WO2013002087A1

WO2013002087A1 - Radiation image capturing device, method, and system

Info

Publication number: WO2013002087A1
Application number: PCT/JP2012/065713
Authority: WO
Inventors: 中津川　晴康; 俊孝阿賀野; 美広岡田
Original assignee: 富士フイルム株式会社
Priority date: 2011-06-30
Filing date: 2012-06-20
Publication date: 2013-01-03
Also published as: JP2013011553A

Abstract

A light source (14) of a radiation generation device (6) is turned on, and visible light reflected by a reflecting mirror (13) is applied to a subject (H). A photosensor array (32) of a radiation image capturing device (7) has a light receiving surface of approximately the same size as an image detection surface of an FPD (31), receives the visible light from the light source (14), and outputs light quantity signals. In the photosensor array (32), the visible light is not incident on a subject region which the subject (H) faces, and the visible light is incident on only a direct incidence region around the subject (H). Discrimination between the two regions is performed on the basis of the difference between a light quantity signal in the subject region and a light quantity signal in the direct incidence region.

Description

Radiation imaging apparatus and method, and system

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

In the medical field, there is known a radiation imaging system for imaging a subject (a region to be imaged of a patient) using radiation (for example, X-rays) to perform image diagnosis. The radiation imaging system includes a radiation generating apparatus for emitting radiation and a radiation imaging apparatus for capturing a radiation image of a subject. The radiation imaging apparatus includes a stationary type incorporated in a standing position imaging table or a reclining position imaging table, and a portable type that can be carried (so-called electronic cassette). A portable radiographic imaging device can be inserted under the patient sleeping on a bed in a hospital room or the like and imaged.

As a radiographic imaging apparatus, what utilized the radiographic image detector which has an image detection surface by which the pixel which accumulate | stores the signal charge according to the incident amount of radiation was arranged in matrix form is utilized. The radiation image detector is generally called FPD (flat panel detector). In FPD, a signal charge is accumulated for each pixel on an image detection surface to detect a radiation image and output it as digital image data.

In FPD, a direct conversion type FPD that converts radiation directly into signal charge in a conversion layer made of amorphous selenium (a-Se) or the like, and an indirect conversion that converts radiation once into visible light and converts visible light into signal charge Type FPD is known. The indirect type FPD is composed of a scintillator that converts radiation into visible light, a sensor panel disposed opposite to the scintillator, and an electrical control circuit. The sensor panel has an image detection surface in which a photoelectric conversion unit that generates signal charges by photoelectric conversion is formed for each pixel on an insulating substrate such as a glass substrate, and converts visible light from the scintillator into signal charges. Accumulate.

As a sensor panel, a TFT panel in which TFTs (thin-film transistors) and photoelectric conversion parts are arranged in a matrix on a glass substrate, or a CMOS image sensor (hereinafter referred to as a CMOS sensor) is used. The TFT is formed of an amorphous semiconductor such as amorphous silicon (a-Si). In a CMOS sensor, photoelectric conversion parts and MOS transistors are formed in a matrix on a single crystal semiconductor substrate of silicon (Si) by a semiconductor process.

Since the MOS transistor of the CMOS sensor is formed of a single crystal semiconductor, its carrier mobility is three to four orders of magnitude higher than that of a TFT panel formed of an amorphous semiconductor, enabling high-speed readout of signal charges. is there. In addition, since the variation of characteristics (for example, the threshold voltage of the MOS transistor and the like) at the time of manufacturing the photoelectric conversion unit and the MOS transistor is small in the CMOS sensor, it is possible to obtain a high S / N image. Thus, the CMOS sensor is suitable for moving image shooting and high-quality shooting.

CMOS sensors can now be manufactured using a 12 inch wafer, with one side of the square having a size of about 200 mm. Therefore, for example, a 17-inch FPD generally used for medical use can be configured using four CMOS sensors.

When performing radiography, it is preferable that the patient not be irradiated with unnecessary radiation as much as possible in order to minimize the exposure dose of the patient. Therefore, the radiation generating apparatus is provided with a collimator that limits the irradiation range of the radiation, and the radiation exposure dose of the patient is reduced by limiting the irradiation range with the collimator. For example, if the subject is smaller than the size of the entire image detection surface of the radiation image detector, for example, when the imaging site captures the hand or foot of a patient, the subject region where the object faces in the image detection surface Reduce the irradiation range according to the size of.

In the image detection plane, an area around the subject where the subject does not face is a blank area where radiation enters without passing through the subject. Radiation incident on a clear area not only does not contribute to the depiction of an object in a radiation image, but also may be irradiated to a patient by scattering or the like. Therefore, if the irradiation range is reduced in accordance with the size of the subject area, the blank area can be reduced, and the exposure dose of the patient can be reduced.

In order to accurately adjust the irradiation range by the collimator, it is necessary to accurately determine and identify the subject area and the blank area in the imaging range. In order to apply accuracy in visual discrimination, in Japanese Patent Application Laid-Open No. 2009-082169 (US Patent Publication No. US 2009/0086885) and Japanese Patent Application Laid-open No. 05-042135, a technique for automatically discriminating between a subject area and a blank area is proposed. It is done.

Japanese Patent Application Laid-Open No. 2009-082169 discloses a technique for determining an object area and a blank area using an optical camera. Specifically, first, the relative position and posture of the radiation image capturing apparatus and the subject are adjusted, and the subject is positioned with respect to the image detection surface. Since the subject and the radiation imaging apparatus are illuminated with room illumination light (such as fluorescent light) and natural light, the subject and the radiation imaging apparatus are imaged by the optical camera under the illumination light. Since the positional relationship between the radiation image capturing apparatus and the subject is recorded in the captured image, image analysis such as pattern recognition and contour extraction is performed on the captured image, and the subject region and blank area in the image detection plane Determine the area. The size of the irradiation area is determined in accordance with the size of the subject area, and the size of the irradiation opening of the collimator is adjusted.

However, since the method of JP 2009-082169 A is a method for capturing an object illuminated with room illumination or natural light and an optical image of a radiation image capturing apparatus, a pattern is used to discriminate between a subject region and a clear region. Image analysis such as recognition and contour extraction is required, and there is a concern that the apparatus configuration becomes complicated. For example, when the color of the subject and the radiographic imaging apparatus are similar such that the clothes of the subject are white and the color of the casing of the radiographic imaging apparatus is also white, highly accurate image analysis is required.

Further, in Japanese Patent Application Laid-Open No. 05-042135, a small amount of radiation is irradiated from a radiation generation apparatus, a subject is pre-photographed by a radiation image photographing apparatus, and a radiographic image obtained by pre-photographing is image analyzed There is disclosed a technique for determining the blank area and the blank area. According to the technique disclosed in Japanese Patent Application Laid-Open No. 05-042135, although the optical camera is not required, there is a problem that the exposure dose of the patient is increased by the preliminary imaging using radiation. Furthermore, even with the technique described in Japanese Patent Application Laid-Open No. 05-042135, since radiation passes through the subject, it is necessary to perform image analysis such as pattern recognition and contour extraction in order to discriminate between the subject area and the clear area. Similar to the technique described in JP-A-2009-082169, there is a concern that the apparatus configuration may be complicated.

An object of the present invention is to distinguish between a subject area and a blank area with a simple configuration without performing preliminary imaging with radiation in radiation imaging.

In order to solve the above-mentioned subject, a radiographic imaging device of the present invention is provided with a radiation generation device, a radiographic imaging device, a light source, and a field distinction part. The radiation generator comprises a radiation source for irradiating the subject with radiation. The radiation imaging apparatus has an image detection surface in which a plurality of pixels are arranged in a matrix, and includes a radiation image detector that receives radiation transmitted through a subject and detects a radiation image of the subject. The light source emits detection light to the subject positioned with respect to the image detection surface. The area determination unit has a light detection unit that receives detection light emitted from the light source and enters the periphery of the subject and outputs a light quantity signal, and based on the light quantity signal, the subject area where the subject faces in the image detection plane And an uncovered area around the subject.

The light detection unit is preferably a photo sensor array having a light receiving surface in which a plurality of photo sensors that output light quantity signals according to the amount of light received of detection light are arranged in a matrix. The photosensor array is preferably disposed closer to the light source than the radiation image detector. The photosensor array is preferably arranged such that the light receiving surface is parallel to the image detection surface of the radiation image detector. Here, parallel includes, in addition to completely parallel, substantially parallel. It is preferable that the area discrimination unit discriminates between the subject area and the blank area by comparing the light amount signals output from the plurality of photosensors with a preset threshold.

The light source is preferably provided in the radiation generation apparatus, and the detection light is preferably emitted to the subject from the same direction as the radiation. The detection light is preferably any one of visible light, infrared light and ultraviolet light. The radiation generating apparatus is composed of a plurality of shielding plates that define an irradiation opening that transmits radiation, and a collimator that limits the irradiation range in the image detection plane by adjusting the size of the irradiation opening by moving the shielding plate. It is preferable to have.

The driving amount of the collimator is set so that the irradiation range determination unit that determines the irradiation range and the irradiation range determined by the irradiation range determination unit so that the blank area is reduced based on the determination result of the area determination unit. It is preferable to include a collimator drive amount determination unit to determine.

The irradiation range determination unit specifies the number of pixels of the radiation image detector which are present in the region of the missing portion based on the determination result of the region determining portion, and the number of the missing pixels does not exceed the preset allowable value. It is preferable to determine the irradiation range as described above. It is preferable that the allowable value of the number of unfiltered pixels is set in accordance with the irradiation dose of radiation emitted from the radiation source.

The photosensor array is configured by tiling a plurality of array units having a light receiving surface in which a plurality of photosensors are arrayed, and the image detection is performed more than the array unit disposed at the center of the image detection surface. The size of the light receiving surface may be smaller in the array unit disposed at the peripheral edge of the surface.

The collimator has at least a pair of shielding plates that change the width of the irradiation opening, and the cross-sectional shape of the shielding plate is a wedge shape whose thickness increases from the edge defining the irradiation opening toward the outside Good.

The collimator may have a first collimator that limits the irradiation range of the radiation, and a second collimator that absorbs an energy component that is relatively low in energy from the radiation that is irradiated to the blank area.

The pair of shielding plates may be movable independently.

