electron beam collides with the target T is called the actual focus, and the substantial size viewed from the subject H side is called the effective focus. Here, the size of the effective focus is the focus size. In general, the shape of the effective focus is a square, and the length of one side is the focus size. When the shape of the focal point is a circle, the diameter is often indicated, and when the shape is a rectangle, the short side is often indicated. When the focal shape is a rectangle, the relationship between the focal point size of the real focal point and the effective focal point is as follows: the real focal point size (width: F _a , length: F _b ), the effective focal point size (width: F _A , long S: F _B ), if the target angle α, then F _A = F _a and F _B = F _b sin α.

  When the radiographic image detector 107 is brought into close contact with the object table 4 and X-ray imaging (close-contact imaging) is performed, the distance between the subject H and the radiographic image detector 107 becomes almost zero, and an image having only absorption contrast is obtained. can get.

  On the other hand, as disclosed in Japanese Patent No. 3,861,572, the distance between the X-ray tube and the subject table 104 and the distance between the radiation image detector 107 and the subject table 104 are set as a predetermined positional relationship as X. When line imaging is performed, phase contrast magnification imaging is performed, and an edge-enhanced (= phase contrast-enhanced) image caused by X-ray refraction can be obtained. X-rays are refracted when passing through the object, and the X-ray density inside the boundary of the object becomes sparse, and further, the outside of the subject H overlaps with X-rays that do not pass through the subject H, so that the X-ray density increases. In this way, the edge that is the boundary portion of the subject H is enhanced as an image. This is considered to be a phenomenon caused by a difference in refractive index with respect to X-ray between the subject H and air.

  However, in capturing a phase contrast image, a so-called penumbra occurs when the focal size of the X-ray is large. The penumbra is a phenomenon in which one point on the subject H is detected as an image having a size on the radiation image detector 107 due to the size of the focus size, and is a so-called blur. Therefore, unlike a synchrotron that emits monochromatic parallel X-rays or a micro-focus X-ray source that can be regarded as a point focus, an ordinary medical X-ray source has a finite focal size. Therefore, the influence of this penumbra becomes a problem.

  When the distance R2 between the subject H and the radiation image detector 107 is large, the blur width due to penumbra increases. For this reason, the lower limit value of the focal spot size is determined in order to obtain a certain amount of X-ray dose, and the object H and the radiation image detector 107 are obtained in order to obtain a sharp image by realizing the refraction contrast without penumbra blur. The upper limit of the focal spot size is determined from the distance R2 between the radiation source 102 and the distance R1 from the radiation source 102 to the subject H. If the focal spot size is too small, X-rays with sufficient intensity cannot be obtained, the imaging time becomes long, and the sharpness deteriorates due to blurring caused by the movement of the subject H. On the other hand, if the focal spot size is too large, blur due to penumbra dominates rather than edge enhancement by phase contrast, and sharpness decreases. Therefore, in order to perform phase contrast magnification imaging in a normal medical facility, the focal spot size needs to be 30 μm or more and 250 μm or less, and preferably 50 μm or more and 100 μm or less.

In this embodiment, the radiation source 102 includes an X-ray tube having at least two focal points having different focal diameters, and the focal diameter is switched according to the setting of the X-ray irradiation conditions.
The focal points having different focal diameters are, for example, a focal point for close-contact photographing (referred to as a large focal point) and a focal point for phase contrast photographing (referred to as a small focal point) having a smaller focal diameter than that for close-contact photographing. The focal point for close-contact photography generally has a focal diameter of about 300 μm, and the focal point for phase contrast photography generally has a small focal diameter of about 100 μm. The focal diameter is determined by a standardized measurement method such as a slit method or a point source method. Standards include JIS Z4102, Z4704, IEC60336, etc., and any standard may be used.

  An irradiation field adjuster 106 (see FIG. 3) that varies the X-ray irradiation field is provided between the radiation source 102 and the subject table 104. This irradiation field adjuster 106 is provided with an X-ray limiter that prevents the X-ray from being irradiated outside the predetermined range of the irradiation field on the subject table 104 by, for example, changing the X-ray aperture. What was used is mentioned.

  The radiation image detector 107 is an FPD (Flat Panel Detector) that detects X-rays emitted from the radiation source 102 with a radiation detection element such as a photodiode and acquires the image as digital image data. In the present embodiment, the radiation image detector 107 includes a scintillator and the like, and will be described as a so-called indirect FPD that obtains an electric signal by converting emitted X-rays into electromagnetic waves of other wavelengths such as visible light. A so-called direct type FPD that directly detects X-rays with a radiation detection element without using a scintillator or the like may be used. Details of the radiation image detector 107 will be described later.

  The housing 110 is provided with a control unit 111, an operation unit 112, a display unit 113, a communication unit 114, a power supply unit 117, and the like illustrated in FIG.

  As shown in FIG. 3, the control unit 111 includes a radiation source 102, an irradiation field adjuster 106, a reading control unit 22, a pressure sensor 108, an operation unit 112, a display unit 113, a communication unit 114, a driving device 115, and position detection. The unit 116 and the power source unit 117 are connected via a bus 118 to centrally control the operation of each unit.

  The control unit 111 includes a CPU (Central Processing Unit), a ROM (Read Only Memory), and a RAM (Random Access Memory). The ROM of the control unit 111 stores a control program for controlling each part of the mammography apparatus 100 and various processing programs. The CPU cooperates with these control programs and each part of the mammography apparatus 100. Overall control of the operation.

  In the present embodiment, the control unit 111 executes enlargement photographing processing, performs pre-photographing for determining photographing conditions at the time of main photographing before the main photographing of the subject H, and pre-photographed images acquired by pre-photographing. The main shooting is performed according to the shooting conditions calculated in the console 200 based on the image data.