The radiation image detector includes a scintillator that absorbs radiation and converts it into light, and a sensor panel that is disposed on the radiation irradiation side of the scintillator and in which a plurality of pixels that detect light converted by the scintillator are arranged in a matrix. It is preferable to include. The sensor panel is preferably composed of a CMOS type image sensor.

The sensor panel may also be used as a light detection unit. The radiation imaging apparatus has an irradiation surface for transmitting radiation and detection light, and a housing for accommodating the sensor panel, and has transparency to radiation while shielding light to the detection light. The light shielding plate may be displaceable between a light shielding position for shielding the detection light incident on the sensor panel and a retracted position for retreating from the light shielding position.

The radiation image capturing apparatus may have an attachment removably attachable to the radiation imaging apparatus, and the attachment may be provided with a light detection unit.

The radiation image capturing apparatus of the present invention includes a radiation image detector, a light detection unit, and an area determination unit. The radiation image detector has an image detection surface in which a plurality of pixels are arranged in a matrix, receives radiation transmitted from the radiation source and transmitted through the subject, and detects a radiation image of the subject. The light detection unit receives the detection light emitted from the light source and incident on the periphery of the subject in a state in which the subject is positioned with respect to the image detection surface, and outputs a light amount signal. The area determination unit determines, based on the light amount signal, a subject area where the subject faces in the image detection plane and a blank area around the subject.

In the radiation image capturing method of the present invention, in the state where the subject is positioned with respect to the image detection surface of the radiation image detector, the step of irradiating the subject with detection light, and the detection light which is incident around the subject is received. And outputting a light amount signal, and determining a subject region facing the subject in the image detection plane based on the light amount signal and a blank region around the subject. It features.

Effect of the present invention

According to the present invention, it is possible to distinguish between a subject area and a blank area with a simple configuration without performing preliminary imaging with radiation.

It is a block diagram of a radiographic imaging system. It is a perspective view of a radiographic imaging apparatus and a collimator. It is explanatory drawing of an array unit. It is sectional drawing of a radiographic imaging apparatus. It is a bottom view which shows the sensor panel bonded together to the top plate. It is sectional drawing which shows the structure of FPD roughly. It is a graph which shows the sensitivity area of a photoelectric converting layer, and the luminescence field of a scintillator. It is a circuit diagram showing composition of a signal output circuit. It is a perspective view which shows the state in which the to-be-photographed object is mounted on the top plate. It is a block diagram which shows the principal part structure of the electric system of a radiographic imaging apparatus. It is a block diagram which shows the principal part structure of an electric system of a console and a radiation generation apparatus. It is explanatory drawing which shows the 1st determination method of irradiation range. It is explanatory drawing which shows the 2nd determination method of irradiation range. It is explanatory drawing which shows the 3rd determination method of irradiation range. It is explanatory drawing of the radiation generation apparatus provided with the 1st and 2nd collimator. It is explanatory drawing which shows the photosensor array which varied the size of the array unit for every area | region. It is an explanatory view showing a radiation generator provided with a swing mechanism. It is a conceptual diagram which shows the state which performed the swing by the swing mechanism of FIG. 15A. It is explanatory drawing of the collimator which enabled it to drive a pair of shielding board separately. It is explanatory drawing of the structure which provides a removable light-shielding plate in a top plate. It is explanatory drawing of the state which inserted the light-shielding plate in the top plate of FIG. 17A. It is explanatory drawing which shows the attachment which can be attached or detached to a radiographic imaging apparatus. It is explanatory drawing which shows the state which attached the attachment to the radiographic imaging apparatus.

First Embodiment
As shown in FIG. 1, a radiation imaging system 5 using the present invention comprises a radiation generator 6 for emitting X-rays to the imaging region of a subject (patient) H and radiation based on the X-rays transmitted through the subject H. A radiation image capturing apparatus 7 for capturing an image, and a console 8 for controlling the radiation generating apparatus 6 and the radiation image capturing apparatus 7 are provided.

The radiation generator 6 includes an X-ray tube 10, a collimator 11, a source filter 12, a reflection mirror 13, and a light source 14. The X-ray tube 10 corresponding to the radiation source of the present invention has a cathode made of a filament that emits thermal electrons, and an anode (target) that emits thermal X rays when the thermal electrons emitted from the cathode collide. X-ray toward the subject H. The x-rays emanate radially from the focal point where the thermal electrons strike at the anode.

The collimator 11 is disposed in front of the irradiation direction of the X-ray tube 10, and shields a part of the radially expanding X-rays to limit the X-ray irradiation range. As shown in FIG. 2, the collimator 11 has, for example, four

shielding plates

17a, 17b, 18a, 18b for shielding X-rays, which define a rectangular irradiation opening 11a for transmitting X-rays. Do. The collimator 11 can adjust the size of the irradiation opening 11a by moving the

shielding plates

17a, 17b, 18a and 18b. The pair of shielding

plates

17a and 17b are movable in the x direction of the xy plane perpendicular to the irradiation direction z of the X-ray. The first shielding portion 17 is configured to change the width of the irradiation opening 11 a in the x direction. The other pair of shielding

plates

18a and 18b is movable in the y direction of the xy plane, and constitutes a second shielding portion 18 that changes the width of the irradiation opening 11a in the y direction. A structure combining such two sets of shielding plates is called a so-called double leaf structure.

The two

shield plates

17a and 17b and the

shield plates

18a and 18b move in conjunction with one another so that the center of the irradiation opening 11a does not change even if the width of the irradiation opening 11a changes. For example, when one of the

shield plates

17a and 17b moves in the x direction, the other moves in the opposite direction by an amount corresponding to the movement amount. Although the width of the irradiation opening 11a changes due to the movement of the

shielding plates

17a and 17b, the center in the width direction of the irradiation opening 11a does not change because they are interlocked.

For the

shield plates

17a, 17b and 18a, 18b, lead or the like excellent in X-ray absorptivity is used, and the cross section thereof is in the z direction from the edge defining the irradiation opening 11a to the outer end Has a thick wedge shape. Thereby, in the

shielding plates

17a, 17b, 18a, and 18b, X-rays are not completely shielded near the end edge on the irradiation opening 11a side, and a part of the X-rays is transmitted.

Further, the collimator 11 has a collimator drive mechanism 19 for driving the first shielding portion 17 and the second shielding portion 18, and the size of the irradiation opening can be changed electrically. The collimator drive mechanism 19 includes, for example, a motor, and a gear and a link mechanism for transmitting the rotational force of the motor to the

shielding plates

17a, 17b, 18a, and 18b. The collimator drive mechanism 19 can drive, for example, the first shielding unit 17 and the second shielding unit 18 independently of each other, and can change the widths of the irradiation opening 11 a in the x direction and the y direction independently. . Therefore, it is possible to make the shape of the irradiation opening 11 a rectangular other than square, and it is also possible to change the ratio of the long side to the short side of the rectangle.

The radiation source filter 12 removes low energy components that scatter from the X-ray emitted from the X-ray tube 10 as they pass through the object H and cause deterioration of the image quality of the radiation image. For the source filter 12, a material having a property of absorbing only low energy components is used. As such a material, for example, aluminum is suitable. The high energy component of the X-ray transmitted through the radiation source filter 6 b is used for imaging the subject H.

The energy distribution of the X-rays emitted from the X-ray tube 10 changes with the tube voltage applied to the X-ray tube 10. For example, when the tube voltage of the X-ray tube 10 is 70 kV, the maximum energy of X-rays emitted from the X-ray tube 10 is 70 KeV, and the energy distribution of the X-rays is approximately 15 to 70 KeV. The source filter 12 absorbs a low energy component of 1/2 or less, for example, 15 to 40 KeV, of the X-ray energy distribution at a tube voltage of 70 kV, and a high energy component of 1/2 or more (40 to 70 KeV) To Penetrate. The source filter 12 is detachably attached to the X-ray generator 6 so that it can be changed according to the tube voltage.

The reflection mirror 13 and the light source 14 are used when determining the irradiation range of the X-ray by the collimator 11. The reflection mirror 13 is disposed between the X-ray tube 10 and the collimator 11, and the light source 14 is disposed to the side of the reflection mirror 13. The light source 14 comprises a lamp or the like for emitting visible light which is detection light. The reflection mirror 13 is inclined to reflect the visible light emitted from the light source 14 toward the collimator 11.

When determining the irradiation range of the X-ray by the collimator 11, the visible light is irradiated from the light source 14 toward the reflection mirror 13 in a state where the subject H is positioned with respect to the radiation imaging device 7. The visible light is reflected by the reflection mirror 13, the optical path is bent, and travels to the collimator 11. The visible light passes through the irradiation opening 11 a of the collimator 11 and is irradiated on the subject H positioned in the radiation imaging device 7. The reflection mirror 13 is formed of, for example, a material that transmits X-rays. Therefore, it is not necessary to retract the reflection mirror 13 even during X-ray irradiation, and the position of the reflection mirror 13 is fixed. Of course, the reflection mirror 13 may be retracted at the time of X-ray irradiation.

As shown in FIG. 2, the radiation image capturing apparatus 7 according to the present invention has an FPD 31 which is a radiation image detector, a photosensor array 32 which is a light detection unit, and a housing 23 which accommodates these. The housing 23 is a portable type, and the radiation imaging device 7 is a so-called electronic cassette. The housing 23 is a box having a flat overall shape, and has a rectangular planar shape. The housing 23 is configured of a housing body 25 and a top plate 24 that seals the opening of the upper portion of the housing body 25. The upper surface of the top plate 24 is an irradiation surface 22 to which X-rays emitted from the radiation generation device 6 are irradiated. When the object H is placed on the irradiation surface 22, the object H with respect to the radiation image capturing device 7 Is positioned. The housing body 25 is made of, for example, an ABS resin or the like, and the top plate 24 is made of a plastic or the like having high transparency to visible light and X-rays. Thereby, the transmittance | permeability of the visible light which the light source 14 mentioned above emits is ensured, suppressing absorption of the radiation by the top plate 24. As shown in FIG. Further, the strength is also secured by using the material of the top plate 24 of a relatively high strength plastic.

The housing 23 is, for example, the same size (for example, 17-inch square) as a conventional film cassette for recording a radiation image on a photosensitive material. The radiographic imaging device 7 has the same portability as the film cassette, and can be used instead of the film cassette.