The operation unit 112 includes an X-ray irradiation switch for instructing X-ray irradiation, a keyboard, a touch panel, and the like for inputting patient information, an imaging direction, an enlargement ratio at the time of enlargement imaging, and the like. A key pressing signal and a mouse operation signal that have been pressed are output to the control unit 111 as input signals. In the present embodiment, the X-ray irradiation switch has a two-stage configuration. The operation unit 112 includes an operation mechanism for adjusting the positions of the subject table 104 and the compression plate 109.
The display unit 113 includes a monitor such as an LCD (Liquid Crystal Display) and displays various screens in accordance with instructions of a display signal input from the control unit 111.

  The communication unit 114 includes a LAN (Local Area Network) card or the like, and performs data transmission / reception with the console 200 via a communication network N such as a LAN.

  The driving device 115 includes a motor that rotates the main body 103 in accordance with a control signal input from the control unit 111, a motor that moves the radiographic image detector 107 up and down in accordance with a control signal input from the control unit 111, and the like.

  The position detection unit 116 detects the position of the subject table 104 and the position of the compression plate 109 and outputs the detected position to the control unit 111. The position detection unit 116 may employ a method using photometry using infrared rays, a method in which a line resistance is provided on the rail of the support shaft that slides the subject table 104 or the compression plate 109, and the position is determined based on the resistance value measurement.

(Configuration of radiation image detector 107)
Hereinafter, the radiation image detector 107 will be described with reference to FIGS. 4 is a diagram illustrating the detection unit 107a of the radiation image detector 107, FIG. 5 is a circuit block diagram of the radiation image detector 107, and FIG. 6 is a circuit block diagram of one pixel in the radiation image detector 107.

As shown in FIG. 5, the radiation image detector 107 includes a detection unit 107 a in which a plurality of radiation detection elements 7 are two-dimensionally arranged. As shown in FIG. 4, the detection unit 107a includes two reading areas (hereinafter referred to as block areas) A and B.
Each radiation detection element 7 is constituted by, for example, a photodiode or the like, and the amount of electromagnetic waves (scintillator) output by converting X-rays incident from the radiation incident surface R by a scintillator (not shown) in the radiation image detector 107. The amount of electric charge (signal amount) that is finally accumulated differs depending on the dose of X-rays incident on. In addition to this, for example, a phototransistor or the like can be used as the radiation detection element 7.

  Each radiation detection element 7 disposed in the block region A has a second electrode 78 (see FIG. 6) connected to the bias line 9A and the connection 10A, respectively. Each radiation detection element 7 disposed in the block region B has a second electrode 78 connected to the bias line 9B and the connection 10B, respectively. Each connection 10A, 10B is connected to a reverse bias power supply 14.

  The reverse bias power source 14 is connected to the reading control unit 22, and in accordance with the control from the reading control unit 22, the reverse bias is applied to each radiation detection element 7 via the connection lines 10A and 10B and the bias lines 9A and 9B. A voltage is applied.

  The current detection means 42A and the current detection means 42B are provided on the connections 10A and 10B of the bias lines 9A and 9B, respectively. The current detection means 42 </ b> A and 42 </ b> B are connected to the reading control unit 22.

  The current detection means 42A and 42B independently detect the current flowing in the connection 10A in which the bias lines 9A are bundled and the current flowing in the connection 10B in which the bias lines 9B are bundled. Specifically, although not shown, the current detection means 42A and 42B measure a voltage between a resistor having a predetermined resistance value connected in series to the connection lines 10A and 10B, and both terminals of the resistor, respectively. And a differential amplifier. Then, by measuring the voltage between the two terminals of the resistor with a differential amplifier, the current flowing through the connections 10A, 10B is detected by converting it to a voltage value, and the voltage value corresponding to the current value flowing through each of the connections 10A, 10B Are output to the reading control unit 22, respectively.

  The first electrode 74 of each radiation detection element 7 is connected to the source electrode 8 s (denoted as S in FIGS. 5 and 6) of the TFT 8. A gate electrode 8 g (denoted as G in FIGS. 5 and 6) of each TFT 8 is connected to each scanning line 5 extending from the scanning drive circuit 15. A drain electrode 8d (denoted as D in FIGS. 5 and 6) of each TFT 8 is connected to each signal line 6.

When a reverse bias voltage is applied to the second electrode 78 of the radiation detection element 7 via the bias lines 9 </ b> A and 9 </ b> B by the reverse bias power supply 14, an electrical gradient is generated in the radiation detection element 7. When X-rays are irradiated from the radiation source 102 in this state, and the X-rays are converted into electromagnetic waves by a scintillator (not shown) and enter the radiation detection element 7, the radiation detection element 7 is proportional to the number of photons of the incident electromagnetic waves. As a result, electron-hole pairs are generated.
Among the generated electron-hole pairs, electrons move to the first electrode 74 side (see FIG. 6), which has a high potential, according to the potential gradient. When the gate of the TFT 8 is closed, the generated electrons are accumulated in the vicinity of the first electrode 74.
On the other hand, of the generated electron-hole pairs, the holes move to the second electrode 78 having a low potential according to the potential gradient, pass through the second electrode 78 to the connections 10A and 10B via the bias lines 9A and 9B. Flows out. Holes flowing through the connection 10 </ b> A are detected as current by the current detection means 42 </ b> A, and a voltage value corresponding to the current is output to the reading control unit 22. Holes flowing through the connection 10 </ b> B are detected as current by the current detection unit 42 </ b> B, and a voltage value corresponding to the current is output to the reading control unit 22.