The top plate 24 of the radiation imaging device 7 is provided with a display unit 28 configured of a plurality of LEDs. The display unit 28 displays an operation mode (e.g., "ready state" or "during data transmission" or the like) of the radiation imaging device 7 and an operation state such as the remaining capacity of the battery. The display unit 28 may be configured of a light emitting element other than an LED, a liquid crystal display, an organic EL display, or the like. In addition, the display unit 28 may be provided on the housing body 25.

In the housing 23 of the radiation imaging device 7, the photosensor array 32 and the FPD 31 are stacked in order from the top plate 24 side. The FPD 31 receives a radiation of X-rays emitted by the radiation generator 6 and detects a radiation image of the subject H.

The photosensor array 32 is disposed between the top 24 and the FPDFPD 31. The photosensor array 32 is configured by tiling a plurality of (in this example, four) array units 33. The array unit 33 has a plurality of photosensors 33a arranged in a matrix, and has a light receiving surface on which the photosensors 33a are arranged. Each photosensor 33a is, for example, a light receiving element composed of a photodiode, receives an visible light emitted from the light source 14, and photoelectrically converts it to an electric signal (light amount signal) such as a voltage signal according to the amount of light received. Output

The photosensor array 32 is an enlargement of the light receiving surface 32 a by tiling a plurality of small-sized (four in this example) array units 33 having a relatively small light receiving surface.

The photosensor array 32 is disposed with its light receiving surface 32 a facing the top 24. The light receiving surface 32a of the photosensor array 32 is substantially parallel to the image detection surface 39a (see FIG. 4) of the FPD 31, and the shape and size of the light receiving surface 32a are the shape and size of the image detection surface 39a of the FPD 31. It is almost the same.

The photosensor array 32 is an area where the subject H faces on the image detection surface 39a of the FPD 31, that is, a subject area where X-rays transmitted through the subject H are incident, and an area around the subject H, that is, the X-ray is the subject H Used in order to distinguish directly from non-passing areas that do not pass through. If the size of the subject H placed on the top 24 is smaller than the size of the image detection surface 39 a of the FPD 31, the entire area of the image detection surface 39 a does not become the subject area. A blank area occurs.

Unlike X-rays, visible light from the light source 14 is blocked without transmitting the subject H. Since the photosensor array 32 is located behind the subject H when viewed from the light source 14 side, visible light does not enter the subject area on the light receiving surface 23 a thereof, but enters only the blank area around the subject H Do. In this case, the output value of the light amount signal output from the photosensor 33a located in the object area and the output value of the light amount signal output from the photosensor 33a located in the blank area are different according to the light amount of visible light . By identifying the respective positions (coordinates in the light receiving surface 32a) of the photosensor 33a located in the subject area and the photosensor 33a located in the blank area based on this difference, the subject area and the blank area are discriminated. can do.

The photosensors 33a do not correspond one-to-one with the pixels 48a of the FPD 31, and the size and arrangement pitch of the light receiving area of each photosensor 33a are the size and arrangement of the pixels 48a (see FIG. 6) of the FPD 31. Large compared to the pitch. Therefore, the number of photosensors 33 of the photosensor array 32 is smaller than the number of pixels of the FPD 31, and the resolution is lower than that of the FPD 31. Since the photosensor array 32 does not detect an image, it is sufficient that the photosensor array 32 has a resolution enough to discriminate between the subject area and the blank area. Since the light receiving area of the photosensor 33a can be increased by the lower resolution, the light receiving sensitivity can be increased. Of course, the resolution of the photosensor array 32 may be the same as that of the FPD 31.

Since the photosensor array 32 is disposed closer to the irradiation surface 22 than the FPD 31, all X-rays incident on the FPD 31 pass through the photosensor array 32. Therefore, the attenuation of X-rays by the photosensor array 32 should be as small as possible. Therefore, as the photosensor array 32, a photosensor formed of an organic photoelectric conversion material (OPC) as a photoelectric conversion material is used. OPC can be reduced in thickness and has good X-ray transmission characteristics with almost no absorption of X-rays.

Inside the housing 23, at one end side along the lateral direction of the irradiation surface 22, a case 36 for containing various electronic circuits including a microcomputer and a rechargeable and detachable battery (secondary battery) is provided. It is arranged. Various electronic circuits of the radiation imaging device 7 including the FPD 31 operate by power supplied from a battery housed in the case 36. Note that, on the irradiation surface 22 side of the case 36 in the housing 23, radiation composed of a lead plate or the like is provided to avoid damage to various electronic circuits accommodated in the case 36 due to the irradiation of X-rays. A shielding member (not shown) is provided.

As shown in FIG. 4, the photosensor array 32 is bonded to the inner surface of the top plate 24 with an adhesive over the entire surface. Further, the FPD 31 is bonded to the lower surface of the photosensor array 32 by an adhesive. As described above, by attaching the photosensor array 32 and the FPD 31 to the top plate 24, the gaps between the respective parts are eliminated, so that the radiation imaging device 7 can be thinned and the FPD 31 can be reinforced by the photosensor array 32. The photosensor array 32 preferably has an outer size equal to or larger than that of the FPD 31 because the FPD 31 is bonded.

In the FPD 31, a sensor panel 39 and a scintillator 40 are stacked in order from the irradiation surface 22 side. A support substrate 41 supporting the scintillator 40 is disposed on the lower surface of the scintillator 40. A sealant 42 is provided on the outer periphery of the FPD 31 to protect the scintillator 40 from moisture and the like. A control board 43 is disposed on the bottom of the housing 23. The control substrate 43, the sensor panel 39 and the photosensor array 32 are electrically connected via

flexible cables

44, 45.

The scintillator 40 transmits the subject H and is irradiated to the irradiation surface 22 of the housing 23, transmits the top plate 24, the photosensor array 32 and the sensor panel 39, absorbs the irradiated X-rays, and emits light. . As the scintillator 40, for example, materials such as CsI: Tl (cesium iodide to which thallium is added), CsI: Na (sodium activated cesium iodide), GOS (Gd2O2S: Tb) and the like can be used. In the present embodiment, CsI: Tl is vapor-deposited on the support substrate 41 as the scintillator 40 to form a plurality of columnar crystals along the light emission direction from the support substrate 41 toward the sensor panel 39. The columnar crystals are approximately uniform in average diameter along the longitudinal direction of the columnar crystals.

The light generated by the scintillator 40 travels in the columnar crystal due to the light guide effect of the columnar crystal and is emitted to the sensor panel 39. At that time, the diffusion of the light emitted to the sensor panel 39 side is suppressed, so that the sharpness of the radiation image detected by the radiation imaging device 7 is improved. Further, the light reaching the deep portion of the scintillator 40 is provided on the inner surface of the support substrate 41 and light oil of the scintillator 40 is reflected again by the reflective layer and reflected to the sensor panel 39 side. The detection efficiency of the light emitted at 40 is improved.

In the present embodiment, the sensor panel 39 is disposed on the radiation-irradiated side of the scintillator 40, but the method of arranging the scintillator and the sensor panel in such a positional relationship is referred to as “surface reading method (ISS: Irradiation It is called "Side Sampling". Since the scintillator emits light more strongly on the X-ray incident side, the front reading method (ISS) in which the sensor panel is disposed on the X-ray incident side of the scintillator is arranged on the opposite side of the scintillator on the X-ray incident side. Since the sensor panel and the light emission position of the scintillator are closer than in the scanning method (PSS: Penetration Side Sampling), the resolution of the radiation image obtained by imaging is higher, and the light reception amount of the sensor panel is increased. As a result, the sensitivity of the radiation imaging apparatus is improved.

Since the FPD 31 uses an indirect type using the scintillator 40, it is assumed that visible light transmitted through the top 24 and the photosensor array 32 from the outside of the radiation imaging device 7 such as visible light emitted by the light source 14 enters the FPD 31 Since the FPD 31 detects visible light, the radiation image may be degraded. Therefore, when the light transmitted through the top plate 24 can not be blocked only by the photosensor array 32, it is preferable to provide a light shielding layer 34 having X-ray transparency between the photosensor array 32 and the FPD 31.

As shown in FIG. 5, the sensor panel 39 is configured by four CMOS type image sensors (hereinafter referred to as CMOS sensors) 48. Each CMOS sensor 48 has a plurality of pixels 48a (see FIG. 8) arranged in a matrix. Each CMOS sensor 48 has a rectangular shape with a side length of about 200 mm. The four CMOS sensors 48 are arranged adjacent to each other vertically and horizontally to form a quadrangle of approximately 17 inches on one side. The 17-inch size is a common size as a medical FPD 31 size.

As shown in FIG. 6, the CMOS sensor 48 has a configuration similar to that disclosed in US Patent Publication 2009/0224162. Specifically, the CMOS sensor 48 includes the single crystal semiconductor substrate 50, the insulating layer 54, the first electrode 51, the photoelectric conversion layer 52, and the second electrode 53.

The single crystal semiconductor substrate 50 is made of single crystal Si. The insulating layer 54 is formed of silicon oxide or the like on the surface of the single crystal semiconductor substrate 50. The first electrode 51 is individually formed on the surface of the insulating layer 54 for each pixel 48 a. The photoelectric conversion layer 52 is provided on the surface of each first electrode 51 in common to each pixel 48 a. The second electrode 53 is provided commonly to the respective pixels 48 a on the surface of the photoelectric conversion layer 52. The aforementioned scintillator 40 is bonded onto the surface of the second electrode 53 by an adhesive (not shown).

The second electrode 53 is formed of a conductive material (for example, indium tin oxide (ITO)) which is transparent to visible light so that visible light generated by the scintillator 40 may be incident on the photoelectric conversion layer 52. In the present embodiment, the second electrode 53 is provided in common to each pixel 48a, but may be provided individually for each pixel 48a.

The photoelectric conversion layer 52 generates a signal charge according to the amount of incident X-rays by the combination with the scintillator 40. The photoelectric conversion layer 52 absorbs visible light generated by the scintillator 26, and the amount of light corresponds to the amount of light. It generates signal charges and is made of an organic or inorganic photoelectric conversion material. As an inorganic photoelectric conversion material, there is, for example, amorphous silicon (a-Si). Moreover, as an organic photoelectric conversion material, there is quinacridone, for example.

As shown in FIG. 7, the sensitivity of the organic photoelectric conversion material (OPC) made of quinacridone is visible when the scintillator 40 made of CsI: Tl is generated as compared to CsI: Na, single crystal Si (c-Si) or the like. It is close to the wavelength range of light. For this reason, in the present embodiment using CsI: Tl as the scintillator 40, the photoelectric conversion layer 52 is preferably formed of quinacridone, and high detection efficiency can be obtained. In order to improve high-quality imaging and moving-image imaging, it is preferable to use an organic photoelectric conversion material having a fast carrier mobility and little variation in manufacturing as the material of the photoelectric conversion layer 52.

The single crystal semiconductor substrate 34 is provided with a signal output circuit 57 for each pixel 48 a. The signal output circuit 57 is formed of a CMOS circuit. The signal output circuit 57 and the first electrode 51 are electrically connected by the contact wiring 58. A bias voltage is applied to the second electrode 53 (see FIG. 8), and the signal charge generated by the photoelectric conversion layer 52 is collected by the first electrode 51 of each pixel 48a. The signal output circuit 57 converts the signal charge collected by the first electrode 36 into a voltage signal corresponding to the signal charge amount and outputs the voltage signal.

As shown in FIG. 8, the signal output circuit 57 includes an output transistor T1, a row selection transistor T2, a reset transistor T3, a row selection line L1, a signal output line L2, and a reset line L3. The output transistor T1, the row selection transistor T2, and the reset transistor T3 are each a MOS transistor. The row selection line L1, the signal output line L2, and the reset line L3 are formed of a metal such as aluminum in the insulating layer 54 described above.

The output transistor T1 is connected to the first electrode 51, and a voltage corresponding to the signal charge collected by the first electrode 51 is applied to the gate. The row selection transistor T2 is turned on by the selection signal applied to the row selection line L1, and a voltage signal controlled according to the gate voltage of the output transistor T1 is applied to the signal output line L2. The reset transistor T3 is turned on by the selection signal applied to the reset line L3, and discards the signal charge collected by the first electrode 51 to the power supply wiring Vdd.

As described above, since silicon is used for the single crystal semiconductor substrate 50 of the CMOS sensor 48, the carrier mobility of each of the transistors T1 to T3 is higher than that of a TFT made of an amorphous semiconductor such as a-Si. Thus, high speed reading is possible by three to four digits or more. Further, the single crystal semiconductor substrate 50 has less variation in characteristics (for example, threshold voltage and the like) at the time of manufacturing than a TFT made of an amorphous semiconductor. Therefore, it is suitable for high image quality shooting and moving image shooting that require high sensitivity, high S / N, and high speed readout. Furthermore, peripheral circuits such as a control unit of the FPD 31 can be mixedly mounted on the single crystal semiconductor substrate 50.

Of the signal output circuit 57, the row selection line L1, the signal output line L2, and the reset line L3 are formed of metal such as aluminum, so the deterioration due to X-ray is small, but the output transistor T1, the row selection transistor T2, the reset Since the transistor T3 is formed of single crystal Si, there is a possibility that the characteristics are degraded (change in threshold voltage and dark current increase) due to X-rays. This is because in the MOS structure using single crystal Si, charges (hereinafter referred to as interface charges) are generated and accumulated on the sea surface of the single crystal semiconductor substrate 50 and the insulating layer 54 by absorption of X-rays.

As shown in FIG. 9, when the imaging region of the subject H is a hand, the hand is smaller than the entire size of the image detection surface 39a (see FIG. 4) of the FPD 31. Therefore, when the entire image detection surface 39a is set as the X-ray irradiation range, all areas other than the subject area where the hand is located become a blank area where X-rays directly enter without passing through the subject H.

The X-ray incident on the blank area is an unnecessary X-ray that does not contribute to the depiction of the subject H in the radiation image, and furthermore, it may be irradiated to the patient by scattering, thus increasing the dose of the patient. Cause it to Therefore, in the present embodiment, the light source 14 and the photosensor array 32 are used to determine the subject area and the blank area, and the irradiation range is determined so as to reduce the blank area.

In addition, characteristic degradation is likely to occur in the CMOS sensor 48 in the blank region. Therefore, reducing the blank area contributes to suppression of the characteristic deterioration of the CMOS sensor 48.

In FIG. 10 schematically showing the electrical configuration of the radiation image capturing apparatus 7, the radiation image capturing apparatus 7 includes a sensor panel 39, a signal processing unit 65, an image memory 66, a control unit 67, and a wireless communication unit 68 which constitute the FPD 31 described above. , And a region determination unit 70 including the power supply unit 69 and the photosensor array 32. The signal processing unit 65 includes an amplifier for amplifying a voltage signal output from each pixel 48 a of the sensor panel 39, a multiplexer, an A / D (analog / digital) converter, etc., and the voltage output from the sensor panel 39 Convert the signal into digital image data.

An image memory 66 is connected to the signal processing unit 65, and the image data output from the A / D converter of the signal processing unit 65 is sequentially stored in the image memory 66. The image memory 66 has a storage capacity capable of storing image data for a plurality of frames, and image data obtained by imaging is sequentially stored in the image memory 66 each time a radiation image is captured.

The image memory 66 is connected to a control unit 67 that controls the overall operation of the radiation imaging device 7. The control unit 67 is configured to include a microcomputer, and includes a CPU 67a, a RAM 67b, and a ROM 67c. The RAM 67 b is a temporary storage memory made of a DRAM or the like. The ROM 67c is a non-volatile memory including a flash memory or the like.

When the console 8 receives the operation instruction for determining the irradiation range by the operation input, the console 8 instructs the radiation generator 6 to turn on the light source 14. At the same time, it instructs the radiation image capturing apparatus 7 to determine the subject area and the blank area. When the control unit 67 receives a command from the console 8, the control unit 67 operates the area determination unit 70 to execute the area determination process.

When each photosensor receives visible light from the light source 14, the photosensor array 32 outputs a light amount signal according to the amount of received light. The visible light from the light source 14 is irradiated to the entire light receiving surface 32 a of the photosensor array 32. When the subject H is placed on the top 24 of the radiographic imaging device 7 and the subject H is positioned with respect to the image detection surface 39a of the FPD 31, visible light is blocked by the subject H in the subject region Therefore, no visible light is incident on the photosensor 33a located in the subject area. On the other hand, visible light is incident on the photosensor 33a positioned in the blank area around the subject H.

The circuit unit 70a determines the subject area and the missing area based on the light amount signal output from each photosensor 33a of the photosensor array 32, generates coordinate information of each area, and generates the generated coordinate information as an area determination result. Output to the control unit 67.

The circuit unit 70a compares the read light amount signal with a preset threshold value and outputs a comparison result by sequentially selecting a light amount signal output from each of the photosensors 33a arranged in a matrix. And a memory for storing the comparison result.

The comparator, for example, compares the magnitudes of the threshold voltage and the input voltage, and outputs the comparison result. Specifically, “Hi” is output when the light amount signal is equal to or greater than the threshold, and “Low” is output when the light amount signal is less than the threshold. The threshold value is that of the visible light emitted by the light source 14 such that the output value of the photosensor 33a receiving the visible light emitted by the light source 14 is equal to or higher than the threshold, and the output value of the photosensor 33a not receiving the visible light is smaller than the threshold. It is set according to the light emission amount.

The circuit unit 70a stores the comparison result output from the comparator in association with the coordinates in the light receiving surface 32a of the photosensor 32 of each photosensor 33a. When the comparison result is "Hi", the circuit unit 70a determines that the coordinates of the photosensor 33a that has output the light amount signal belongs to the blank area, and identifies the coordinates and the blank area. Information is stored in memory in association with it. On the other hand, if the comparison result is "Low", it is determined that the coordinates of the photosensor 33a that has output the light quantity signal belongs to the subject area, and the coordinates are associated with identification information indicating that it is the subject area. Store in memory. The circuit unit 70a performs such processing on all the photosensors 33a, and outputs coordinate information of each of the subject area and the blank area stored in the memory to the control unit 67 as an area discrimination result.

A wireless communication unit 68 is connected to the control unit 67. The wireless communication unit 68 is an IEEE (Institute
Electrical and Electronics Engineers) Supports wireless LAN (Local Area Network) standards represented by 802.11a / b / g / n etc., and controls transmission of various information to / from external equipment by wireless communication . The control unit 67 can wirelessly communicate with the console 8 (see FIG. 11) via the wireless communication unit 68, and can transmit and receive various types of information to and from the console 8.

In addition, the radiation image capturing apparatus 7 is provided with a power supply unit 69, and the various electronic circuits described above (the signal processing unit 65, the image memory 66, the control unit 67, the wireless communication unit 68, etc.) are connected to the power supply unit 69, respectively. (Not shown) and operated by the power supplied from the power supply 69. The power supply unit 69 incorporates the above-described battery (secondary battery) so as not to impair the portability of the radiation imaging device 7, and supplies power from the charged battery to various electronic circuits. The signal processing unit 65, the image memory 66, the control unit 67, and the wireless communication unit 68 are provided in the case 36 or the control board 43 described above.

As shown in FIG. 11, the console 8 comprises a computer, a CPU 74 which controls the operation of the entire apparatus, a ROM 75 in which various programs including control programs are stored in advance, a RAM 76 temporarily storing various data, and various data , And are connected to one another via a bus. The communication I / F 78 and the wireless communication unit 79 are connected to the bus, the display 80 is connected via the display driver 81, and the operation panel 82 is connected via the operation input detection unit 83.

The communication I / F 78 is connected to the communication I / F 87 of the radiation generation device 6 via the connection terminal 78 a, the communication cable 86, and the connection terminal 87 a of the radiation generation device 6. The console 8 (the CPU 74 thereof) instructs the radiation generating apparatus 6 to turn on and off the light source 14 through the communication I / F 87, and transmits and receives various information such as the irradiation condition and the collimator driving amount to the radiation generating apparatus 6. The wireless communication unit 79 has a function of performing wireless communication with the wireless communication unit 68 of the radiation image capturing apparatus 7. The console 8 (the CPU 74 thereof) instructs the radiation image capturing apparatus 7 to perform area discrimination processing by the wireless communication unit 79, and transmits and receives various information such as image data and area discrimination results.