When the voltage value output from the current detection unit 42A exceeds a predetermined threshold value, the reading control unit 22 detects that the block region A has been irradiated with X-rays. Further, when the voltage value from the current detection unit 42B exceeds a predetermined threshold value, the reading control unit 22 detects that the block region B has been irradiated with X-rays.
In the present embodiment, the reading control unit 22 is described as applying a signal reading voltage to the gate voltage 8g of the TFT 8 based on the reading control signal transmitted from the control unit 111. However, the reading control unit 22 A signal readout voltage may be applied to the gate electrode 8g of the TFT 8 in the block region where the X-ray irradiation is automatically detected.

  When a signal reading voltage is applied from the scanning drive circuit 15 to the gate electrode 8 g of the TFT 8 through the scanning line 5 by the control from the reading control unit 22, the gate of the TFT 8 is opened and accumulated in the radiation detection element 7. The charged charge, that is, an electric signal is read from the drain electrode 8d to the signal line 6 through the source electrode 8s of the TFT 8. Each signal line 6 in the block area A is connected to the readout circuit 17A. Each signal line 6 in the block region B is connected to the readout circuit 17B.

As shown in FIG. 5, the readout circuits 17A and 17B are composed of an amplifier circuit 18, a correlated double sampling circuit 19, an analog multiplexer 20, and an A / D converter 21, The charges read from the radiation detection elements 7 through the signal lines 6 are converted into electric signals by performing charge voltage conversion and amplification etc. for each radiation detection element 7.
In FIGS. 5 and 6, the correlated double sampling circuit is expressed as CDS. In FIG. 6, the analog multiplexer 20 is omitted.

The amplifier circuit 18 is constituted by, for example, a charge amplifier circuit, and is constituted by connecting an operational amplifier 18a, capacitors C1 to C2 and a charge reset switch SW1 in parallel with the operational amplifier 18a. A switch SW2 is connected in series to the capacitor C2, and on / off of the charge reset switch SW1 and the switch SW2 is controlled by the read control unit 22.
The signal line 6 is connected to the inverting input terminal 18a1 on the input side of the operational amplifier 18a of the amplifier circuit 18, and the non-inverting input terminal 18a2 on the input side of the amplifier circuit 18 is grounded (GND). That is, this embodiment corresponds to the case where the initial voltage is set to 0 [V].

  Hereinafter, the case where the non-inverting input terminal 18a2 on the input side of the amplifier circuit 18 is grounded as described above will be described. However, a predetermined initial voltage is applied to the non-inverting input terminal 18a2 on the input side of the amplifier circuit 18. It is also possible to configure as described above.

  In the amplifier circuit 18, when the gate of the TFT 8 of the radiation detection element 7 is opened with the charge reset switch SW1 turned off (that is, when a signal readout voltage is applied to the gate electrode 8g of the TFT 8), the capacitor C1 The charges read from the radiation detection element 7 are accumulated in the capacitors C1 and C2 when the switch SW2 is turned on, and a voltage value corresponding to the accumulated charge amount is output from the output terminal 18a3 of the operational amplifier 18a. It is output. In this way, the amplification circuit 18 outputs a voltage value according to the amount of charge output from each radiation detection element 7 and amplifies the voltage by voltage conversion.

In the present embodiment, capacitors having capacitances of 0.5 pF and 0.5 pF are used as the capacitors C1 and C2, respectively, and the capacitance of the capacitors C1 and C2 as a whole is controlled by turning on / off the switch SW2 under the control of the reading control unit 22. Can be set in two stages of 0.5 pF or 1 pF. When the switch SW2 is turned off, the gain can be largely adjusted.
On the other hand, when the charge reset switch 18c is turned on by the control from the reading control unit 22, the input side and the output side of the amplifier circuit 18 are short-circuited, and the charge accumulated in the capacitor 18b is discharged and amplified. Circuit 18 is reset.
It should be noted that the range width of the capacitance set for the entire capacitor in the amplifier circuit 18 and the capacitance to be set are appropriately set according to the performance required for the radiation image detector 107, and the capacitance of each capacitor for realizing it. Etc. are also determined as appropriate.

  A correlated double sampling circuit 19 is connected to the output side of the amplifier circuit 18. The correlated double sampling in the correlated double sampling circuit 19 is performed as follows.

  In the radiation image detector 107, the capacitors C1 and C2 are turned on and the charge reset switch SW1 is turned on before the gate of the TFT 8 of each radiation detection element 7 is opened for signal readout. A so-called reset is performed to release the charges accumulated in the capacitors C1 and C2. Thereafter, the charge reset switch SW1 is turned off, and a signal reading standby state is set. The correlated double sampling circuit 19 first holds the voltage value output from the output side terminal of the amplifier circuit 19 at that stage. In addition, the gates of the TFTs 8 of the radiation detection elements 7 are opened, electric signals are read from the radiation detection elements 7, charges are accumulated in the capacitors C1 and / or C2 of the amplifier circuit 18, and then the gates of the TFTs 8 are closed. At this stage, the correlated double sampling circuit 19 again holds the voltage value output from the output side terminal of the amplifier circuit 18. The correlated double sampling circuit 19 calculates the difference between the two voltage values held in this way and outputs an analog value of the electrical signal from the radiation detection element 7. In this way, the correlated double sampling circuit 19 reduces noise when the capacitor is reset.

  The electric signal output from the correlated double sampling circuit 19 is sequentially transmitted to the A / D converter 21 via the analog multiplexer 20 and converted into a digital value by the A / D converter 21. The A / D converter 21 sequentially outputs the electrical signal of each radiation detection element 7 converted into a digital value to the reading control unit 22.

  Here, the block areas A and B are individually switched to the reading mode or the standby mode according to the control of the reading control unit 22. In the reading mode, power is supplied from the power supply unit 117 to the reading circuits 17A and 17B under the control of the reading control unit 22, and the charge amplifier circuit constituting the amplifier circuit 18 is put into an operating state. On the other hand, in the standby mode, the read circuits 17A and 17B themselves are not energized, and the charge amplifier circuit that constitutes the amplifier circuit 18 is inactivated.