The display driver 81 generates and outputs a signal for displaying various information to the display 80, and (the CPU 74 of the console 8) causes the display 80 to display an operation menu, a radiographic image taken, etc. via the display driver 81. . The operation panel 82 is configured to include a plurality of keys, and various information and operation instructions are input. The operation input detection unit 83 detects an operation on the operation panel 82 and notifies the CPU 74 of the detection result.

The console 8 is provided with a collimator drive amount determination unit 91 that determines the drive amount of the collimator 11 based on the area determination result transmitted from the radiation imaging device 7. The collimator drive amount determination unit 91 includes an irradiation range determination unit 91a. As shown in FIG. 12A and FIG. 12B, the irradiation range determination unit 91a determines the irradiation range so that the blank area decreases, based on the result of the area determination. The irradiation range determination unit 91a has table data storing the correspondence between the position in the light receiving surface 32a of the photosensor array 32 and the position of each pixel in the image detection surface 39a of the FPD 31. The irradiation range is determined based on the and table data.

Since the irradiation opening 11a of the collimator 11 is rectangular, as shown in FIGS. 12A and 12B, the irradiation areas A1 and A2 on the light receiving surface 32a of the photosensor array 32 and the image detection surface 39a of the FPD 31 are also rectangular. FIG. 12A is an example in which the irradiation range is determined such that the entire outline of the subject H fits, and FIG. 12B is an example in which the irradiation range is determined such that the irradiation range fits in the outline of the subject H.

In FIG. 12A, the irradiation range A1 is determined based on the result of area determination so as to be the smallest size among sizes in which the subject area corresponding to the entire outline of the subject H can be contained. Specifically, the irradiation range determination unit 91a is configured such that the outermost part of the outline of the subject H and the outline of the irradiation range A1 substantially match, that is, the longitudinal direction and the lateral direction of the subject H The width of the irradiation range A1 is determined in accordance with the respective maximum widths of. In this way, the area of the blank area B indicated by hatching in the irradiation range A1 can be reduced as compared to the case where the irradiation range is the entire image detection surface 39a. In addition, in the radiographic image to be captured, it is possible to visualize the entire image of the subject H.

However, since the shape of the subject H is not rectangular like the irradiation range A1, if the irradiation range A1 is determined so that the contour of the subject H is contained, the blank area B remains in the irradiation range A1. Specifically, in each of the array units 33a to 33j in which the outline of the subject H is located, a part is a missing area. It is better for the blank area to be smaller.

Therefore, when the region of interest is a part of the entire subject H, the irradiation range A2 may be determined so as to be within the contour of the subject H as shown in FIG. 12B. Thus, in the irradiation range A2, a region such as the blank region B in FIG. 12A can be substantially eliminated, so the area of the blank region can be further reduced compared to the irradiation range A1.

The irradiation range A2 is determined as follows. The irradiation range determination unit 91a specifies the contour position of the subject H from the coordinate information of the subject area included in the area determination result. Then, the size of the irradiation range A2 is determined so as to be within the contour of the subject H.

If the irradiation range A2 is determined, the contour portion of the subject H will extend out of the irradiation range A2, so there is a concern that the contour portion of the subject H will not be drawn in the radiographed image. However, as described above, each of the

shielding plates

17a, 17b, 18a and 18b of the collimator 11 has a wedge-shaped cross section (see FIG. 2), and the thickness thereof is thinner on the irradiation opening 11a side. A small amount of X-rays is transmitted near the opening 11a. Therefore, as in the area C surrounded by the alternate long and two short dashes line L larger than the irradiation range A2 outside the irradiation range A2, X-rays are slightly emitted in the vicinity of the outline of the object H that protrudes from the outside of the irradiation range A2. Ru. Therefore, in the radiation image, the subject H can be depicted to such an extent that the outline thereof can be determined.

In the area C, an area outside the contour of the subject H is a blank area where X rays are directly incident on the irradiation surface 22 without transmitting the subject H. However, since the X-rays entering the region C are transmitted through the

shielding plates

17a, 17b, 18a and 18b, their energy and dose are greatly attenuated. Therefore, in comparison with the X-ray incident to the region B in FIG. 12A, the influence on the dose of the patient or the characteristic deterioration of the CMOS sensor 48 is small.

Whether the irradiation range is determined as the irradiation range A1 or the irradiation range A2 can be selected, for example, by manual setting. The irradiation range determination unit 91a determines the irradiation range according to the selected method.

The collimator drive amount determination unit 91 determines the drive amount of the collimator 11 based on the irradiation range determined by the irradiation range determination unit 91 a. Since the X-ray irradiated from the X-ray tube 10 is a divergent beam having a spread angle, the irradiation range in the radiographic imaging device 7 depends on the imaging distance between the X-ray tube 10 and the radiographic imaging device 7 The size is larger than the size of the irradiation opening 11 a of the collimator 11. The collimator drive amount determination unit 91 determines the size of the irradiation opening 11 a of the collimator 11 based on the irradiation range A1 determined by the irradiation range determination unit 91 a and the imaging distance between the radiographic imaging device 7 and the X-ray tube 10. The amount of driving of the collimator 11 is determined so that the irradiation opening 11a has the size.

The collimator drive amount determination unit 91 transmits the determined collimator drive amount to the radiation generator 6 via the communication I /

Fs

78 and 87.

In FIG. 11, the radiation generator 6 communicates with the console 8 along with the X-ray tube 10, the collimator 11, the light source 14, and the collimator drive mechanism 19 to transmit and receive various information such as exposure conditions and the amount of collimator drive. / F 87 and a radiation source control unit 89. When receiving a command for area determination from the console 8, the radiation source control unit 89 controls the collimator drive mechanism 19 to maximize the irradiation opening 11 a of the collimator 11 and turn on the light source 14. Further, when receiving the collimator driving amount from the console 8, the radiation source control unit 89 controls the collimator driving mechanism 19 based on the collimator driving amount. The collimator drive mechanism 19 drives the collimator 11 to adjust the size of the irradiation opening 11 a. Further, the radiation source control unit 89 controls the X-ray tube 10 based on the irradiation condition (the irradiation condition includes the information of the tube voltage and the tube current) received from the console 8.

Next, the operation of the present embodiment will be described. When imaging a radiographic image using the radiographic imaging device 7, a radiographer (for example, a radiographer or the like) is a radiographic imaging device with the irradiation surface 22 facing upward between the subject H and the imaging table 7 is inserted, and the orientation of the subject H with respect to the irradiation surface 22 and the position of the subject H in the irradiation surface 22 are adjusted. For example, as shown in FIG. 9, when the subject H is a hand, it is placed directly on the center of the top of the radiation image capturing apparatus with the palm directed to the irradiation surface 22. Thereby, the positioning of the subject H with respect to the image detection surface 39a is completed.

When the positioning is completed, the photographer operates the operation panel 82 to instruct area determination. Thus, the console 8 transmits a region discrimination command to the radiation generation device 6 and the radiation imaging device 7. The radiation source control unit 89 of the radiation generating apparatus 6 receiving the command from the console 8 controls the collimator drive mechanism 19 to maximize the irradiation opening 11 a of the collimator 11 and turn on the light source 14. The visible light emitted from the light source is reflected by the reflection mirror 13 and emitted to the subject H and the radiation imaging device 7. Since the irradiation opening 11 a is at the maximum, the irradiation range of visible light is the entire surface of the irradiation surface 22. Since the visible light does not pass through the subject H, only the visible light irradiated around the subject H on the illumination surface 22 passes through the illumination surface 22.

The control unit 67 of the radiation image capturing apparatus 7 having received an instruction for area determination from the console 8 operates the area determination unit 70 to execute area determination processing of the subject area and the missing area.

In the photosensor array 32, visible light does not enter the subject region, but visible light enters only the clear region around the subject H. The photosensor array 32 outputs a light amount signal corresponding to the amount of light received by each photosensor 33a to the circuit unit 70a. The circuit unit 70a compares the light amount signal of each photosensor 33a with the threshold, and determines that the coordinates of the photosensor 33a whose light amount signal is equal to or greater than the threshold is a missing area, and the light amount signal of the photosensor 33a is less than the threshold The coordinates are determined to be the subject area. The circuit unit 70a generates coordinate information in which the information indicating each of the determined areas is associated with the coordinates of the photosensor 33a, and outputs the generated information to the control unit 67 as an area determination result. The control unit 67 transmits the area determination result to the console 8 by the

wireless communication units

68 and 79.

In the console 8, the irradiation range determination unit 91 a determines the irradiation range as the irradiation range A1 illustrated in FIG. 12A or the irradiation range A2 illustrated in FIG. 12B according to the selected determination method.
As a result, compared to the case where the irradiation range is the entire image detection surface 39a, it is possible to reduce the X-ray clear area. Moreover, since the irradiation range is determined based on the highly accurate area discrimination result by the light source 14 and the area discrimination unit 70, it is more accurate than when the irradiation range is determined visually.

The collimator drive amount determination unit 91 determines a collimator drive amount based on the irradiation range determined by the irradiation range determination unit 91a. The determined amount of driving of the collimator is transmitted to the radiation generator 6 by the communication I /

Fs

78 and 87. The collimator drive mechanism 19 drives the collimator 11 based on the received collimator drive amount to turn off the light source 14. Thereby, the size of the irradiation opening 11a is adjusted so that it may become the irradiation range which the irradiation range determination part 91a determined.

After the adjustment of the collimator 11 is completed, the photographer operates the operation panel 82 and instructs start of imaging. Thereby, the console 8 transmits an instruction signal instructing the start of exposure to the radiation generator 6, and the radiation generator 6 causes the X-ray tube 10 to emit radiation. The X-rays emitted from the X-ray tube 10 are narrowed by passing through the irradiation opening 11 a of the collimator 11. Then, the X-rays that have passed through the irradiation opening 11 a of the collimator 11 are transmitted through the radiation source filter 12 to remove low energy components, and are irradiated onto the subject H on the irradiation surface 22 of the radiographic imaging device 7. Since the irradiation range of X-rays is determined so that the blank area is reduced, it is possible to reduce X-rays directly incident on the irradiation surface 22 of the radiation imaging device 7 without transmitting the subject H. it can. Since wasteful X-rays are reduced, the amount of exposure to the patient is reduced, and it also contributes to the reduction of characteristic deterioration of the CMOS sensor 48 constituting the sensor panel 39.