  When the charge amplifier circuit constituting the amplifier circuit 18 is deactivated, a current flows between the inverting input terminal 18a1 and the non-inverting input terminal 18a2 (see FIG. 6) on the input side of the operational amplifier 18a of the amplifier circuit 18. Disappear. Therefore, since the loop electrically connected to ground (GND) → reverse bias power supply 14 → (current detection means 42A (42B)) → radiation detection element 7 → TFT 8 → signal line 6 is broken at the operational amplifier 18a, radiation detection is performed. A closed loop including the element 7 and the reverse bias power source 14 cannot be formed.

  Therefore, in this embodiment, as shown in FIG. 6, each operational amplifier 18a of the amplifier circuit 18 of the readout circuit 17A (17B) is grounded to the inverting input terminal 18a1 to which the signal line 6 is connected as described above. A mode changeover switch 24 that is connected to the non-inverting input terminal 18a2 and switches between short-circuit and cancellation of the short-circuit is provided on the upstream side of each operational amplifier 18a.

  In this embodiment, the mode changeover switch 24 is composed of a MOSFET (MOS field effect transistor). A gate electrode (not shown) of the MOSFET, which is the mode switching switch 24, is connected to the reading control unit 22, and the mode switching is performed by switching the application of the voltage from the reading control unit 22 to the gate electrode and stopping of the application. The on / off state of the switch 24 is controlled.

  When the reading circuit 17A (17B) is set to the standby mode, the reading control unit 22 does not energize the reading circuit 17A (17B), sets the charge amplifier circuit constituting the amplifier circuit 18 in a non-operating state, The mode changeover switch 24 is turned on to short-circuit the inverting input terminal 18a1 and the non-inverting input terminal 18a2 of each operational amplifier 18a of the amplifier circuit 18 of the readout circuit 17A (17B).

  Thus, when each mode changeover switch 24 is turned on, as shown in FIG. 6, even if the charge amplifier circuit constituting the amplifier circuit 18 is in the non-operating state, ground (GND) → reverse bias power supply A closed loop of 14 → (current detection means 42) → image sensor 7 → TFT 8 → mode changeover switch 24 → ground (GND) is formed.

  Further, when the reading circuit 17A (17B) is set to the reading mode described above, the reading circuit 17A (17B) is energized, the charge amplifier circuit constituting the amplifier circuit 18 is set in the operating state, and the mode changeover switch is set. 24 is turned off.

  Image data output from the readout circuits 17A and 17B is stored in the RAM in the reading control unit 22, and is output to the console 200 via the wireless communication unit 29 in the radiation image detector 107.

(Console 200 configuration)
Next, the configuration of the console 200 will be described.
The console 200 receives image data of a pre-photographed image acquired by pre-photographing from the mammography apparatus 100, and calculates a photographing condition for the main exposure based on the received image data. Further, the image data of the actual captured image acquired by the actual imaging is received from the breast image capturing apparatus 100, and the received image data is subjected to image processing and displayed.

  FIG. 7 is a block diagram showing a functional configuration of the console 200. As shown in FIG. 7, the console 200 includes a control unit 211, an operation unit 212, a display unit 213, a storage unit 214, a wireless communication unit 215, and a communication unit 216, and each unit is connected via a bus 217. ing.

  The control unit 211 includes a CPU, a ROM, and a RAM. The ROM of the control unit 211 stores a control program for controlling each part of the console 200 and various processing programs, and the CPU controls the operation of each part of the console 200 in cooperation with these control programs. Control.

  The operation unit 212 includes a keyboard having cursor keys, numeric input keys, various function keys, and the like, and a pointing device such as a mouse, and controls an instruction signal input by key operation or mouse operation on the keyboard. 211 is output. The operation unit 212 may include a touch panel on the display screen of the display unit 213, and in this case, an instruction signal input via the touch panel is output to the control unit 211.

  The display unit 213 is configured by a monitor such as an LCD or a CRT (Cathode Ray Tube), and displays an input instruction, data, and the like from the operation unit 212 in accordance with an instruction of a display signal input from the control unit 211.

  The storage unit 214 is configured by a nonvolatile semiconductor memory, a hard disk, or the like. The storage unit 214 stores image data transmitted from the mammography apparatus 100 and subjected to image processing.

  The wireless communication unit 215 includes a wireless LAN card or the like, and performs data transmission / reception with the radiation image detector 107 wirelessly.

  The communication unit 216 is configured by a LAN card or the like, and performs data transmission / reception with the mammography apparatus 100 via the communication network N.

(Operation of Radiation Imaging System 1)
Next, operation | movement of the radiographic imaging system 1 is demonstrated.
8A to 8B are diagrams illustrating a sequence of enlargement imaging processing executed by the mammography apparatus 100 and the console 200 when enlargement imaging is instructed by the operation unit 112 or the operation unit 212. The processing on the mammography apparatus 100 side shown in FIGS. 8A to 8B is executed under the control of the control unit 111, and the processing on the console 200 side is executed under the control of the control unit 211.

  In the breast image photographing apparatus 100, when a photographing direction and an enlargement ratio are input by the operation unit 112 (step S1), the driving unit 115 is controlled by the control unit 111 based on the input photographing direction, and The rotation is controlled (step S2). For example, when the input photographing direction is MLO for photographing the breast from the oblique direction, the driving unit 115 rotates the main body 103 by a predetermined amount so as to be inclined.

  Next, when the position of the subject table 104 is adjusted in the vertical direction in accordance with the operation of the operation unit 112 in view of the height of the subject by an engineer or the like, the radiation image detector 107 is controlled by the control of the driving device 115 by the control unit 111. Is adjusted to a position corresponding to the enlargement ratio input in step S1 (step S3). When the position of the compression plate 109 is adjusted by the operation of the operation unit 112 and the subject H is pressed and fixed (step S4), pre-shooting shooting conditions are set by the control unit 111 (step S5). In step S5, for example, the breast thickness of the subject H is calculated based on the position of the compression plate 109 and the position of the subject table 104 output from the position detection unit 116. Then, the imaging condition corresponding to the calculated breast thickness is read from the table stored in the ROM of the control unit 111, and the imaging condition is set. The imaging conditions are mainly X-ray irradiation conditions, such as tube voltage of the radiation source 102, tube current, filter presence / absence, X-ray irradiation time, and the like. Note that the pre-photographing is photographing for determining the photographing conditions of the main photographing, and the pre-photographed image acquired by the pre-photographing is not used for diagnostic interpretation. Therefore, in X-ray irradiation (pre-irradiation) in pre-imaging, X-rays are irradiated with a smaller dose than in main imaging in order to suppress the exposure amount of the subject H.

  Next, the reading conditions at the time of pre-imaging are transmitted from the control unit 111 to the reading control unit 22, and the reading conditions for pre-imaging are set in the radiation image detector 107 (step S6). Specifically, when the switch SW2 of each pixel of the radiation image detector 107 is set to OFF, the reading gain is set to be larger. This is to make it easier to analyze the histogram when calculating the photographing condition by increasing the density (signal value) resolution of the pre-photographed image. Next, the irradiation field adjuster 106 is controlled in accordance with the operation of the photographer, and the irradiation field on the radiation image detector 107 is adjusted (step S7).

  When the X-ray irradiation switch (first stage) of the operation unit 112 is pressed (step S8; YES), the control unit 111 transmits a first reset start signal to the reading control unit 22 of the radiation image detector 107, Reset (first reset) is performed in the radiation image detector 107 (step S9). In the radiation image detector 107, when the first reset start signal is received, the mode switch 24 is turned on under the control of the reading control unit 22, and each radiation detection element 7 is connected from the reverse bias power supply 14. A reverse bias voltage is applied to. Next, the charge reset switches SW1 and SW2 of all the amplifier circuits 18 are turned on. In addition, a voltage for signal readout is applied from each scanning drive circuit 15 to the gate electrodes 8g of the TFTs 8 of all the radiation detection elements 7 via the scanning lines 5, and all the TFTs 8 are turned on. By this reset, excess charges accumulated in the radiation detection element 7 and the capacitors C1 and C2 can be discharged to the bias lines 9A and 9B.

  When the predetermined time has elapsed, the control unit 111 transmits a first reset stop signal to the reading control unit 22, and the radiation image detector 107 is shifted to the first accumulation mode (step S10). In the radiation image detector 107, when the first reset stop signal is received, the charge reset switch SW1 and the switch SW2 are turned off under the control of the reading control unit 22. Further, when the off voltage is applied to the gate voltage 8g of all the TFTs 8 and all the TFTs 8 are turned off, the mode shifts to the first accumulation mode.

  Next, when the X-ray irradiation switch (second stage) of the operation unit 112 is pressed (step S11; YES), the radiation source 102 is controlled by the control unit 111 based on the imaging conditions set in step S5. Pre-irradiation is performed using a large focal point in the source 102 (step S12). In mammography, the breast is pressed from pre-photographing to main-photographing. Therefore, X-ray irradiation is performed with a large focal point in order to shorten the pre-photographing time as much as possible. In the radiation image detector 107, when X-rays are irradiated, charges corresponding to the X-ray dose transmitted through the subject H are accumulated in each radiation detection element 7.

When the pre-imaging is completed, a first reading control signal is transmitted from the control unit 111 to the reading control unit 22, and the radiographic image detector 107 reads the pre-photographed image (step S13).
Specifically, the reading control unit 22 is instructed by the first reading control signal as to the reading interval (how many lines (scanning lines) are read). The reading interval is not limited to an equal line interval. For example, the reading interval may be controlled so that the reading interval increases from the chest wall side useful for calculation of imaging conditions toward the opposite side (so that the thinning rate increases). Good. For example, the reading interval is set to 2 ^{n in} order from the chest wall side (n is an integer equal to or greater than zero). In this way, the data amount of the pre-photographed image can be reduced.

  For example, when the area of the detection unit 107a of the radiation image detector 107 is sufficiently large and the subject H is arranged only in one of the block regions A or B, the control unit 111 places the subject H. An instruction to read only the block area to be processed may be transmitted to the reading control unit 22.

  In the radiation image detector 107, when the first reading control signal is received, each mode changeover switch 24 is turned on and the reading circuit 17A or 17B is energized under the control of the reading control unit 22. Is called. Then, the scanning drive circuit 15 applies an on voltage to the gate voltage 8g of the TFT 8 at a reading interval specified by the first reading control signal, and the TFT 8 is turned on. Then, the charges accumulated in the radiation detection element 7 are read out and digitally converted by the readout circuits 17A and 17B, so that a pre-photographed image is read out.

  When the reading of the pre-photographed image is completed, the radiation image detector 107 is reset (second reset) (step S14). Further, in the radiation image detector 107, image data of the pre-captured image is transmitted to the console 200 by the wireless communication unit 29 (step S15). Of course, wired communication may be used.

  In the console 200, when the image data of the pre-captured image is received by the wireless communication unit 215, the control unit 211 analyzes the pre-captured image (step S16).