The X-rays transmitted through the subject H are transmitted through the top 24, the photosensor array 32 and the sensor panel 39 of the radiation imaging device 7, and are applied to the scintillator 40.

Most of the X-rays irradiated to the scintillator 40 are converted into light in the vicinity of the X-ray incident surface of the scintillator 40, that is, in the vicinity of the sensor panel 39, and travel toward the sensor panel 39. Further, among the light generated by the scintillator 40, the light directed to the support substrate 41 side is reflected by the reflective layer of the support substrate 41 and is directed to the sensor panel 39. Thereby, the resolution of the radiographic image obtained by imaging | photography is high, and the light reception amount of the sensor panel 39 increases, As a result, the sensitivity of the radiographic imaging apparatus 7 improves. In addition, light directed from the scintillator 40 to the sensor panel 39 is guided by the columnar crystals of the scintillator 40 made of CsI: Tl, so image blurring is suppressed.

The visible light converted by the scintillator 40 passes through the second electrode 53 and enters the photoelectric conversion layer 52, where it is converted into a signal charge. After completion of the exposure, the signal charge generated in the photoelectric conversion layer 52 is converted into a voltage signal by the signal output circuit 57, and this voltage signal is sequentially output from each pixel. The output voltage signal is converted into image data by the signal processing unit 65 and stored in the image memory 66. The CPU 67 a transmits the image data stored in the image memory 66 to the console 8 by the wireless communication unit 68. The CPU 74 of the console 8 stores the image data received from the radiation imaging device 7 in the HDD 77 via the RAM 76. Further, the CPU 74 causes the display 80 to display a radiation image composed of image data stored in the HDD 77 via the display driver 81.

Even when the irradiation range is determined to be within the contour of the subject H as in the irradiation range A2 shown in FIG. 12B, each of the

shield plates

17a, 17b, 18a, 18b of the collimator 11 has a wedge-shaped cross section. Because of this, X-rays are slightly transmitted through the thin portions of the

shielding plates

17a, 17b and 18a, 18b on the irradiation opening 11a side. Therefore, since the X-ray is slightly irradiated to the outline portion of the subject H, the outline of the subject H is depicted on the radiation image. Thus, even if X-rays are irradiated near the contour, the amount of X-rays is small, so the effect on the increase in the exposure dose of the patient is small.

As in the case of follow-up observation, when the radiation image taken by the radiation imaging device 7 is compared with the radiation image in the past, the one where the outline of the subject H is reflected in the radiation image identifies the imaging site. Cheap. According to the present embodiment, it is possible to reduce the exposure dose of the patient without impairing the convenience in comparison with the past radiation image.

In the present embodiment, the subject area and the blank area are discriminated by using the light source 14 for emitting visible light and the area discrimination unit 70 having the photosensor array 32. Compared to the prior art (Japanese Patent Application Laid-Open No. 05-042135) for determining the subject area and the blank area, the amount of exposure to the patient can be reduced because there is no pre-imaging using X-rays.

In addition, a conventional technique for imaging an object positioned on a radiographic imaging device using an optical camera that uses indoor lighting or natural light, and performing image analysis on the captured image (Japanese Patent Laid-Open No. 2009-082169) This embodiment can have a simple configuration as compared with (US Patent Publication No. US 2009/0086885.) In the case of this prior art, it is possible to use indoor illumination or reflected light that is reflected by natural light from an object. The difference between light and dark areas between the subject area and the blank area in the photographed image is difficult: For example, the color of the clothes of the subject is white, and the color of the irradiation surface of the radiographic imaging device is white as well. In this case, there is almost no difference in brightness between the subject area and the blank area, so pattern recognition or contour extraction is necessary to distinguish between the subject area and the blank area. Image analysis is required.

On the other hand, in the present embodiment, the visible light is emitted from the light source 14 to the subject positioned on the radiation imaging apparatus, and the light is incident on the periphery of the subject by the photosensor array 32 disposed behind the subject. Visible light is detected. Since visible light does not pass through the subject, the photosensor array 32 can output a light amount signal having a large contrast between the subject area and a clear area around the subject. If there is a difference between light and dark in the two regions, it is possible to determine the two by comparing the magnitude with one threshold. Therefore, in the present embodiment, there is no need to perform image analysis such as pattern recognition and contour extraction, so the configuration can be simplified.

Further, as the difference in brightness between the two areas to be determined is larger, the determination of the two areas is easier, and the determination accuracy of the two areas is also improved. In the present invention, since the light source 14 is used, the contrast can also be adjusted by a simple method of adjusting the light emission amount of the light source 14. Furthermore, since the detection light emitted from the light source 14 is visible light, there is no concern that the patient is exposed.

Further, the FPD 31 of this embodiment is an ISS type in which the sensor panel 48 and the scintillator 40 are arranged in order from the irradiation surface 22 side, but in the ISS type, the scintillator 40 and sensor panel 39 are arranged in order from the irradiation surface side There is a great need to reduce the blank area more than the PSS type. That is, in the case of the PSS type, X-rays transmitted without being converted into light by the scintillator 40 enter the sensor panel 39, while in the case of the ISS type, before entering the scintillator 40. X-rays enter the sensor panel 39. Therefore, even in the same blank area, the energy and dose of X-rays incident on the sensor panel 39 are larger in the ISS type, and the damage of the blank area to the sensor panel 39 is larger in the ISS type. It is from. Therefore, the present invention is highly useful for ISS type radiation imaging apparatus.

Further, as described above, in the CMOS sensor made of single crystal Si, characteristic deterioration such as a change in threshold voltage Vth of the MOS transistor and an increase in dark current occurs due to the X-ray irradiation. This is because the interface charge increases in the MOS structure using single crystal Si. In the CMOS sensor, the influence of the characteristic deterioration due to such X-ray irradiation is large as compared with the TFT panel. Therefore, the present invention is particularly effective for a radiation imaging apparatus using a CMOS sensor.

Further, among the X-rays irradiated to the radiation imaging apparatus, it is considered that the low energy component of the X-rays not used for the radiation imaging influences the characteristic deterioration of the CMOS sensor. In other words, the high energy component of the X-ray passes through the CMOS sensor, but the low energy component of the X-ray is absorbed by the CMOS sensor because there is not enough energy to pass through the CMOS sensor. It is because there is a possibility of increasing the charge. In the present embodiment, since the low energy component of the X-ray is cut by the radiation source filter 12, characteristic deterioration of the CMOS sensor can be further suppressed.

Further, in the ISS type FPD 31 using a TFT panel, X-ray absorption of alkali-free glass used as a substrate of the TFT panel is large, and therefore, it is difficult to use for mammography with a low tube voltage. On the other hand, application to mammography has been expected because the single crystal Si substrate of the CMOS sensor has low absorption of X-rays. As in the present embodiment, by reducing the area where light passes through, characteristic deterioration due to X-ray irradiation in the CMOS sensor can also be suppressed, so it becomes possible to use the CMOS sensor as an ISS type FPD, and application to mammography is easy Become.

Further, in the present example, as a method of determining the irradiation range, as in the irradiation range A1 shown in FIG. 12A, a method of placing the entire subject H within the irradiation range A1 and a subject H as in the irradiation range A2 shown in FIG. Although two methods of the method of putting irradiation range A2 within the outline of are illustrated, it may be made to be able to determine in the middle size of irradiation range A1 and irradiation range A2 like irradiation range A3 shown to FIG. 12C.

In FIGS. 12A and 12B, the subject H is simplified and shown as an ellipse, but the actual subject H is not a simple outline of the actual subject H as shown in FIG. 12C, and the subject H is a hand. In the case, there is also a gap between the fingers. In the case of such a subject H, it is often difficult to fit the entire irradiation range within the contour of the subject H, including the region of interest, as in the case of the irradiation range A2 shown in FIG. 12B. Therefore, as in the irradiation range A3 shown in FIG. 12C, the irradiation range may be determined such that a part of the outline of the subject H is contained and a part thereof protrudes.

In this case, the irradiation range determination unit 91a specifies each array unit 33 in which the outline of the subject H is located and each blank area in each array unit 33 based on the area determination result.

The irradiation range determination unit 91a counts the number of pixels of the sensor panel 39 corresponding to the blank area of each array unit 33, that is, the number of blank pixels. Then, the irradiation range A3 is determined so that the number of missing pixels does not exceed a preset allowable value. That is, the irradiation range A3 is determined so that the number of missing pixels located in the region B in the irradiation range A3 does not exceed the allowable value. By doing this, it is possible to determine the irradiation range A3 that does not disturb the observation of the region of interest while reducing the blank region.

The allowable value of the number of missing pixels is stored in advance in, for example, the ROM 75 or the HDD 77, and can be changed by setting. The allowable value may be changed according to the imaging site such as the head, chest, abdomen, and hand. For example, in the case of relatively simple shapes such as the head, chest, and abdomen, the tolerance is low (relatively low pixel count is relatively low), and in the case of a complicated shape such as a hand, the tolerance is high ( Relatively large number of missing pixels). In this case, table data in which the relationship between the imaging region and the tolerance value is recorded is stored in the ROM 75, the HDD 77, and the like. Then, the irradiation range determination unit 91a determines the irradiation range with reference to the table data according to the selected imaging region.

Further, the tolerance value is not uniform for the X-ray irradiation dose, and the tolerance value may be changed according to the irradiation dose. For example, when the irradiation dose is large, the tolerance is low, and when the irradiation dose is small, the tolerance is high. In this case, the ROM 75 or the HDD 77 stores, as table data, an allowable value of the number of missing pixels corresponding to the X-ray irradiation dose. For example, when the irradiation dose is large, since the characteristic deterioration of the CMOS sensor 48 is large, the allowable value table of the number of missing pixels is set such that the number of missing pixels becomes “0”.

When the number of missing pixels is set to “0”, the irradiation range is determined as an irradiation range A2 shown in FIG. 12B. As described above, if the cross sections of the

shielding plates

17a and 17b and 18a and 18b of the collimator 11 have a wedge shape, X-rays are made in a range slightly larger than the irradiation range A2 as in the area C shown in FIG. Since it is slightly transmitted, it is possible to roughly determine the contour of the object H.