In step S16, the subject region and the mammary gland region are detected from the pre-photographed image. As a method for detecting the subject area and the mammary gland area, the following processing can be used.
In the detection of the subject area, first, a differential filter such as pre-witt is applied to the pre-captured image in the nipple-chest wall direction X (the short axis direction of the image) shown in FIG. The position within the range is set as a skinline candidate point. Skin line candidate points are obtained for all rows in the long axis direction Y of the image, and a line connecting the skin line candidate points is detected as a skin line Sa. Then, an area between the detected skin line Sa and the chest wall is detected as a subject area Ha.
In detection of a mammary gland region, first, a local region is set in the detected subject region Ha, and the local region is binarized with a threshold value determined from the pixel signal value of each pixel in the local region. As the threshold value, a value obtained by subtracting a certain amount from the average value of the pixel signal values of each pixel in the local region or a value multiplied by a coefficient less than 1 is used. A local area is set for each predetermined interval in the subject area Ha, threshold value determination and binarization are repeated for each local area, and an area having a smaller pixel signal value than the threshold is extracted from the obtained binarized image. To do. Thereby, first, a low signal region including the mammary gland region Hb and the pectoral muscle region Hc is extracted. Next, in the above-mentioned image obtained by applying a differential filter such as prewitt to the pre-photographed image, an edge row along a diagonal line between the skin line Sa and the chest wall is detected as a pectoral muscle line Hd, A region surrounded by the pectoral muscle line Hd and the chest wall is recognized as a pectoral muscle region Hc. Then, the breast region Hb is detected by removing the pectoral muscle region Hc from the low signal region.

Next, the photographing conditions for the main photographing are calculated based on the analysis result by the control unit 211 (step S17).
In step S17, for example, a histogram of the pixel signal values in the mammary gland region Hb detected in step S17 is created, and a pixel signal value having a predetermined ratio of the cumulative frequency from the minimum value is calculated. Then, from the calculated pixel signal value, pre-imaging X-ray irradiation time, and pixel signal value desired in the main image (pixel signal value corresponding to the calculated pixel signal value), the X-ray irradiation time in main imaging is determined. calculate. The X-ray irradiation time at the time of pre-imaging is acquired from the mammography apparatus 100. When the calculated X-ray irradiation time is longer than a predetermined value, imaging is performed so that imaging is performed with a higher tube voltage, a filter that transmits higher energy X-rays, and an anode that emits X-rays with higher energy distribution. Change the condition. At that time, when the X-ray irradiation is performed under the changed imaging conditions, the X-ray irradiation time is calculated so that the calculated pixel signal value becomes a desired pixel signal value.

  The calculated imaging conditions are transmitted to the mammography apparatus 100 by the communication unit 216 (step S18).

  In the mammography apparatus 100, when the imaging condition is received by the communication unit 114, the imaging condition received by the control unit 111 is set (step S19). Further, the control unit 111 transmits a second reset stop signal and a control signal for setting a smaller gain to the reading control unit 22. The radiation image detector 107 receives the second reset stop signal and shifts to the second accumulation mode (step S20). Specifically, when the radiographic image detector 107 receives the second reset stop signal, the charge reset switch SW1 and the switch SW2 are turned off by the control of the reading control unit 22. Reset is stopped and the second accumulation mode is entered. Further, the switch SW2 is turned on, and the gain is set to be smaller.

  When the radiation image detector 107 shifts to the second accumulation mode, the radiation source 102 is controlled by the control unit 111, and main irradiation (X-ray irradiation in main imaging) is performed using a small focal point under the set imaging conditions. (Step S21).

  When the X-ray irradiation time and the second accumulation mode time elapse, the control unit 111 transmits a second reading control signal to the reading control unit 22 and the radiographic image detector 107 reads the actual captured image ( Step S22). In step S23, pixels in the entire area of the detection unit 107a are read.

  When the radiographic image detector 107 finishes reading the actual captured image, the radiographic image detector 107 is reset (third reset) under the control of the reading control unit 22 (step S23). Further, the image data of the actual captured image is transmitted to the console 200 by the wireless communication unit 29 (step S24).

  In the console 200, when image data of the actual captured image is received by the wireless communication unit 215, the control unit 211 creates a thumbnail image by thinning out the actual captured image at a predetermined pixel interval (step S25). Note that the actual captured image before thinning is stored in the RAM of the control unit 211.

  Next, correction processing is performed on the thumbnail image by the control unit 211 (step S26).

  In step S <b> 26, first, offset correction, gain correction, and image missing correction are performed on the actual captured image using a correction image prepared in advance stored in the storage unit 214. Even if X-rays with the same dose are evenly irradiated, the pixel signal values output by the pixels are not the same due to variations in the pixels of the radiation image detector 107. The correction process corrects the variation of each pixel. The correction image prepared in advance is a dark image acquired at the time of manufacture or most recently, an image uniformly irradiated with a specified dose of X-rays, an image showing a defective pixel position, and the like. The dark image is an image acquired when reading is performed by the radiation image detector 107 without X-ray irradiation.

  Next, unevenness correction is performed. Here, as described above, at the time of pre-imaging, not all the pixels of the detection unit 107a of the radiation image detector 107 are read, but thinning-out reading is performed at a predetermined line interval. Therefore, as shown in FIG. 10A, in the radiographic image detector 107 after the pre-imaging, the residual potential of the pixel located at the line L1 read at the pre-imaging is Lread, and is not read at the pre-imaging. If the residual potential of the pixel located on the line L2 is L, Lread <L. When the second reset is performed, the difference between Lread and L decreases as shown in FIG. However, since the pre-imaging is X-ray irradiation with a small dose, as shown in FIG. 11, the readout efficiency of the accumulated charges in the radiographic image detector 107 is low, and a residual potential remains even after the second reset. When the main shooting is performed in such a state of Lread <L, the pixel signal value at the position corresponding to the line L1 read at the time of pre-shooting in the main shooting image is a line that has not been read at the time of pre-shooting. Unevenness occurs in which the pixel signal value at the position corresponding to L2 becomes smaller. Therefore, in this embodiment, unevenness correction processing for correcting this unevenness is performed.

Here, as a method of unevenness correction processing, the correction method disclosed in JP-A-2005-303448 can be used. In the following description, the case where the unevenness correction process is performed using the image 40 shown in FIG. 12 as the actual captured image will be described as an example.
First, a plurality of regions (for example, R1, R2, and R3 shown in FIG. 12) having a predetermined width I1 in the main scanning direction perpendicular to the sub-scanning direction in which unevenness occurs are extracted from the corrected captured image. A plurality of projections are created in which the pixel signal values of each pixel are aggregated in the main scanning direction corresponding to the plurality of extraction regions. Here, the main scanning direction refers to the row direction, that is, the direction parallel to the reading line (scanning line), and the sub-scanning direction refers to the column direction, that is, the direction orthogonal to the row direction. FIG. 13A shows an example of projections corresponding to R1, R2, and R3. Next, an edge (a portion where the pixel signal value changes rapidly, for example, a portion where the pixel signal value changes more than a predetermined amount, E1 to E4 in FIG. 13A) is detected from the actual captured image.
Next, a predetermined high pass filter is applied to each projection to cut off low frequency components. For example, the projections R11, R12, and R13 after removal of the low-frequency components in the regions R1, R2, and R3 shown in FIG. 13C can be obtained by the high-pass filter shown in FIG. 13B. The center of the pixel signal value of each projection after removal of the low frequency component is set to zero. Further, the detected edge position and the value in the vicinity thereof are set to 0 for each projection after the removal of the low frequency component. This is because the edge portion that should not be corrected is excluded from the correction target. FIG. 14A shows the low-frequency components R1, R2, and R3 and the projections R11, R12, and R13 after removal of the edges and the vicinity thereof.
Next, all the projections after the removal of the low frequency component and the edge portion are averaged, and a high pass component is cut off by applying a predetermined low pass filter to the averaged projection, and the projection from which the high frequency component is cut off is obtained. Obtained as correction data. For example, the correction data shown in FIG. 14C is obtained by cutting off the high frequency component of the projection shown in FIG. 14A by the low-pass filter shown in FIG. 14B.
Then, the actual captured image is corrected using the obtained correction data. For example, a difference obtained by subtracting correction data from 0 is added to each pixel.
In addition, as the unevenness correction, a method of correcting the pixel signal value of the line read in the pre-photographing to the average value of the pixel signal values of the pixels adjacent to the read line may be used.

  Returning to FIG. 8B, in step S28, the control unit 211 performs gradation processing, frequency enhancement processing, and dynamic range compression processing on the thumbnail image that has been subjected to the correction processing (step S27). Then, the thumbnail image after the image processing is displayed on the display unit 213 (step S28). The photographer checks the displayed thumbnail image. Here, the process of step S25-step S27 is performed in order to display a main captured image rapidly on a display part, and to enable a photographer to confirm a main captured image.

  In the mammography apparatus 100, when a predetermined time has elapsed from the start of the third reset, the control unit 111 transmits a third reset stop signal to the reading control unit 22. The radiation image detector 107 is shifted to the third accumulation mode (step S29).

When the same time as the second accumulation mode has elapsed since the transition to the third accumulation mode, the control unit 111 transmits a third reading control signal to the reading control unit 22, and the radiographic image detector 107 accumulates the accumulated charges. Is read. Thereby, a dark image is acquired (step S30).
The acquired dark image is transmitted to the console 200 by the wireless communication unit 29 (step 31).

In the console 200, when the dark image is received by the wireless communication unit 215, the control unit 211 performs a correction process on the actual captured image (step S32). In step S <b> 32, the actual captured image stored in the RAM of the control unit 211 is read out, and offset correction, gain correction, and image missing correction are performed on the actual captured image using the dark image. In addition, unevenness correction is performed.
Next, image processing such as gradation processing, frequency enhancement processing, dynamic range compression processing, and the like is performed on the corrected main captured image by the control unit 211 (step S33). Then, the actual captured image after the image processing is displayed on the display unit 213 (step S34), and the enlargement capturing process ends.

<Embodiment 2>
Next, Embodiment 2 of the present invention will be described.
Since the configuration of the second embodiment is the same as the configuration of the first embodiment, the description will be used and the operation of the second embodiment will be described below.
FIG. 15 is a diagram illustrating a flow up to step S24 of the enlargement imaging process B executed by the mammography apparatus 100 and the console 200 in the second embodiment. Since the process after step S25 is the same as the process shown to FIG. 8B, FIG. 8B is used.
The processing on the mammography apparatus 100 side shown in FIGS. 15 and 8B is executed by the control unit 111, and the processing on the console 200 side is executed by the control unit 211.

Here, in the first embodiment, the second reset for removing the electric charge accumulated in the detection unit 107a of the radiation image detector 107 by the pre-imaging is performed. However, in the second embodiment, the second reset is performed. The difference is not done.
That is, the enlarged photographing process B shown in FIGS. 15A and 8B is different in that step S14 of the enlarged photographing process of the first embodiment shown in FIGS. 8A to 8B is omitted. The process and the flow of the process in the other steps of the enlarged photographing process B are the same as the enlarged photographing process shown in FIGS. 8A to 8B.

  FIG. 16 shows a timing sequence of each step after the X-ray irradiation switch is pressed (after step S8 in FIG. 8A) in the magnification imaging process (with the second reset) of the first embodiment. FIG. 17 shows a timing sequence of each step after the X-ray irradiation switch is pressed (after step S8 in FIG. 15A) in the magnification imaging process (without the second reset) of the second embodiment.

As shown in FIG. 17, in the magnification imaging process B, the first accumulation mode in step S20 is from immediately before the second stage X-ray irradiation switch is pressed to after the main irradiation in step S21. Since the second reset is not performed in the magnified imaging process B, the charge accumulated in the radiation detection element 7 by the main irradiation in step S21 is superimposed on the charge accumulated in the radiation detection element 7 by the pre-irradiation in step S12. Is done. For this reason, taking the first accumulation mode time into consideration when calculating the photographing conditions of the main photographing in step S17, photographing is performed with a smaller irradiation dose than in the case of performing the second reset, and the exposure dose of the subject H Can be reduced.
Further, when the second reset is performed by increasing the reading interval of each reading line in step S13 to significantly reduce the data amount of the pre-photographed image and reducing the calculation time of the photographing condition in the console 200. Compared with this, the processing time until the end of the main photographing can be shortened. As a result, the time for pressing the breast, which is the subject H, can be shortened.

On the other hand, the charge accumulation time in the dark image needs to be the same as in the acquisition of the actual captured image. That is, when the second reset is not performed, the third accumulation mode time (S29) is the same as the first accumulation mode time (S10) as shown in FIG. When the second reset is performed as shown in FIG. 16, the third accumulation mode time (S29) is the same as the second accumulation mode time (S20). Here, as shown in FIGS. 16 and 17, the first accumulation mode time without the second reset (S10 in FIG. 17) is the first accumulation mode time with the second reset (FIG. 17). 16 S10) and the second accumulation mode time (S20 in FIG. 16) is longer than the sum. Therefore, the total time until the final captured image is finally displayed on the display unit 213 of the console 200 becomes long.
In addition, in the case without the second reset, the non-uniformity of the actual captured image before the non-uniformity correction becomes larger than in the case with the second reset. Further, when there is no second reset, the accumulation mode time for the main photographing is long, so that the noise increases as compared with the case with the second reset.

  As described above, according to the radiographic image capturing system 1, the radiographic image detector 107 reads a region including the subject H in the detection unit 107a at a predetermined line interval and acquires a pre-captured image during pre-imaging. At the time of shooting, the entire photographed image is acquired by reading all lines of the area including the subject of the detection unit 107a. The console 200 corrects the unevenness of the pixel signal value generated between the position corresponding to the line read during the pre-photographing in the main captured image and the position corresponding to the line not read, and displays the corrected image on the display unit 213.

  Therefore, as compared with the conventional case, the value of each pixel acquired from the radiation image detector 107 by pre-imaging and the value of each pixel acquired after the actual imaging are added for each pixel to obtain the actual captured image. Thus, it is possible to quickly obtain a main captured image without unevenness of pixel signal values generated between a position corresponding to a line read during pre-photographing and a position corresponding to a line not read.

  In addition, the radiation image detector 107 performs a reset process for removing the charge remaining in the detection unit 107a after reading in the pre-photographing, so that unevenness of the pixel signal value of the captured image that occurs before correction, Noise can be reduced. In addition, it is possible to further shorten the time from the start of pre-shooting to the acquisition of the main shot image.

  In addition, since an image obtained by enlarging and photographing a breast has a large amount of data, the time from the start of pre-imaging to acquisition of the actual captured image can be significantly shortened compared to the conventional method.

In addition, the description in this Embodiment mentioned above is an example of the suitable radiographic imaging system which concerns on this invention, and is not limited to this.
For example, in the first and second embodiments, pixels in the entire area of the detection unit 107a are read in the main shooting. However, the subject area Ha detected from the pre-captured image is acquired from the console 200, and the subject area Ha is acquired. It is good also as reading the pixel of the area | region corresponding to. Thereby, it is possible to reduce the time required for transmission of the captured image from the radiation image detector 107 to the console 200, correction processing, and image processing.

  In the above description, an example in which a hard disk, a semiconductor nonvolatile memory, or the like is used as a computer-readable medium for the program according to the present invention is disclosed, but the present invention is not limited to this example. As another computer-readable medium, a portable recording medium such as a CD-ROM can be applied. A carrier wave is also applied as a medium for providing program data according to the present invention via a communication line.

  In addition, the detailed configuration and detailed operation of each device constituting the radiographic imaging system can be changed as appropriate without departing from the spirit of the present invention.

DESCRIPTION OF SYMBOLS 1 Radiographic imaging system 100 Mammography apparatus 102 Radiation source 103 Main body part 104 Subject stand 105 Support shaft 106 Irradiation field adjuster 107 Radiation image detector 107a Detection part 108 Pressure sensor 109 Compression plate 110 Case 111 Control part 112 Operation Unit 113 display unit 114 communication unit 115 drive unit 116 position detection unit 117 power supply unit 118 bus 22 reading control unit 29 wireless communication unit 200 console 211 control unit 212 operation unit 213 display unit 214 storage unit 215 wireless communication unit 216 communication unit 217 bus

Claims (3)

A radiation source for irradiating the subject with radiation;
A radiation detection element that outputs charges according to the amount of radiation that has passed through the subject has a detection unit that is arranged in a two-dimensional manner, and obtains a captured image of the subject by reading the charge generated by the detection unit A radiological image detector;
Control means for irradiating a predetermined amount of radiation from the radiation source in advance before the main photographing of the subject and causing the radiation image detector to acquire a pre-photographed image for determining photographing conditions at the time of the main photographing. A radiographic imaging system,
The radiological image detector obtains a pre-captured image by reading an area including the subject in the detection unit at a predetermined line interval during pre-imaging, and all lines of the region including the subject of the detection unit during main imaging. To obtain the actual image,
A radiographic imaging system comprising correction means for correcting unevenness of pixel signal values generated between a position corresponding to a line read during the pre-imaging in the main captured image and a position corresponding to a line not read.   The radiographic image capturing system according to claim 1, wherein the radiographic image detector performs a reset process for removing charges remaining in the detection unit after reading in pre-imaging.   The radiographic image capturing system according to claim 1, wherein the radiographic image capturing system is a system that performs magnified image capturing of the subject.