Second Embodiment
In the above embodiment, the collimator 11 having a wedge-shaped cross section is used. However, as in a radiation generating apparatus 100 shown in FIG. 13, a normal first collimator 101 having a constant thickness in the cross section and a wedge-shaped cross section The second collimator 102 may be used. In this case, the first collimator 101 is formed of a lead plate or the like to limit the X-ray irradiation range as usual. The second collimator 102 is made of aluminum or the like excellent in absorption of the low energy component of the X-ray, and used for absorption of the low energy component in the blank area. Then, when determining the irradiation range of the X-ray, the second collimator 102 is fully opened, and the opening degree of the first collimator 101 and the second collimator 102 is controlled according to the determined irradiation range, Adjust the size of each irradiation opening.

For example, the irradiation aperture of the first collimator 101 is adjusted such that the irradiation range has a size indicated by a two-dot chain line L in FIG. 12B, and the opening range of the second collimator 102 is irradiated with a size of the irradiation range It is adjusted to be in the range A2. Also in this case, the same effect as that of the first embodiment can be obtained.

Third Embodiment
Further, in the above embodiment, although the size of the array unit 33 constituting the photosensor array 32 is made uniform, the size of the light receiving surface of the array unit may be different. For example, as in the case of the photosensor array 105 shown in FIG. 14, the area of the light receiving surface of the array unit 106 disposed in the central portion is increased over the entire light receiving surface 105 a, and the light receiving surface of the array unit 107 disposed in the peripheral portion The area of may be reduced. Since the light receiving surface 105 a of the photosensor panel 105 corresponds to the image detection surface 39 a of the FPD 31, the array unit 106 is located at the center of the image detection surface 39 a and the array unit 107 is located at the periphery of the image detection surface 39 a. Since the peripheral part often corresponds to the contour part of the subject, the resolution of that part is increased by reducing the area of the array unit 107 in the peripheral part, so the boundary between the subject area and the blank area can be made more precisely. It can be determined.

Fourth Embodiment
Further, as shown in FIGS. 15A and 15B, the radiation generating apparatus 110 is provided with a swing mechanism 112 for rotating the X-ray tube 10 about the focal point of the X-ray tube 10 and used together with the control of the opening degree of the collimator 11. May be The swinging angle changes the X-ray irradiation angle. When a lesion G to be a region of interest is present at the end of the subject H, as shown in FIG. 15A, the lesion G faces the center of the irradiation path of the X-ray tube 10 when no swing is performed. Position the subject H to do so. Then, the opening degree of the collimator 11 is controlled to narrow the irradiation range so that the blank area is reduced. As described above, the collimator 11 interlocks the pair of shielding plates to adjust the width of the irradiation opening 11 a so that the center of the irradiation opening 11 a does not change. Therefore, when the irradiation opening 11a is narrowed so as to reduce the blank area in the irradiation range, the lesion G is located at the end of the subject H, so the irradiation range becomes very narrow.

In such a case, as shown in FIG. 15B, if the irradiation angle of X-rays is changed by swinging, the irradiation opening 11a is narrowed so that the blank area near the end of the object H is reduced. Even in this case, the irradiation range can be expanded as compared with the example shown in FIG. 15A. If the irradiation range is expanded, the outline of the subject H around the lesion G can be imprinted in the radiation image. It is easier to compare the radiation image taken in the past if the contour is known.

Furthermore, even if the subject H is intended to be positioned at the center of the radiation generator 110, a slight deviation may occur. Also in such a case, by slightly swinging the radiation generator 110, it is easy to reduce the blank area within the irradiation range. The determination as to whether or not to cause the radiation generator 110 to swing is when there is a significant difference in the size of the left and right blank areas in the width direction of the irradiation range. Specifically, as shown in FIG. 15A, when the irradiation range is expanded without swinging, many missing areas exist on the right side of the subject H with the lesion G, and on the left side There is a subject H, and there is only a few blank areas. If the difference is large, as shown in FIG. 15B, the swing is performed so that the blank area on the right side decreases.

Fifth Embodiment
In the fourth embodiment, the radiation generating apparatus is provided with a swing mechanism, but as shown in FIG. 16, the pair of

shield plates

115 a and 115 b of the collimator 115 can be individually driven, and the central axis CX of the X-ray bundle On the other hand, the collimator openings a and b may be different on the left and right. According to this, the same effect as the swing mechanism can be obtained.

Sixth Embodiment
In the first embodiment, the light transmitted through the top plate 24 is prevented from entering the FPD 31 by the light shielding property of the photosensor array 32 itself or the light shielding layer disposed between the photosensor array 32 and the FPD 31. A light shielding plate may be used to provide the top plate with a light shielding property as needed.

The top plate 120 of the present embodiment shown in FIG. 17 is made of plastic or the like having high visible light and X-ray permeability, as in the first embodiment. The top plate 120 is provided with a slit 122 in which the light shielding plate 121 can be inserted and removed from the side. The light shielding plate 121 has high light shielding properties for visible light, and has high transparency for X-rays.

When performing region determination by the radiation image capturing apparatus using the top 120 of the present embodiment, as shown in FIG. 17A, the light shielding plate 121 can be pulled out from the top 120 and visible light can be transmitted through the top 120 Make it Further, when radiation imaging is performed, as shown in FIG. 17B, the light shielding plate 121 is inserted into the top plate 120 so that visible light does not enter the FPD 31.

By using such a top plate 120, it is possible to prevent visible light from being incident on the FPD 31. In addition, although the form which provides the light-shielding plate 121 in the slit 122 of the top plate 120 so that insertion and removal is possible is illustrated, another form may be sufficient. For example, one end of the light shielding plate 121 is rotatably attached to the top plate 120 or the housing body via a hinge, and between the light shielding position for shielding visible light incident on the FPD 31 and the retracted position for retreating from the light shielding position. You may make it displace.

Further, if the top plate 120 is provided, the sensor panel 39 of the FPD 31 can also be used as a light detection unit. Similar to the photosensor array 32, the sensor panel 39 also has a function of detecting visible light and outputting a light amount signal. If the top plate 120 is provided, the light shielding plate 121 is removed from the slit 122 to distinguish visible light from the sensor panel 39 in the area determination, and the light shielding plate 121 is slit in the radiation imaging. The light blocking state of the sensor panel 39 can be switched such that the visible light is blocked by inserting the light source into the light source 122. As described above, when the sensor panel 39 is also used as a light detection unit, the photosensor array 32 can be omitted, so that the configuration can be simplified as compared with the prior art.

Seventh Embodiment
In each of the above-described embodiments, the radiation image capturing apparatus itself is provided with the area determining function. However, the radiation image capturing apparatus may be provided with the area determining function. 18A and 18B, the radiation imaging system 130 of the present embodiment includes a radiation imaging device 131 and an attachment 132 which can be detachably attached to the radiation imaging device 131. The radiographic imaging device 131 is not provided with a configuration for realizing the area discrimination function such as the transparent top plate 24 and the photosensor array 32 as the radiographic imaging device 7 of the first embodiment is not provided. It is a conventional radiographic imaging device 131 in which a housing including a top plate has a light shielding property.

The attachment 132 has a jacket shape having a slit 133 into which the radiographic imaging device 131 can be inserted and removed, and covers the outer peripheral surface of the radiographic imaging device 131 when the radiographic imaging device 131 is inserted into the slit 133. The attachment 132 is formed of, for example, a transparent plastic that is highly transparent to visible light and X-rays.

The attachment 132 is provided with a photosensor array 134, a circuit unit (not shown) similar to the circuit unit 70a, a wire 135, and a connector 136. The photosensor array 134 is the same as the photosensor array 32 according to the first embodiment, and is disposed at a position facing the upper surface of the radiographic imaging device 131 when the radiographic imaging device 131 is inserted into the slit 133.

The wiring 135 and the connector 136 electrically connect the photosensor array 134 and the radiographic imaging device 131 when the radiographic imaging device 131 is inserted into the slit 133.

When the area is determined by the radiation imaging system 130 according to the present embodiment, as shown in FIG. 18A, the radiation imaging apparatus 131 is inserted into the slit 133 to insert the conventional radiation imaging apparatus 131 into the slit 133. Attachment 132 is attached to the state shown in FIG. 18B. In this state, the subject H is positioned to emit visible light from the light source 14. In the passthrough region around the subject H, visible light passes through the attachment 132 and enters the photosensor array 134. In the radiographic imaging device 131, since the casing has a light shielding property, visible light for area determination does not enter the built-in FPD 31.

The photosensor array 134 outputs a light amount signal to the circuit unit. The circuit unit transmits the area determination result to the radiation image capturing apparatus 131 via the wiring 135 and the connector 136. The radiographic imaging device 131 transmits the area discrimination result to the console 8. As in the first embodiment, the irradiation range and the collimator driving amount are determined by the collimator driving amount determination unit 91 of the console 8, and the collimator 11 of the radiation generating device 6 is adjusted so as to reduce the blank area.

Radiography is performed after adjustment of the collimator 11 is completed. Since the attachment 132 is made of a material having high X-ray transparency, as shown in FIG. 18B, radiation imaging can be performed with the attachment 132 attached.

In this example, although the region discrimination result is transmitted to the console 8 via the radiation imaging device 131, it may be directly transmitted to the console 8 without passing through the radiation imaging device 131. In this case, the attachment 132 is provided with a wired or wireless transmission unit for transmitting to the console 8 instead of the wire 135 and the connector 136. As in the present embodiment, the photosensor array may be provided separately from the radiation imaging device.

In the above embodiments, the region determination is always performed at the time of X-ray imaging, but may be performed only when the X-ray dose is a predetermined dose or more. The purpose of reducing the blank area is to reduce the exposure dose of the patient and the characteristic deterioration of the CMOS sensor, so the higher the dose, the more problematic.

In each of the above embodiments, the photosensor array 32 is used as the light detection unit of the present invention, and the area determination is performed by the photosensor array 32. However, as in the sixth embodiment, the light shielding plate 121 or the like When using a structure capable of switching between a state in which visible light from the outside can be incident on the FPD 31 and a light blocking state in which the FPD 31 is blocked, false detection by visible light from the outside can be prevented during radiation imaging. The sensor panel 39 may also be used as a light detection unit for area determination. When the sensor panel 39 is used for area discrimination, it is preferable that the light receiving sensitivity of the sensor panel 39 be high.

In order to increase the light reception sensitivity of the sensor panel 39, it is conceivable to reduce the thickness of the substrate or the like to suppress the light loss and to increase the amount of light incident on the photoelectric conversion portion. As a method of reducing the thickness of the substrate, a flexible CMOS sensor in which a CMOS sensor is formed on a flexible plastic substrate such as a transparent plastic film is used as the sensor panel 39 instead of the CMOS sensor using a single crystal semiconductor substrate. There is a way.

An organic thin film transistor can be used as a transistor of the flexible CMOS sensor. For organic thin film transistors, see “Tsuyoshi Sekitani,” “Flexible organic transistors and circuits with extreme bending.
As described in detail in “Stability”, Nature Materials 9, November 7, 2010, pp. 1015-1022 ”, the detailed description is omitted.

Alternatively, a structure in which a photodiode and a transistor formed of single crystal Si are provided over a plastic substrate may be used. The arrangement of photodiodes and transistors on a plastic substrate can be performed, for example, by spraying a device block having a size of several tens of microns in a solution and placing it at a required position on any substrate. The Assembly (FSA) method can be used. Regarding the FSA method, "Koichi Maezawa," Resonant tunnel device block fabrication technology for Fluidic Self-Assembly ", Technical Report of IEICE, ED, Electronic Device, The Institute of Electronics, Information and Communication Engineers, 2008 6 Since it is described in detail on May 6, 108, 87, pp. 67-71, the detailed description is omitted.

In each of the above embodiments, an example using visible light as detection light used for area discrimination has been described, but it may be other than visible light and may be electromagnetic waves other than radiation that does not expose a subject such as infrared light or ultraviolet light. It is also good. As a matter of course, as the light detection unit, one having sensitivity characteristics according to the wavelength of detection light to be used is used, but instead of matching the sensitivity characteristics of the photoelectric conversion unit itself to the wavelength of detection light A member may be used.

For example, the scintillator 40 made of CsI emits a slight amount of light even when irradiated with ultraviolet light, in addition to radiation. Since the emitted light has a wavelength in the visible light range, this light may be detected by the sensor panel 39 or the photosensor array 32 and used for area determination.

In each of the above embodiments, the FPD is described using an example in which a CMOS sensor is used, but the present invention may use a CCD image sensor using a single crystal semiconductor substrate as in the CMOS sensor. In addition, a TFT type FPD using a glass substrate may be used, and either an indirect conversion type or a direct conversion type may be used. Also, although the ISS type FPD has been described as an example, the present invention is also applicable to the PSS type FPD.

However, as described above, unlike FPDs that use a single crystal semiconductor substrate such as a CMOS sensor, unlike TFTs that use a glass substrate, there is a concern that the characteristics will deteriorate due to X-ray irradiation. Sex is high. Further, in the case of the ISS type, since the X-rays are directly incident on the substrate without passing through the scintillator, the concern about the characteristic deterioration becomes more problematic. Therefore, the present invention is particularly effective for ISS type FPDs using a single crystal semiconductor substrate.

In each of the above-described embodiments, in the case where the light detection unit is provided separately from the FPD, an example using a photosensor array as the light detection unit has been described. However, the light receiving surface in which elements having a light detection function are arranged in a matrix Any sensor may be used, and a CMOS sensor or a CCD sensor may be used instead of the photosensor array. In the present invention, the detection light is emitted to the subject from the light source 14 and the detection light incident on the periphery of the subject is received, so that a light quantity signal with a large difference in brightness between the subject area and the clear area is obtained. Therefore, since the area determination can be performed by comparing the light amount signal with the threshold, image analysis such as pattern recognition and contour extraction is unnecessary. Therefore, even if a CMOS sensor or a CCD sensor is used as the light detection unit, the image analysis function is not necessary, and the effect of the present invention of simplifying the configuration can be obtained. Of course, it is preferable to use a photosensor array as the light detection unit because the photosensor array is simpler and cheaper than the CMOS sensor or the CCD sensor.

Although the light source 14 is provided in the radiation generating apparatus in each of the above embodiments, the light source 14 may not be provided in the radiation generating apparatus. Since visible light from the light source 14 is irradiated on the entire light receiving surface of the light detection unit including the subject, it is not necessary to narrow the irradiation range with a collimator. Therefore, there is no problem even if the light source 14 is provided separately from the radiation generator.

In each of the above embodiments, the irradiation range is determined based on the area discrimination result, and the collimator is automatically controlled based on the determined irradiation range. However, the control of the collimator may not be performed automatically. For example, the determination result of the area and the determined irradiation range may be displayed on the console display, and the opening degree of the collimator may be manually adjusted while confirming the display on the display. Of course, automatic control of the collimator is preferable because it reduces the time and effort. In addition, even when the control of the collimator is automatically performed, it is preferable to be able to perform fine adjustment manually.

In addition, although a portable example in which the FPD is incorporated in a cassette-sized casing has been described, the FPD can be incorporated in a standing or lying-down stationary imaging apparatus or a mammography apparatus. Also, although X-rays have been described as an example of radiation, the present invention may use radiation other than X-rays, such as γ-rays. In addition, the configuration of the radiation image capturing apparatus according to the present invention described in the above embodiment is merely an example, and it goes without saying that the configuration can be appropriately changed without departing from the scope of the present invention.

Claims

A radiation generator having a radiation source for irradiating the subject with radiation;
A radiation imaging apparatus having an image detection surface in which a plurality of pixels are arranged in a matrix, and having a radiation image detector that receives radiation transmitted through the subject and detects a radiation image of the subject;
A light source that emits detection light to the subject positioned with respect to the image detection surface;
A light detection unit that receives the detection light emitted from the light source and enters the periphery of the subject and outputs a light amount signal, and based on the light amount signal, the subject facing the subject in the image detection plane An area determination unit that determines an area and a missing area around the subject;
A radiation imaging system comprising:
The light detection section is a photo sensor array having a light receiving surface in which a plurality of photo sensors for outputting light quantity signals according to the light receiving amount of the detection light are arranged in a matrix. The radiographic imaging system as described in a term.
The radiation imaging system according to claim 1, wherein the photo sensor array is disposed closer to the light source than the radiation image detector.
The radiation imaging system according to claim 3, wherein the light receiving surface of the photo sensor array is disposed in parallel with the image detection surface of the radiation image detector.
The area determination unit determines the subject area and the blank area by comparing the light amount signals output from the plurality of photosensors with a preset threshold value. The radiation imaging system according to claim 2.
The radiation image capturing system according to claim 1, wherein the light source is provided in the radiation generating apparatus, and the detection light is emitted to the subject from the same direction as the radiation.
The radiation generating apparatus is composed of a plurality of shielding plates that define an irradiation opening that transmits the radiation, and the size of the irradiation opening is adjusted by the movement of the shielding plate, and the irradiation range in the image detection plane Have a collimator that limits
An irradiation range determination unit that determines the irradiation range so that the blank area is reduced based on the determination result of the area determination unit;
The radiation image photographing according to claim 1, further comprising: a collimator drive amount determination unit that determines a drive amount of the collimator so as to be the irradiation range determined by the irradiation range determination unit. system.
The irradiation range determination unit specifies the number of unfiltered pixels of the radiation image detector present in the unfiltered region based on the determination result of the region judging unit, and the number of unfiltered pixels is set in advance. The radiation imaging system according to claim 7, wherein the irradiation range is determined so as not to exceed an allowable value.
9. The radiation image capturing system according to claim 8, wherein the allowable value of the unfiltered number of pixels is set in accordance with an irradiation dose of radiation emitted from the radiation source.
The photosensor array is configured by tiling a plurality of array units each having a light receiving surface in which a plurality of photosensors are arrayed,
The size of the light receiving surface of the array unit arranged at the peripheral portion of the image detecting surface is smaller than that of the array unit arranged at the central portion of the image detecting surface. The radiographic imaging system as described in a term.
The collimator has at least a pair of the shielding plates that changes the width of the irradiation opening, and the cross-sectional shape of the shielding plate is a wedge shape whose thickness increases from the edge defining the irradiation opening toward the outside. The radiation imaging system according to claim 7, which is a shape.
The collimator is characterized by having a first collimator which limits an irradiation range of radiation, and a second collimator which absorbs an energy component whose energy is relatively low from radiation irradiated to the blank area. Range The radiographic imaging system of Claim 7.
The radiation imaging system according to claim 11, wherein the pair of shielding plates are movable independently of each other.
The radiation image detector is a scintillator that absorbs radiation and converts it into light, and a sensor that is disposed on the radiation irradiation side of the scintillator and in which a plurality of pixels that detect light converted by the scintillator are arranged in a matrix The radiation imaging system according to claim 1, further comprising a panel.
The radiation image capturing system according to claim 14, wherein the sensor panel is configured of a CMOS type image sensor.
The radiation imaging system according to claim 14, wherein the sensor panel is also used as the light detection unit.
The radiation image capturing apparatus is formed with an irradiation surface that transmits the radiation and the detection light, and a housing that accommodates the sensor panel;
A light shielding plate having transparency to the radiation, and a light shielding property to the detection light;
The light shielding plate is provided so as to be displaceable between a light shielding position for shielding the detection light incident on the sensor panel and a retracted position for retreating from the light shielding position.
The radiographic imaging system of Claim 16 characterized by the above-mentioned.
An attachment removably attachable to the radiation imaging apparatus;
The radiation imaging system according to claim 1, wherein the light detection unit is provided in the attachment.
A radiation image detector having an image detection surface in which a plurality of pixels are arranged in a matrix, receiving radiation transmitted from the radiation source and transmitted through the subject to detect a radiation image of the subject;
A light detection unit for receiving the detection light emitted from the light source and incident on the periphery of the subject in a state in which the subject is positioned with respect to the image detection surface, and outputting a light amount signal; An area determination unit that determines, based on the subject area where the subject faces in the image detection plane, and the blank area around the subject;
A radiation image capturing apparatus comprising:
Irradiating the subject with detection light while the subject is positioned with respect to the image detection surface of the radiation image detector;
Receiving the detection light incident on the periphery of the subject and outputting a light amount signal;
Determining a subject area facing the subject in the image detection plane based on the light amount signal and a blank area around the subject;
A radiation image capturing method comprising: