WO2012165156A1 - Radiation-image-capturing device - Google Patents

Radiation-image-capturing device Download PDF

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Publication number
WO2012165156A1
WO2012165156A1 PCT/JP2012/062624 JP2012062624W WO2012165156A1 WO 2012165156 A1 WO2012165156 A1 WO 2012165156A1 JP 2012062624 W JP2012062624 W JP 2012062624W WO 2012165156 A1 WO2012165156 A1 WO 2012165156A1
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WO
WIPO (PCT)
Prior art keywords
radiation
absorbing
scintillator
image
light
Prior art date
Application number
PCT/JP2012/062624
Other languages
French (fr)
Japanese (ja)
Inventor
白水　豪
中津川　晴康
俊孝 阿賀野
美広 岡田
岩切　直人
Original Assignee
富士フイルム株式会社
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Publication of WO2012165156A1 publication Critical patent/WO2012165156A1/en

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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/16—Measuring radiation intensity
    • G01T1/20—Measuring radiation intensity with scintillation detectors
    • G01T1/2018—Scintillation-photodiode combinations
    • G01T1/20188—Auxiliary details, e.g. casings or cooling
    • G01T1/20189—Damping or insulation against damage, e.g. caused by heat or pressure
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/16—Measuring radiation intensity
    • G01T1/20—Measuring radiation intensity with scintillation detectors
    • G01T1/2018—Scintillation-photodiode combinations
    • G01T1/20188—Auxiliary details, e.g. casings or cooling
    • G01T1/2019—Shielding against direct hits
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/42—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment with arrangements for detecting radiation specially adapted for radiation diagnosis
    • A61B6/4283—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment with arrangements for detecting radiation specially adapted for radiation diagnosis characterised by a detector unit being housed in a cassette
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/42—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment with arrangements for detecting radiation specially adapted for radiation diagnosis
    • A61B6/4291—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment with arrangements for detecting radiation specially adapted for radiation diagnosis the detector being combined with a grid or grating
    • G—PHYSICS
    • G03—PHOTOGRAPHY; CINEMATOGRAPHY; ANALOGOUS TECHNIQUES USING WAVES OTHER THAN OPTICAL WAVES; ELECTROGRAPHY; HOLOGRAPHY
    • G03B—APPARATUS OR ARRANGEMENTS FOR TAKING PHOTOGRAPHS OR FOR PROJECTING OR VIEWING THEM; APPARATUS OR ARRANGEMENTS EMPLOYING ANALOGOUS TECHNIQUES USING WAVES OTHER THAN OPTICAL WAVES; ACCESSORIES THEREFOR
    • G03B42/00—Obtaining records using waves other than optical waves; Visualisation of such records by using optical means
    • G03B42/02—Obtaining records using waves other than optical waves; Visualisation of such records by using optical means using X-rays
    • G03B42/04—Holders for X-ray films

Definitions

  • the present invention relates to a radiation imaging apparatus for capturing a radiation image.
  • the radiation imaging system for imaging a subject (a region to be imaged by a patient) using radiation (for example, X-rays) to perform image diagnosis.
  • the radiation imaging system has a radiation generating device for emitting radiation and a radiation imaging device for capturing a radiation image of an imaging region.
  • 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.
  • an FPD flat panel detector
  • the FPD detects a radiation image by accumulating signal charges for each pixel on the detection surface, and outputs this as digital image data.
  • 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 detection panel disposed opposite to the scintillator, and an electrical control circuit.
  • the detection panel has a detection surface in which a photoelectric conversion unit that generates a signal charge by photoelectric conversion is formed for each pixel, converts visible light from the scintillator into a signal charge, and stores the signal charge.
  • 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).
  • a-Si amorphous silicon
  • photoelectric conversion parts and MOS transistors are formed in a matrix on a single crystal semiconductor substrate of silicon (Si) by a semiconductor process.
  • 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.
  • the variation of characteristics for example, the threshold voltage of the MOS transistor and the like
  • 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, an FPD having a size of 17 inches, which is generally used for medical use, can be configured using four CMOS sensors.
  • the radiation image capturing apparatus is set under the patient A protective structure for preventing damage in the case is required, and the radiographic imaging device becomes large.
  • Japanese Patent Application Laid-Open No. 2005-249639 a scattered X-ray absorbing grid for absorbing X-rays scattered by a subject is disposed between the subject and the radiation imaging apparatus, and the outer periphery of the scattered X-ray absorbing grid is disposed What provided the absorption part which absorbs the X-ray irradiated to the blank area is disclosed. Further, Japanese Patent Application Laid-Open No. 2010-075553 discloses that a radiation generating apparatus is provided with a filter in which the amount of absorption of X-rays irradiated to the blank area is increased.
  • both the absorbing unit described in JP-A-2005-249639 and the filter described in JP-A-2010-075553 absorb X-ray components necessary for imaging, and the image quality of the radiation image is degraded.
  • neither JP-A-2005-249639 nor JP-A-2010-075553 describes the prevention of deterioration of the characteristics of the CMOS sensor and the prevention of damage.
  • An object of the present invention is to provide a radiation image capturing apparatus capable of preventing characteristic deterioration of a radiation detector due to radiation and preventing damage due to a load or an impact from a subject.
  • a radiation imaging device of the present invention is provided with a radiation detector and a radiation absorption part.
  • the radiation detector detects radiation transmitted through the subject and generates image data.
  • the radiation absorbing portion is disposed on the radiation incident side of the radiation detector, absorbs a part of the radiation incident on the radiation detector, and is deformable by the load of the subject.
  • the radiation detector has a single crystal semiconductor substrate.
  • the radiation absorbing portion has a bag and a liquid or gel radiation absorbing fluid sealed in the bag.
  • a liquid scintillator that absorbs radiation and converts it into light may be used.
  • a gelled water glass may be used as the radiation absorbing fluid.
  • the radiation absorbing fluid may also have a liquid or gel fluid and a radiation absorbing material dispersed in the fluid.
  • the radiation absorbing portion may have an elastic body which is deformed by a load from a subject, and a radiation absorbing material dispersed in the elastic body. Also, a plurality of radiation absorbing portions may be provided.
  • the radiation detector has 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 that detects light converted by the scintillator.
  • the radiation absorber preferably has a bag and a liquid scintillator enclosed in the bag.
  • a signal output circuit includes a plurality of pixels including a photoelectric conversion layer that converts light converted by a scintillator into electric charge and a signal output circuit that outputs a signal according to the electric charge converted by the photoelectric conversion layer as a sensor panel Any of those provided over a single crystal semiconductor substrate can be used. As such a sensor panel, one configured by a CMOS type image sensor is used.
  • the photoelectric conversion layer is preferably formed of amorphous silicon or quinacridone.
  • the wavelength range of light generated by the liquid scintillator be different from the wavelength range of light detectable by the sensor panel.
  • the radiation absorber is preferably larger in size than the radiation detector. It is preferable to include an image correction unit that corrects image data according to the area difference of the radiation absorption amount of the radiation absorption unit.
  • the radiation absorbing portion preferably absorbs the low energy component of the radiation.
  • the low energy component of the radiation absorbed by the radiation absorbing portion is an energy component of 1/2 or less of the energy distribution of the radiation.
  • the radiation absorbing unit absorbs a part of the radiation, it is possible to suppress the characteristic deterioration of the radiation detector. Further, the radiation detector can be reinforced by the radiation absorbing portion to prevent damage to the radiation detector.
  • the radiation absorbing portion is deformed by the load from the subject, the radiation absorption distribution is optimized according to the position, size, and range of the subject. As a result, the radiation transmitted through the subject does not deteriorate the image quality of the radiation image because the amount of radiation absorbed by the radiation absorbing portion is small, and the amount of radiation absorbed in the blank area directly irradiated with radiation is large, and the characteristics of the radiation detector deteriorate Is suppressed.
  • a radiation imaging system 5 comprises: a radiation generator 6 for emitting X-rays as radiation toward the imaging region of a subject (patient) H; a radiation imaging apparatus 7 for imaging a radiation image of the subject H; A console 8 for controlling the radiation generator 6 and the radiation imaging device 7 is provided.
  • the radiation generator 6 has a radiation source 6a and a source filter 6b.
  • the radiation source 6a has an X-ray tube 6c for emitting X-rays and a radiation field limiter (collimator) 6d for limiting the radiation field of X-rays emitted by the X-ray tube 6c.
  • the X-ray tube 6c has a cathode formed of a filament that emits thermoelectrons, and an anode (target) that emits thermorays which collide with the thermoelectrons emitted from the cathode.
  • the irradiation field limiter 6d is, for example, a plurality of lead plates for shielding X-rays arranged on each side of a square and an irradiation opening for transmitting X-rays formed at the center, and the position of the lead plate is moved By changing the size of the irradiation opening, the irradiation field is limited.
  • the radiation source filter 6b removes low energy components which are scattered when passing through the imaging region from the X-rays emitted from the radiation source 6a and cause deterioration of the radiation image.
  • a material having a property of absorbing only the low energy component of the X-ray is used.
  • 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 6c is, for example, approximately when the tube voltage of the X-ray tube 6c is 70 kV and the maximum energy of the X-rays emitted from the X-ray tube 6c is about 70 KeV 15 to 70 KeV.
  • about 1/2 or less (15 to 40 KeV) of the energy distribution of the X-ray is a low energy component, and 1/2 or more (40 to 70 KeV) is a high energy component.
  • the source filter 6b absorbs low energy components of 15 to 40 KeV.
  • the radiation image capturing apparatus 7 includes an FPD 19, a radiation absorbing unit 20, a light absorbing layer 21, an electric circuit unit 23, and a portable case 12.
  • the housing 12 has a top plate 13 and a flat box-shaped housing body 14, and accommodates the FPD 19 and the electric circuit unit 23.
  • the top plate 13 seals the opening 14 a at the top of the case body 14.
  • the upper surface of the top plate 13 is an irradiation surface 11 to which X-rays emitted from the radiation generator 6 are irradiated.
  • the top plate 13 is formed of carbon or the like having high X-ray permeability. Since carbon has high strength, it is suitable as a material of the top 13 to which the weight of the subject H is applied.
  • the housing body 14 is formed of ABS resin or the like.
  • the housing 12 is the same size (e.g., 17 inches by 17 inches) as a conventional radiation film cassette for recording a radiation image on a photosensitive material.
  • the radiographic imaging device 7 has the same portability as the radiographic film cassette, and is used in place of the radiographic film cassette.
  • the top plate 13 is provided with a display unit 16 configured of a plurality of LEDs.
  • the display unit 16 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 a remaining capacity of a battery.
  • the display unit 16 may be configured by a light emitting element other than an LED, a liquid crystal display, an organic EL display, or the like.
  • the display unit 16 may be provided in the housing body 14.
  • the light absorbing layer 21 and the radiation absorbing portion 20 are stacked in this order on the top plate 13.
  • the radiation absorbing unit 20 faces the top 13 via the light absorbing layer 21.
  • the radiation absorbing unit 20 absorbs a part of X-rays (for example, low energy components of X-rays) irradiated to the irradiation surface 11 and protects the FPD 19 from a load, an impact or the like applied to the top plate 13.
  • the light absorbing layer 21 absorbs the light generated by the radiation absorbing unit 20 by the X-ray irradiation.
  • the electric circuit unit 23 accommodates a signal processing unit 50, an image memory 51, a control unit 52, a wireless communication unit 53, a battery 54 and the like (all refer to FIG. 9).
  • the electric circuit unit 23 is disposed at one end side along the short direction inside the housing 12.
  • the FPD 19 operates by the power supplied from the battery 54.
  • a radiation shielding member (not shown) such as a lead plate is provided on the top plate 13 side of the electric circuit section 23 in order to prevent the electric circuit section 23 from being damaged by X-rays.
  • a light absorbing layer 21 is attached to the top surface of the top plate 13 with an adhesive (not shown).
  • the radiation absorbing portion 20 is attached onto the light absorbing layer 21 by an adhesive (not shown).
  • the light absorbing layer 21 and the radiation absorbing unit 20 cover the irradiation surface 11. Thereby, the radiation imaging device 7 can be thinned and the radiation detector 19 can be reinforced.
  • the FPD 19 is one in which a sensor panel 25 and a scintillator 26 are sequentially stacked from the top 13 side along the direction in which the radiation is irradiated.
  • a support substrate 27 supporting the scintillator 2 is provided on the lower surface of the scintillator 26.
  • a sealing agent 28 is provided on the outer periphery of the FPD 19 in order to protect the scintillator 26 from moisture and the like.
  • a drive circuit board 29 of the FPD 19 is disposed on the bottom of the housing 12. The drive circuit board 29 and the sensor panel 25 are electrically connected via the flexible cable 30.
  • the scintillator 26 transmits the subject H and is irradiated to the irradiation surface 11 of the housing 12, and transmits the top plate 13, the radiation absorbing unit 20, and the sensor panel 25 to absorb incident X-rays and generate visible light.
  • CsI: Tl cesium iodide to which thallium is added
  • CsI: Na cesium iodide to which sodium is added
  • GOS Gd 2 O 2 S: Tb
  • CsI: Tl is used as the scintillator 26.
  • the scintillator 26 is formed by vapor deposition of CsI: Tl on the support substrate 27.
  • the scintillator 26 has a columnar crystal structure, and has a plurality of columnar crystals (not shown) along the direction from the support substrate 27 toward the sensor panel 25.
  • the columnar crystal has a flat diameter substantially uniform along the longitudinal direction of the columnar crystal.
  • the light generated by the scintillator 26 propagates in the columnar crystal due to the light guide effect of the columnar crystal, and is emitted toward the sensor panel 25 from the tip of the columnar crystal.
  • the scintillator 26 in a columnar crystal structure, diffusion of visible light emitted from the scintillator 26 to the sensor panel 25 side is suppressed, so the sharpness of the radiation image captured by the radiation imaging device 7 Improve.
  • a reflective layer (not shown) is provided on the surface of the support substrate 27 on the scintillator 26 side.
  • the visible light emitted by the scintillator 26 and propagated to the support substrate 27 is reflected to the sensor panel 25 by the reflective layer, so the amount of light incident on the sensor panel 25 (detection efficiency of light emitted by the scintillator 26) is improved. Do.
  • the configuration in which the sensor panel 25 is disposed on the X-ray incident side of the scintillator 26 is referred to as an ISS (Irradiation Side Sampling) method.
  • a configuration in which the sensor panel is disposed on the opposite side to the X-ray incident side of the scintillator is called a PSS (Penetration Side Sampling) method.
  • PSS Pulsion Side Sampling
  • the sharpness and sensitivity of the radiation image are improved compared to the PSS method.
  • the sensor panel 25 is configured by four CMOS type image sensors (hereinafter referred to as CMOS sensors) 33.
  • CMOS sensors CMOS type image sensors
  • Each CMOS sensor 33 has a plurality of pixels 33a (see FIG. 7) arranged in a matrix.
  • Each CMOS sensor 33 has a rectangular shape with a side length of about 200 mm.
  • the four CMOS sensors 33 are arranged adjacent to each other vertically and horizontally, and form a square of approximately 17 inches on a side.
  • the CMOS sensor 33 has a configuration similar to that disclosed in U.S. Patent Publication 2009/0224162. Specifically, as shown in FIG. 5, the CMOS sensor 33 is constituted of a single crystal semiconductor substrate 34, an insulating layer 35, a first electrode 36, a photoelectric conversion layer 37, and a second electrode 38. There is.
  • the single crystal semiconductor substrate 34 is made of single crystal Si.
  • the insulating layer 35 is formed on the surface of the single crystal semiconductor substrate 34, and silicon oxide or the like is used for the insulating layer 35.
  • the first electrode 36 is individually formed for each pixel 33 a on the surface of the insulating layer 35.
  • the photoelectric conversion layer 37 is provided on the surface of each first electrode 36 in common to each pixel 33 a.
  • the second electrode 38 is provided on the surface of the photoelectric conversion layer 37 in common to each pixel 33 a.
  • the aforementioned scintillator 26 is bonded onto the surface of the second electrode 38 by an adhesive (not shown).
  • the second electrode 38 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 26 is incident on the photoelectric conversion layer 37.
  • ITO indium tin oxide
  • the second electrode 38 is provided in common to each pixel 33a, but may be provided individually for each pixel 33a.
  • the photoelectric conversion layer 37 generates a signal charge according to the amount of incident X-rays, in combination with the scintillator 26.
  • the photoelectric conversion layer 37 absorbs visible light generated by the scintillator 26, generates a signal charge according to the amount of absorbed light, and is formed of an organic or inorganic photoelectric conversion material.
  • the inorganic photoelectric conversion material include amorphous silicon (a-Si).
  • An example of the organic photoelectric conversion material is quinacridone.
  • the sensitivity of the organic photoelectric conversion material (OPC) made of quinacridone is closer to the wavelength range of visible light generated by CsI: Tl than by CsI: Na or single crystal Si (c-Si). .
  • CsI: Tl as the scintillator 26
  • a signal output circuit 41 is provided for each pixel 33a.
  • the signal output circuit 41 is formed of a CMOS circuit.
  • the signal output circuit 41 and the first electrode 36 are electrically connected by the contact wiring 42.
  • a bias voltage is applied to the second electrode 38 (see FIG. 7), and the signal charge generated by the photoelectric conversion layer 37 is collected by the first electrode 36 of each pixel 33 a.
  • the signal output circuit 41 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.
  • the signal output circuit 41 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 35 described above.
  • the output transistor T1 is connected to the first electrode 36, and a voltage corresponding to the signal charge collected by the first electrode 36 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 36 to the power supply wiring Vdd.
  • 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.
  • a control unit of the FPD 19 can be mixedly mounted on the single crystal semiconductor substrate 34.
  • the row selection line L1, the signal output line L2, and the reset line L3 are formed of a metal such as aluminum, so there is little deterioration due to X-rays, 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 may be degraded (change in threshold voltage or increase in dark current) due to X-rays. This is because in the MOS structure using single crystal Si, charges (hereinafter referred to as interface charge) are generated and accumulated at the interface between the single crystal semiconductor substrate 34 and the insulating layer 35 by absorption of X-rays.
  • interface charge charges
  • the radiation absorbing portion 20 absorbs a part of the X-rays transmitted through the top plate 13 (a low energy component affecting the characteristic deterioration of the MOS transistor).
  • the high energy component of the X-ray passes through the CMOS sensor 33, but the low energy component of the X-ray does not have enough energy to pass through the CMOS sensor 33, and is absorbed by the CMOS sensor 33.
  • the radiation absorbing portion 20 is composed of a bag 47 and a radiation absorbing fluid 48.
  • the bag body 47 has elasticity that can be deformed by the load of the subject H.
  • the radiation absorbing fluid 48 is enclosed in a bag 47.
  • the bag body 47 is formed in a bag shape, for example, of an elastic material such as rubber or plastic material having a thickness of about several mm to 1 cm.
  • the radiation absorbing fluid 48 is liquid or gel and has a property of absorbing a part of X-rays (low energy component). In the present embodiment, a liquid scintillator that absorbs X-rays and converts them into light is used as the radiation absorbing fluid 48.
  • the liquid scintillator is formed of a solute composed of a phosphor, a solvent in which the solute is mixed, a surfactant, and the like.
  • the properties of the liquid scintillator depend on the type and amount of solvent, solute, surfactant and the like.
  • mixed xylene (isomer), toluene, pseudocumene, dioxane or the like is used as the solvent.
  • solute for example, PPO, bulty-PBD, DMPOPOP, Bis-MSB or the like is used.
  • the solution mass is, for example, 0.5 to 5 g / liter to construct a liquid scintillator.
  • PPO is used as a solute, an emission spectrum of 440 nm or less is obtained. It is also possible to shift the emission spectrum to the long wave side by mixing a wavelength conversion material in the solute.
  • the light absorbing layer 21 absorbs the light converted from the X-rays by the radiation absorbing unit 20 to prevent the light of the radiation absorbing unit 20 from being incident on the sensor panel 25. This prevents the image quality of the radiation image.
  • the light absorption layer 21 is formed of a material capable of absorbing light emitted by the liquid scintillator.
  • the imaging region of the subject H when the imaging region of the subject H is a hand, the imaging region is smaller than the imaging range of the FPD 19. Moreover, in such a case, the imaging region is mostly imaged in a state of being disposed at the center of the imaging range of the FPD 19. For this reason, the peripheral portion other than the central portion where the imaging region is arranged becomes a blank area where the X-rays are directly irradiated to the FPD 19, and characteristic deterioration easily occurs in the CMOS sensor 33.
  • the X-ray incident on the FPD 19 has a small amount of X-ray absorption by the radiation absorbing portion 20 in the central portion that affects the image quality of the radiation image, and in the peripheral portion corresponding to one missing region Preferably, the amount of absorption is large.
  • the radiation absorbing unit 20 can be deformed by the weight of the subject H. Specifically, as shown in FIG. 9, when the subject H is placed at the central portion of the radiation absorbing unit 20, the bag 47 is deformed by its weight, and the radiation absorbing fluid 48 is a bag. It flows in the bag 47 in accordance with the deformation of the body 47. As a result, in the radiation absorbing unit 20, the thickness of the central portion immediately below the subject H is reduced, and the thickness of the peripheral portion is increased.
  • the X-ray intensity I after passing through the object is represented by the following equation (1), and decreases exponentially as the thickness T of the object increases when the attenuation coefficient ⁇ is constant. Therefore, when the radiation absorbing portion 20 is deformed as shown in FIG. 9, the X-ray absorption amount at the peripheral portion becomes high, so that the characteristic deterioration of the CMOS sensor 33 in the blank area is suppressed. Further, the amount of absorption of X-rays at the central portion of the radiation absorbing portion 20 is reduced, and the subject H is in a state of being close to the top 13, so X-ray scattering by the radiation absorbing portion 20 hardly occurs. Image quality deterioration is suppressed.
  • I I'exp (- ⁇ T) (1)
  • I X-ray intensity after transmission
  • I ' X-ray intensity before transmission
  • ⁇ Attenuation coefficient (different according to the type of substance and density)
  • T Transmission thickness
  • the single crystal semiconductor substrate 34 of the CMOS sensor 33 is easily broken in material and thin as about several tens of ⁇ m, it is preferable that the single crystal semiconductor substrate 34 have a protective structure that prevents damage due to the load of the object H applied to the top plate 13 .
  • the sensor panel 25 configured of four CMOS sensors 33, when an impact or load is applied to the central portion of the sensor panel 25, all four CMOS sensors 33 may be damaged, and the repair cost is very high. large.
  • the radiation absorbing portion 20 is deformed by the load from the subject H to absorb the load and the impact applied to the central portion of the top 13, so that the CMOS sensor 33 can be prevented from being damaged.
  • the radiation absorbing unit 20 can suppress the heat of the subject H from being transmitted to the sensor panel 25, and can also suppress the image quality deterioration of the radiation image due to the temperature unevenness of the sensor panel 25.
  • the electric circuit of the radiation image capturing apparatus 7 includes a sensor panel 25, a signal processing unit 50, an image memory 51, a control unit 52, a wireless communication unit 53, a battery 54 and the like.
  • the signal processing unit 50 includes an amplifier for amplifying a pixel signal output from each pixel 33 a of the sensor panel 25, an A / D (analog / digital) converter, and the like, and the pixel signal output from the sensor panel 25 Into digital image data.
  • the signal processing unit 50 further includes an image correction unit 50 a that corrects a radiation image according to the absorption distribution of low energy components of X-rays by the radiation absorption unit 20. Since the region to be imaged is mostly placed at the central portion, the central portion of the radiation absorbing unit 20 becomes thinner at the time of imaging, and the X absorption amount is smaller than that at the peripheral portion. The state of depression at the center depends on the imaging site. Therefore, the thickness of each part with respect to the imaging site is determined experimentally. Then, a correction coefficient for obtaining an X-ray absorption amount for each pixel or area corresponding to the thickness amount is set in the image correction unit 50a.
  • the image correction unit 50a corrects the image data using the correction coefficient by designating the imaging region with the console 8, thereby removing the influence of the X-ray absorption amount of the radiation absorption unit 20.
  • the accuracy of the correction can be improved by inputting the gender or the weight and the height.
  • An image memory 51 is connected to the signal processing unit 50, and the image data output from the image correction unit 50 a of the signal processing unit 50 is stored in the image memory 51.
  • the image memory 51 has a storage capacity capable of storing image data of a plurality of frames. Image data obtained by imaging is sequentially stored in the image memory 51 each time a radiographic image is captured.
  • the control unit 52 includes a CPU 52a, a RAM 52b, and a ROM 52, and controls the overall operation of the radiation imaging device 7.
  • the RAM 52 b is a temporary storage memory made of a DRAM or the like.
  • the ROM 52 is a non-volatile memory including a flash memory or the like.
  • the wireless communication unit 53 is an IEEE (Institute of Electrical and Electronics). Engineers) correspond to wireless LAN (Local Area Network) standards represented by 802.11a / b / g / n etc., and enable wireless communication of various information with external devices.
  • the control unit 52 performs wireless communication with the console 8 via the wireless communication unit 53, and transmits and receives various types of information to and from the console 8.
  • the battery 54 supplies power to each part in the radiation imaging device 7.
  • the battery 54 is a rechargeable secondary battery, and is detachable from the radiation imaging device 7.
  • the signal processing unit 50, the image memory 51, the control unit 52, and the wireless communication unit 53 may be provided in the drive circuit board 29. These are connected to one another via a bus.
  • the console 8 includes a CPU 57, a ROM 58, a RAM 59, and an HDD 60, which are connected to one another via a bus 67.
  • the CPU 57 controls each part of the console 8.
  • the ROM 58 stores various programs including a control program.
  • the RAM 59 temporarily stores various data.
  • the HDD 60 stores various data.
  • a communication I / F 61, a wireless communication unit 62, a display driver 64, and an operation input detection unit 66 are connected to the bus 67.
  • a display 63 is connected to the display driver 64.
  • An operation panel 65 is connected to the operation input detection unit 66.
  • the communication I / F 61 is connected to the communication I / F 70 of the radiation generation device 6 via the connection terminal 61 a, the communication cable 69, and the connection terminal 70 a of the radiation generation device 6.
  • the CPU 57 of the console 8 performs transmission and reception of various information such as the irradiation condition with the radiation generator 6 via the communication I / F 61.
  • the wireless communication unit 62 has a function of performing wireless communication with the wireless communication unit 53 of the radiation image capturing apparatus 7.
  • the CPU 57 of the console 8 exchanges various information such as image data with the radiation imaging device 7 via the wireless communication unit 62.
  • the display driver 64 generates and outputs a signal for displaying various information on the display 63.
  • the CPU 57 of the console 8 causes the display 63 to display the operation menu, the captured radiographic image, and the like via the display driver 64.
  • the operation panel 65 has a plurality of keys, and can input various information and operation instructions.
  • the operation input detection unit 66 detects an operation performed on the operation panel 65, and transmits the detection result to the CPU 57.
  • the radiation generating apparatus 6 is configured of a communication I / F 70, a radiation source control unit 72, and a radiation source 6a.
  • the communication I / F 70 transmits and receives various information such as an irradiation condition to and from the console 8.
  • the radiation source control unit 72 controls the radiation source 6 a based on the irradiation conditions (including the information of the tube voltage and the tube current) received from the console 8.
  • the photographer inserts the radiation imaging device 7 between the imaging region of the subject H and the imaging table with the irradiation surface 11 facing the imaging region.
  • the orientation, position, and the like of the radiation imaging device 7 are finely adjusted.
  • the radiation imaging device 7 when the imaging region is directly placed on the radiation imaging device 7, the radiation imaging device 7 is subjected to an impact or a load when mounting the imaging region.
  • the impact and load are absorbed by deformation of the radiation absorbing portion 20.
  • the radiation image capturing apparatus 7 of the present embodiment is an ISS method, and the sensor panel 25 is disposed on the top plate 13 side of the scintillator 26, so that breakage of the sensor panel 25 is effectively prevented.
  • the photographer When the preparation work is completed, the photographer operates the operation panel 65 to instruct start of imaging.
  • the console 8 transmits an instruction signal instructing the start of exposure to the radiation generator 6, and the radiation generator 6 causes the radiation source 6a to emit X-rays.
  • the low energy component is absorbed by the radiation source filter 6b, the X-ray emitted from the radiation source 6a is transmitted through the imaging region of the subject H, and is irradiated to the irradiation surface 11 through the radiation absorbing portion 20 and the light absorbing layer 21. Ru.
  • the X-rays pass through the top plate 13 and the sensor panel 25 and enter the scintillator 26.
  • the low energy component of the X-rays transmitted through the object H is absorbed by the radiation absorbing unit 20, so that the characteristic deterioration of each CMOS sensor 33 due to the sensor panel 25 absorbing the X-rays is suppressed.
  • the radiation absorbing portion 20 is deformed by the weight of the object H, and the thickness of the peripheral portion is thicker than that of the central portion. Therefore, the characteristic deterioration due to direct irradiation of the X-ray to the peripheral portion Is effectively suppressed.
  • the radiation absorbing portion 20 has a central portion thinner than the peripheral portion, X-rays transmitted through the imaging region of the subject H are not absorbed more than necessary, and the image quality of the radiation image is greatly degraded. There is no.
  • the X-rays incident on the scintillator 26 are converted into visible light in the vicinity of the X-ray incident surface of the scintillator 26, that is, on the sensor panel 25 side.
  • the visible light generated by the scintillator 26 the visible light propagating toward the sensor panel 25 is incident on the sensor panel 25.
  • the visible light propagated toward the support substrate 27 is reflected by the reflective layer of the support substrate 27 and then passes through the scintillator 26 again toward the sensor panel 25, and the sensor The light is incident on the panel 25. Since the visible light emitted by the scintillator 26 is guided by the columnar crystal made of CsI: Tl, image blurring is suppressed.
  • the visible light incident on the scintillator 26 is transmitted through the second electrode 38 and incident on the photoelectric conversion layer 37, and is converted into signal charge by the photoelectric conversion layer 37.
  • the signal charge generated in the photoelectric conversion layer 37 is collected by the first electrode 36 and converted to a voltage signal by the signal output circuit 41 after the end of the X-ray emission.
  • This voltage signal is sequentially output from each pixel 33a as a pixel signal.
  • the output pixel signals are converted into image data by the signal processing unit 50.
  • the image data is corrected by the image correction unit 50a according to the amount of X-ray absorption which changes with the thickness of each portion of the radiation absorption unit 20.
  • the image data after correction is stored in the image memory 51.
  • the CPU 52 a transmits the image data stored in the image memory 51 to the console 8 via the wireless communication unit 53.
  • the CPU 57 of the console 8 stores the image data received from the radiation imaging device 7 in the HDD 60 via the RAM 59. Further, the CPU 57 causes the display 63 to display a radiation image based on the image data stored in the HDD 60 via the display driver 64.
  • the characteristic deterioration of the CMOS sensor 33 is suppressed.
  • the radiation absorbing unit 20 is deformed according to the position of the subject H placed on the radiation imaging device 7, and the amount of absorption of the low energy component changes, so that the image quality of the radiation image is not deteriorated. Characteristic deterioration of the CMOS sensor 33 in the missing region is suppressed.
  • the conventional ISS-type radiation imaging apparatus using a TFT panel is not easy to apply to mammography with a low tube voltage because the X-ray absorption of alkali-free glass used as a substrate of the TFT panel is large.
  • a CMOS sensor with low X-ray absorption can be used for a radiation imaging apparatus of ISS method, application to mammography is easy.
  • the light absorbing layer 21 is separately provided to absorb the light emitted from the radiation absorbing unit 20, but the light is applied to the bonding surface of the bag 47 with the top plate 13 or the entire bag 47. It may have absorbency.
  • the wavelength range of the light emitted from the radiation absorbing fluid 48 may be different from the wavelength range detectable by the sensor panel 25 by appropriately selecting the type of solute or the wavelength conversion material. Thereby, the light absorption layer 21 can be omitted.
  • a liquid scintillator is used as the radiation absorbing fluid 48, but a paraffin scintillator may be used.
  • the paraffin scintillator is waxy at low temperature and has fluidity at about 40 ° C.
  • the paraffin scintillator is described in detail in Japanese Patent Application Laid-Open No. 07-301676, so the detailed description is omitted.
  • gelled water glass sodium silicate
  • the gelation of water glass occurs, for example, by adding an acid to water glass and stirring.
  • the radiation absorbing fluid 48 may be configured by dispersing an X-ray absorbing material in a liquid or gel fluid.
  • an X-ray absorbing material metals of atomic numbers 20 to 31 (for example, copper) having a high effect of absorbing low energy components of X-rays and little back scattering are preferable.
  • the radiation absorption part 20 is comprised with the bag body 47 and the radiation absorptive fluid body 48, you may form with elastic bodies, such as rubber
  • the one radiation absorption part 20 is provided in the radiographic imaging apparatus 7, you may provide several radiation absorption parts.
  • a plate-like low energy absorbing plate 75 for absorbing low energy components of X-rays is disposed between the top plate 13 and the sensor panel 25.
  • the low energy absorbing plate 75 may be disposed between the top plate 13 and the radiation absorbing unit 20, or the low energy absorbing plate 75 may be used instead of the top plate 13.
  • the low energy absorbing plate 75 for example, aluminum, glass or the like can be used, but the same material as the above-mentioned source filter 6b may be included.
  • the transmitted X-ray energy distribution becomes the same.
  • X-rays that can pass through the radiation source filter 6b are difficult to be absorbed by the radiation absorbing unit 20 as well, so that it is possible to suppress X-rays including information of the object H from being absorbed by the radiation absorbing unit 20. As a result, the utilization efficiency of X-rays is improved.
  • the CMOS sensor may be configured by an organic thin film transistor formed on a plastic film so as to impart flexibility to the CMOS sensor.
  • organic thin film transistors see “Tsuyoshi Sekitani, Flexible organic”. Since this is described in detail in “Transistors and circuits with extreme bending stability, Nature Materials 9, November 7, 2010, pp. 1015-1022”, the detailed description will be omitted.
  • a photodiode and a transistor formed of single crystal Si may be provided over a flexible plastic substrate.
  • a device block of about several tens of microns in size is dispersed in a solution, and FAS (Fluidic Self- Assembly) can be used.
  • FSA Fluidic Self- Assembly
  • “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 May 6, 108, 87, p. Detailed description will be omitted as it is described in detail in 67-71.
  • the sensor panel is constituted by a CMOS sensor, but the present invention is also applicable to a radiation image photographing apparatus in which the sensor panel is constituted by a CCD image sensor formed by a single crystal semiconductor substrate. .
  • the present invention is not limited to the ISS method, and is also applicable to a PSS type radiation imaging apparatus.
  • the present invention is also applicable to a direct conversion type radiation imaging apparatus that converts radiation directly into electric charges.
  • the FPD is incorporated into a cassette-sized casing
  • a standing-type or lying-down type imaging device or a mammography device it is also possible to be incorporated into a standing-type or lying-down type imaging device or a mammography device.
  • the present invention is also applicable to a radiation imaging apparatus using radiation other than X-rays, such as ⁇ -rays.
  • the configuration of the radiation image capturing apparatus according to the present invention described in the above embodiment is an example, and it goes without saying that the configuration can be appropriately changed without departing from the scope of the present invention.

Abstract

A radiation-image-capturing device (7) has a sensor panel (25) made of a plurality of CMOS sensors (33). The sensor panel (25) has a signal-outputting circuit formed on a monocrystal semiconductor substrate. The top surface of a top plate (13) irradiated by x-rays is provided with a radiation-absorbing unit (20) for absorbing some of the x-rays (low-energy components) responsible for characteristics degradation of the signal-outputting circuit. The radiation-absorbing unit (20) is made of a bag (47) and a radiation-absorbing fluid (48) accommodated within the bag (47). The radiation-absorbing unit (20) is deformed under loading from a subject (H).

Description

放射線画像撮影装置Radiation imaging device
　本発明は、放射線画像を撮影する放射線画像撮影装置に関する。 The present invention relates to a radiation imaging apparatus for capturing a radiation image.
　医療分野において、画像診断を行うために、放射線（例えば、Ｘ線）を利用して被写体（患者の撮影部位）を撮影する放射線撮影システムが知られている。放射線撮影システムは、放射線を照射する放射線発生装置と、撮影部位の放射線画像を撮影する放射線画像撮影装置とを有する。放射線画像撮影装置には、立位撮影台や臥位撮影台に組み込まれた据え置き型のものや、持ち運び可能な可搬型のもの（いわゆる電子カセッテ）がある。可搬型の放射線画像撮影装置は、病室等でベッドに寝ている患者の下に挿入して撮影することができる。 2. Description of the Related Art In the medical field, there is known a radiation imaging system for imaging a subject (a region to be imaged by a patient) using radiation (for example, X-rays) to perform image diagnosis. The radiation imaging system has a radiation generating device for emitting radiation and a radiation imaging device for capturing a radiation image of an imaging region. 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.
　放射線画像撮影装置には、放射線の入射量に応じた信号電荷を蓄積する画素がマトリクス状に配列された検出面を有するＦＰＤ（flat panel detector）を放射線検出器として用いたものが実用化されている。ＦＰＤでは、検出面において画素毎に信号電荷を蓄積することで、放射線画像を検出し、これをデジタルの画像データとして出力する。 As a radiation imaging apparatus, one using an FPD (flat panel detector) having a detection surface in which pixels accumulating signal charges according to the amount of incident radiation are arranged in a matrix as a radiation detector is put to practical use. There is. The FPD detects a radiation image by accumulating signal charges for each pixel on the detection surface, and outputs this as digital image data.
　ＦＰＤには、アモルファスセレン（ａ－Ｓｅ）等からなる変換層で放射線を直接信号電荷に変換する直接変換型ＦＰＤと、放射線を一旦可視光に変換し、可視光を信号電荷に変換する間接変換型ＦＰＤが知られている。間接型ＦＰＤは、放射線を可視光に変換するシンチレータと、このシンチレータに対向して配置された検出パネルと、電気制御回路とで構成されている。検出パネルは、光電変換により信号電荷を発生する光電変換部を画素ごとに形成した検出面を有しており、シンチレータからの可視光を信号電荷に変換して蓄積する。 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 detection panel disposed opposite to the scintillator, and an electrical control circuit. The detection panel has a detection surface in which a photoelectric conversion unit that generates a signal charge by photoelectric conversion is formed for each pixel, converts visible light from the scintillator into a signal charge, and stores the signal charge.
　検出パネルとしては、ガラス基板上にＴＦＴ（thin-film transistor）と光電変換部とをマトリクス状に配列したＴＦＴパネルや、ＣＭＯＳ型イメージセンサ（以下、ＣＭＯＳセンサという）が用いられる。ＴＦＴは、アモルファスシリコン（ａ－Ｓｉ）等の非晶質半導体により形成されている。ＣＭＯＳセンサは、シリコン（Ｓｉ）の単結晶半導体基板に、半導体プロセスにより光電変換部とＭＯＳトランジスタとがマトリクス状に形成されている。 As a detection 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.
　ＣＭＯＳセンサのＭＯＳトランジスタは、単結晶半導体により形成されているため、非晶質半導体で形成されたＴＦＴパネルに比べて、キャリア移動度が３～４桁以上高く、信号電荷の高速読み出しが可能である。また、ＣＭＯＳセンサは、光電変換部やＭＯＳトランジスタの製造時の特性（例えば、ＭＯＳトランジスタの閾値電圧等）のばらつきが小さいため、高Ｓ／Ｎの画像を得ることが可能である。このように、ＣＭＯＳセンサは、動画撮影や高画質撮影に適している。 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センサは、現在では１２インチウエハを用いて、四角形の一辺が約２００ｍｍのサイズを有するものが製造可能である。このため、例えば、医療用として一般的な一辺が１７インチのサイズを有するＦＰＤは、４枚のＣＭＯＳセンサを用いて構成することができる。 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, an FPD having a size of 17 inches, which is generally used for medical use, can be configured using four CMOS sensors.
　一般的に、単結晶半導体基板は脆く割れやすいため、可搬型の放射線画像撮影装置のＦＰＤとして、単結晶半導体基板を有するＣＭＯＳセンサを用いる場合には、患者の下に放射線画像撮影装置をセットした場合の破損を防止するための保護構造が必要となり、放射線画像撮影装置が大型化してしまう。 Generally, since a single crystal semiconductor substrate is fragile and easily broken, when using a CMOS sensor having a single crystal semiconductor substrate as an FPD of a portable radiation image capturing apparatus, the radiation image capturing apparatus is set under the patient A protective structure for preventing damage in the case is required, and the radiographic imaging device becomes large.
　また、単結晶半導体基板では、放射線照射により、ＭＯＳトランジスタの閾値電圧の変化や、暗電流の増加等の特性劣化が発生することが知られている。これは、単結晶半導体基板を用いたＭＯＳ構造では、放射線の吸収によって単結晶半導体基板と酸化膜との界面に電荷（以下、界面電荷という）が生じて蓄積されるためである。放射線照射による単結晶半導体の特性劣化は、被写体のサイズがＦＰＤの検出面よりも小さい場合には、被写体を通過しない放射線が入射する領域（素抜け領域）で界面電荷が大きくなるため、特に問題となる。 In addition, in single crystal semiconductor substrates, it is known that radiation irradiation causes characteristic changes such as a change in threshold voltage of a MOS transistor and an increase in dark current. This is because in the MOS structure using a single crystal semiconductor substrate, charges (hereinafter referred to as interface charges) are generated and accumulated at the interface between the single crystal semiconductor substrate and the oxide film by absorption of radiation. The characteristic deterioration of the single crystal semiconductor due to the radiation irradiation is particularly problematic when the size of the object is smaller than the detection surface of the FPD, since the interface charge becomes large in the region where the radiation which does not pass through the object is incident (clear region). It becomes.
　特開２００５－２４９６３９号公報には、被写体と放射線画像撮影装置との間に、被写体によって散乱されたＸ線を吸収する散乱Ｘ線吸収グリッドを配置し、この散乱Ｘ線吸収グリッドの外周に、素抜け領域に照射されたＸ線を吸収する吸収部を設けたものが開示されている。また、特開２０１０－０７５５５３号公報には、放射線発生装置に、素抜け領域に照射されるＸ線の吸収量を増加させたフィルタを配置することが開示されている。しかし、特開２００５－２４９６３９号公報に記載の吸収部と、特開２０１０－０７５５５３号公報に記載のフィルタとは、いずれも撮影に必要なＸ線成分を吸収してしまい、放射線画像の画質劣化をもたらす。また、特開２００５－２４９６３９号公報と特開２０１０－０７５５５３号公報とには、いずれもＣＭＯＳセンサの特性劣化の防止や、破損防止については記載がない。 In Japanese Patent Application Laid-Open No. 2005-249639, a scattered X-ray absorbing grid for absorbing X-rays scattered by a subject is disposed between the subject and the radiation imaging apparatus, and the outer periphery of the scattered X-ray absorbing grid is disposed What provided the absorption part which absorbs the X-ray irradiated to the blank area is disclosed. Further, Japanese Patent Application Laid-Open No. 2010-075553 discloses that a radiation generating apparatus is provided with a filter in which the amount of absorption of X-rays irradiated to the blank area is increased. However, both the absorbing unit described in JP-A-2005-249639 and the filter described in JP-A-2010-075553 absorb X-ray components necessary for imaging, and the image quality of the radiation image is degraded. Bring Further, neither JP-A-2005-249639 nor JP-A-2010-075553 describes the prevention of deterioration of the characteristics of the CMOS sensor and the prevention of damage.
　本発明は、放射線による放射線検出器の特性劣化を抑制するととともに、被写体からの荷重や衝撃による破損を防止することができる放射線画像撮影装置を提供することを目的とする。 An object of the present invention is to provide a radiation image capturing apparatus capable of preventing characteristic deterioration of a radiation detector due to radiation and preventing damage due to a load or an impact from a subject.
　上記課題を解決するために、本発明の放射線画像撮影装置は、放射線検出器と放射線吸収部とを備える。放射線検出器は、被写体を透過した放射線を検出して画像データを生成する。放射線吸収部は、放射線検出器の放射線入射側に配置され、放射線検出器に入射する放射線の一部を吸収するとともに、被写体の荷重によって変形可能である。放射線検出器は、単結晶半導体基板を有する。 In order to solve the above-mentioned subject, a radiation imaging device of the present invention is provided with a radiation detector and a radiation absorption part. The radiation detector detects radiation transmitted through the subject and generates image data. The radiation absorbing portion is disposed on the radiation incident side of the radiation detector, absorbs a part of the radiation incident on the radiation detector, and is deformable by the load of the subject. The radiation detector has a single crystal semiconductor substrate.
　放射線吸収部は、袋体と、この袋体に封入された液状またはジェル状の放射線吸収性流動体とを有する。放射線吸収性流動体としては、放射線を吸収して光に変換する液状シンチレータを用いてもよい。この場合、放射線検出器と放射線吸収部との間に、液状シンチレータにより変換された光を吸収する光吸収層を設けることが好ましい。 The radiation absorbing portion has a bag and a liquid or gel radiation absorbing fluid sealed in the bag. As the radiation absorbing fluid, a liquid scintillator that absorbs radiation and converts it into light may be used. In this case, it is preferable to provide, between the radiation detector and the radiation absorbing portion, a light absorbing layer that absorbs the light converted by the liquid scintillator.
　放射線吸収性流動体としては、ゲル化された水ガラスを用いてもよい。また、放射線吸収性流動体は、液状またはジェル状の流動体と、流動体中に分散された放射線吸収材とを有するものであってもよい。また、放射線吸収部は、被写体からの荷重で変形する弾性体と、この弾性体中に分散された放射線吸収材とを有するものであってもよい。また、放射線吸収部を複数設けてもよい。 A gelled water glass may be used as the radiation absorbing fluid. The radiation absorbing fluid may also have a liquid or gel fluid and a radiation absorbing material dispersed in the fluid. Further, the radiation absorbing portion may have an elastic body which is deformed by a load from a subject, and a radiation absorbing material dispersed in the elastic body. Also, a plurality of radiation absorbing portions may be provided.
　放射線検出器は、放射線を吸収して光に変換するシンチレータと、シンチレータの放射線照射側に配置され、シンチレータにより変換された光を検出するセンサパネルとを有するものである。 The radiation detector has 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 that detects light converted by the scintillator.
　シンチレータ及びセンサパネルを収容する筐体を有し、この筐体上に放射線吸収部が取り付けられている。放射線吸収部は、袋体と、この袋体に封入された液状シンチレータとを有することが好ましい。 It has a housing which accommodates a scintillator and a sensor panel, and a radiation absorption part is attached on this housing. The radiation absorber preferably has a bag and a liquid scintillator enclosed in the bag.
　センサパネルとして、シンチレータで変換された光を電荷に変換する光電変換層と、光電変換層で変換された電荷に応じた信号を出力する信号出力回路とを含む画素を複数有し、信号出力回路は、単結晶半導体基板に設けられているものを用いることができる。このようなセンサパネルとして、ＣＭＯＳ型のイメージセンサにより構成されたものが用いられる。光電変換層は、アモルファスシリコンまたはキナクリドンで形成されていることが好ましい。 A signal output circuit includes a plurality of pixels including a photoelectric conversion layer that converts light converted by a scintillator into electric charge and a signal output circuit that outputs a signal according to the electric charge converted by the photoelectric conversion layer as a sensor panel Any of those provided over a single crystal semiconductor substrate can be used. As such a sensor panel, one configured by a CMOS type image sensor is used. The photoelectric conversion layer is preferably formed of amorphous silicon or quinacridone.
　放射線吸収部として、液状シンチレータを用いる場合には、液状シンチレータが発生する光の波長域とセンサパネルが検出可能な光の波長域とが異なるように構成することが好ましい。 When a liquid scintillator is used as the radiation absorbing portion, it is preferable that the wavelength range of light generated by the liquid scintillator be different from the wavelength range of light detectable by the sensor panel.
　放射線吸収部は、放射線検出器よりもサイズが大きいことが好ましい。放射線吸収部の放射線吸収量の領域差に応じて画像データを補正する画像補正部を備えることが好ましい。放射線吸収部は、放射線の低エネルギ成分を吸収することが好ましい。放射線吸収部が吸収する放射線の低エネルギ成分は、放射線のエネルギ分布の１／２以下のエネルギ成分である。 The radiation absorber is preferably larger in size than the radiation detector. It is preferable to include an image correction unit that corrects image data according to the area difference of the radiation absorption amount of the radiation absorption unit. The radiation absorbing portion preferably absorbs the low energy component of the radiation. The low energy component of the radiation absorbed by the radiation absorbing portion is an energy component of 1/2 or less of the energy distribution of the radiation.
　本発明の放射線画像撮影装置によれば、放射線吸収部により放射線の一部を吸収するので、放射線検出器の特性劣化を抑制することができる。また、放射線吸収部により放射線検出器を補強し、放射線検出器の破損を防止することができる。 According to the radiation image capturing apparatus of the present invention, since the radiation absorbing unit absorbs a part of the radiation, it is possible to suppress the characteristic deterioration of the radiation detector. Further, the radiation detector can be reinforced by the radiation absorbing portion to prevent damage to the radiation detector.
　また、被写体からの荷重より放射線吸収部が変形するので、被写体の位置、大きさ、範囲に合わせて放射線の吸収分布を適正化される。これにより、被写体を透過した放射線は、放射線吸収部による吸収量が小さいため、放射線画像の画質は劣化せず、放射線が直接照射される素抜け領域では吸収量が多く、放射線検出器の特性劣化が抑制される。 Further, since the radiation absorbing portion is deformed by the load from the subject, the radiation absorption distribution is optimized according to the position, size, and range of the subject. As a result, the radiation transmitted through the subject does not deteriorate the image quality of the radiation image because the amount of radiation absorbed by the radiation absorbing portion is small, and the amount of radiation absorbed in the blank area directly irradiated with radiation is large, and the characteristics of the radiation detector deteriorate Is suppressed.
放射線撮影システムの構成図である。It is a block diagram of a radiography system. 放射線画像撮影装置を一部破断して示す斜視図である。It is a perspective view which fractures | ruptures and shows a radiographic imaging apparatus partially. 放射線画像撮影装置の断面図である。It is sectional drawing of a radiographic imaging apparatus. センサパネルの構成を示す平面図である。It is a top view which shows the structure of a sensor panel. ＦＰＤの構成を示す断面図である。It is sectional drawing which shows the structure of FPD. 光電変換層の感度域及びシンチレータの発光領域を示すグラフである。It is a graph which shows the sensitivity area of a photoelectric converting layer, and the luminescence field of a scintillator. 画素の構成を示す回路図である。It is a circuit diagram showing composition of a pixel. 放射線吸収部上に被写体が載置された状態を示す斜視図である。It is a perspective view which shows the state in which the subject was mounted on the radiation absorption part. 被写体の荷重により放射線吸収部が変形した状態を示す放射線画像撮影装置の断面図である。It is sectional drawing of the radiographic imaging apparatus which shows the state which the radiation absorption part deform | transformed by the load of a to-be-photographed object. 放射線画像撮影装置の電気的構成を示すブロック図である。It is a block diagram which shows the electric constitution of a radiographic imaging apparatus. コンソール及び放射線発生装置の電気的構成を示すブロック図である。It is a block diagram which shows the electric constitution of a console and a radiation generation apparatus. 放射線吸収部を複数設けた放射線画像撮影装置の構成を示す断面図である。It is sectional drawing which shows the structure of the radiographic imaging apparatus which provided multiple radiation absorption parts.
　図１において、放射線画像撮影システム５は、被写体（患者）Ｈの撮影部位に向けて放射線としてＸ線を照射する放射線発生装置６と、被写体Ｈの放射線画像を撮影する放射線画像撮影装置７と、放射線発生装置６と放射線画像撮影装置７とを制御するコンソール８とを備える。 In FIG. 1, a radiation imaging system 5 comprises: a radiation generator 6 for emitting X-rays as radiation toward the imaging region of a subject (patient) H; a radiation imaging apparatus 7 for imaging a radiation image of the subject H; A console 8 for controlling the radiation generator 6 and the radiation imaging device 7 is provided.
　放射線発生装置６は、放射線源６ａと、線源フィルタ６ｂとを有する。放射線源６ａは、Ｘ線を放射するＸ線管６ｃと、Ｘ線管６ｃが放射するＸ線の照射野を限定する照射野限定器（コリメータ）６ｄとを有している。 The radiation generator 6 has a radiation source 6a and a source filter 6b. The radiation source 6a has an X-ray tube 6c for emitting X-rays and a radiation field limiter (collimator) 6d for limiting the radiation field of X-rays emitted by the X-ray tube 6c.
　Ｘ線管６ｃは、熱電子を放出するフィラメントからなる陰極と、陰極から放出された熱電子が衝突してＸ線を放射する陽極（ターゲット）とを有している。照射野限定器６ｄは、例えば、Ｘ線を遮蔽する複数枚の鉛板を四角形の各辺に配置し、Ｘ線を透過させる照射開口を中央に形成したものであり、鉛板の位置を移動することで照射開口の大きさを変化させて、照射野を限定する。 The X-ray tube 6c has a cathode formed of a filament that emits thermoelectrons, and an anode (target) that emits thermorays which collide with the thermoelectrons emitted from the cathode. The irradiation field limiter 6d is, for example, a plurality of lead plates for shielding X-rays arranged on each side of a square and an irradiation opening for transmitting X-rays formed at the center, and the position of the lead plate is moved By changing the size of the irradiation opening, the irradiation field is limited.
　線源フィルタ６ｂは、放射線源６ａから放射されたＸ線から、撮影部位を透過する際に散乱して放射線画像を劣化させる原因となる低エネルギ成分を除去する。線源フィルタ６ｂには、Ｘ線の低エネルギ成分のみを吸収する性質を有する材料が用いられる。このような材料としては、例えばアルミニウムが好適である。線源フィルタ６ｂを透過したＸ線の高エネルギ成分が被写体Ｈの撮影に用いられる。 The radiation source filter 6b removes low energy components which are scattered when passing through the imaging region from the X-rays emitted from the radiation source 6a and cause deterioration of the radiation image. For the source filter 6b, a material having a property of absorbing only the low energy component of the X-ray 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.
　Ｘ線管６ｃから放射されるＸ線のエネルギ分布は、例えば、Ｘ線管６ｃの管電圧が７０ｋＶであり、Ｘ線管６ｃから放射されるＸ線の最大エネルギが７０ＫｅＶ程度であるときには、おおよそ１５～７０ＫｅＶである。本実施形態では、このＸ線のエネルギ分布の約１／２以下（１５～４０ＫｅＶ）を低エネルギ成分とし、１／２以上（４０～７０ＫｅＶ）を高エネルギ成分とする。線源フィルタ６ｂは、１５～４０ＫｅＶの低エネルギ成分を吸収する。 The energy distribution of the X-rays emitted from the X-ray tube 6c is, for example, approximately when the tube voltage of the X-ray tube 6c is 70 kV and the maximum energy of the X-rays emitted from the X-ray tube 6c is about 70 KeV 15 to 70 KeV. In this embodiment, about 1/2 or less (15 to 40 KeV) of the energy distribution of the X-ray is a low energy component, and 1/2 or more (40 to 70 KeV) is a high energy component. The source filter 6b absorbs low energy components of 15 to 40 KeV.
　図２において、放射線画像撮影装置７は、ＦＰＤ１９と、放射線吸収部２０と、光吸収層２１と、電気回路部２３と、可搬型の筐体１２とで構成されている。筐体１２は、天板１３と、扁平な箱形状の筐体本体１４とを有し、ＦＰＤ１９と電気回路部２３とを収容している。天板１３は、筐体本体１４の上部の開口部１４ａを封止している。 In FIG. 2, the radiation image capturing apparatus 7 includes an FPD 19, a radiation absorbing unit 20, a light absorbing layer 21, an electric circuit unit 23, and a portable case 12. The housing 12 has a top plate 13 and a flat box-shaped housing body 14, and accommodates the FPD 19 and the electric circuit unit 23. The top plate 13 seals the opening 14 a at the top of the case body 14.
　天板１３の上面は、放射線発生装置６から射出されたＸ線が照射される照射面１１である。このため、天板１３は、Ｘ線の透過性が高いカーボン等で形成されている。カーボンは高強度であるため、被写体Ｈの体重がかかる天板１３の材料として好適である。筐体本体１４はＡＢＳ樹脂等で形成されている。 The upper surface of the top plate 13 is an irradiation surface 11 to which X-rays emitted from the radiation generator 6 are irradiated. For this reason, the top plate 13 is formed of carbon or the like having high X-ray permeability. Since carbon has high strength, it is suitable as a material of the top 13 to which the weight of the subject H is applied. The housing body 14 is formed of ABS resin or the like.
　筐体１２は、放射線画像を感光材料に記録する従来の放射線フィルムカセッテと同じサイズ（例えば、１７インチ×１７インチ）である。放射線画像撮影装置７は、放射線フィルムカセッテと同様に可搬性を有し、放射線フィルムカセッテに代えて用いられる。 The housing 12 is the same size (e.g., 17 inches by 17 inches) as a conventional radiation film cassette for recording a radiation image on a photosensitive material. The radiographic imaging device 7 has the same portability as the radiographic film cassette, and is used in place of the radiographic film cassette.
　天板１３には、複数個のＬＥＤにより構成された表示部１６が設けられている。表示部１６には、放射線画像撮影装置７の動作モード（例えば「レディ状態」や「データ送信中」等）やバッテリの残容量等の動作状態が表示される。なお、表示部１６を、ＬＥＤ以外の発光素子や、液晶ディスプレイや有機ＥＬディスプレイ等で構成してもよい。また、表示部１６を、筐体本体１４に設けてもよい。 The top plate 13 is provided with a display unit 16 configured of a plurality of LEDs. The display unit 16 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 a remaining capacity of a battery. The display unit 16 may be configured by a light emitting element other than an LED, a liquid crystal display, an organic EL display, or the like. In addition, the display unit 16 may be provided in the housing body 14.
　天板１３上には、光吸収層２１と放射線吸収部２０とがこの順に積層されている。放射線吸収部２０は、光吸収層２１を介して天板１３に対面している。放射線吸収部２０は、照射面１１に照射されるＸ線の一部（例えば、Ｘ線の低エネルギ成分）を吸収するとともに、天板１３に加えられる荷重や衝撃等からＦＰＤ１９を保護する。光吸収層２１は、Ｘ線照射によって放射線吸収部２０が発生した光を吸収する。 The light absorbing layer 21 and the radiation absorbing portion 20 are stacked in this order on the top plate 13. The radiation absorbing unit 20 faces the top 13 via the light absorbing layer 21. The radiation absorbing unit 20 absorbs a part of X-rays (for example, low energy components of X-rays) irradiated to the irradiation surface 11 and protects the FPD 19 from a load, an impact or the like applied to the top plate 13. The light absorbing layer 21 absorbs the light generated by the radiation absorbing unit 20 by the X-ray irradiation.
　電気回路部２３は、信号処理部５０、画像メモリ５１、制御部５２、無線通信部５３、バッテリ５４等（いずれも図９参照）を収容している。電気回路部２３は、筐体１２の内部の短手方向に沿った一端側に配置されている。ＦＰＤ１９は、バッテリ５４から供給される電力によって作動する。電気回路部２３の天板１３側には、電気回路部２３がＸ線により損傷することを防止するため、鉛板等の放射線遮蔽部材（図示せず）が設けられている。 The electric circuit unit 23 accommodates a signal processing unit 50, an image memory 51, a control unit 52, a wireless communication unit 53, a battery 54 and the like (all refer to FIG. 9). The electric circuit unit 23 is disposed at one end side along the short direction inside the housing 12. The FPD 19 operates by the power supplied from the battery 54. A radiation shielding member (not shown) such as a lead plate is provided on the top plate 13 side of the electric circuit section 23 in order to prevent the electric circuit section 23 from being damaged by X-rays.
　図３において、天板１３の上面には光吸収層２１が接着剤（図示せず）により貼り付けられている。光吸収層２１の上には、放射線吸収部２０が接着剤（図示せず）により貼り付けられている。光吸収層２１及び放射線吸収部２０は、照射面１１を覆っている。これにより、放射線画像撮影装置７を薄型化し、かつ放射線検出器１９を補強することができる。 In FIG. 3, a light absorbing layer 21 is attached to the top surface of the top plate 13 with an adhesive (not shown). The radiation absorbing portion 20 is attached onto the light absorbing layer 21 by an adhesive (not shown). The light absorbing layer 21 and the radiation absorbing unit 20 cover the irradiation surface 11. Thereby, the radiation imaging device 7 can be thinned and the radiation detector 19 can be reinforced.
　ＦＰＤ１９は、放射線が照射される方向に沿って、天板１３側から、センサパネル２５とシンチレータ２６とが順に積層されたものである。シンチレータ２６の下面には、シンチレータ２を支持する支持基板２７が設けられている。ＦＰＤ１９の外周には、シンチレータ２６を湿気等から保護するために封止剤２８が設けられている。筐体１２内の底面には、ＦＰＤ１９の駆動回路基板２９が配置されている。駆動回路基板２９とセンサパネル２５とは、フレキシブルケーブル３０を介して電気的に接続されている。 The FPD 19 is one in which a sensor panel 25 and a scintillator 26 are sequentially stacked from the top 13 side along the direction in which the radiation is irradiated. A support substrate 27 supporting the scintillator 2 is provided on the lower surface of the scintillator 26. A sealing agent 28 is provided on the outer periphery of the FPD 19 in order to protect the scintillator 26 from moisture and the like. A drive circuit board 29 of the FPD 19 is disposed on the bottom of the housing 12. The drive circuit board 29 and the sensor panel 25 are electrically connected via the flexible cable 30.
　シンチレータ２６は、被写体Ｈを透過して筐体１２の照射面１１に照射され、天板１３、放射線吸収部２０、センサパネル２５を透過して入射したＸ線を吸収して可視光を発生する。シンチレータ２６としては、ＣｓＩ：Ｔｌ（タリウムを添加したヨウ化セシウム）や、ＣｓＩ：Ｎａ（ナトリウムを添加したヨウ化セシウム）、ＧＯＳ（Ｇｄ２Ｏ２Ｓ：Ｔｂ）等が用いられる。本実施形態では、シンチレータ２６として、ＣｓＩ：Ｔｌを用いる。シンチレータ２６は、ＣｓＩ：Ｔｌを支持基板２７に蒸着することにより形成される。このシンチレータ２６は、柱状結晶構造であり、支持基板２７からセンサパネル２５に向かう方向に沿って、複数の柱状結晶（図示せず）を有している。柱状結晶は、その平径が柱状結晶の長手方向に沿ってほぼ均一である。 The scintillator 26 transmits the subject H and is irradiated to the irradiation surface 11 of the housing 12, and transmits the top plate 13, the radiation absorbing unit 20, and the sensor panel 25 to absorb incident X-rays and generate visible light. . As the scintillator 26, CsI: Tl (cesium iodide to which thallium is added), CsI: Na (cesium iodide to which sodium is added), GOS (Gd 2 O 2 S: Tb) or the like is used. In the present embodiment, CsI: Tl is used as the scintillator 26. The scintillator 26 is formed by vapor deposition of CsI: Tl on the support substrate 27. The scintillator 26 has a columnar crystal structure, and has a plurality of columnar crystals (not shown) along the direction from the support substrate 27 toward the sensor panel 25. The columnar crystal has a flat diameter substantially uniform along the longitudinal direction of the columnar crystal.
　シンチレータ２６で発生した光は、柱状結晶のライトガイド効果によって柱状結晶内を伝搬し、柱状結晶の先端部からセンサパネル２５に向かって射出される。このように、シンチレータ２６を柱状結晶構造とすることにより、シンチレータ２６からセンサパネル２５側へ射出される可視光の拡散が抑制されるので、放射線画像撮影装置７によって撮影される放射線画像の鮮鋭度が向上する。 The light generated by the scintillator 26 propagates in the columnar crystal due to the light guide effect of the columnar crystal, and is emitted toward the sensor panel 25 from the tip of the columnar crystal. As described above, by forming the scintillator 26 in a columnar crystal structure, diffusion of visible light emitted from the scintillator 26 to the sensor panel 25 side is suppressed, so the sharpness of the radiation image captured by the radiation imaging device 7 Improve.
　支持基板２７のシンチレータ２６側の表面には、反射層（図示せず）が設けられている。シンチレータ２６で発光し、支持基板２７側に伝搬した可視光は、反射層によりセンサパネル２５側へ反射されるので、センサパネル２５への入射光量（シンチレータ２６で発光した光の検出効率）が向上する。 A reflective layer (not shown) is provided on the surface of the support substrate 27 on the scintillator 26 side. The visible light emitted by the scintillator 26 and propagated to the support substrate 27 is reflected to the sensor panel 25 by the reflective layer, so the amount of light incident on the sensor panel 25 (detection efficiency of light emitted by the scintillator 26) is improved. Do.
　Ｘ線によるシンチレータ２６の発光は、主にシンチレータ２６の表面側（Ｘ線入射側）で生じるため、本実施形態のように、シンチレータ２６のＸ線入射側にセンサパネル２５を配置した場合には、シンチレータ２６での発光位置とセンサパネル２５との距離が近いため、光検出の分解能及び受光量が高い。このように、シンチレータ２６のＸ線入射側にセンサパネル２５を配置する構成は、ＩＳＳ（Irradiation Side Sampling）方式と称される。逆に、シンチレータのＸ線入射側とは反対側にセンサパネルを配置する構成は、ＰＳＳ（Penetration Side Sampling）方式と称される。ＩＳＳ方式では、ＰＳＳ方式より放射線画像の鮮鋭度及び感度が向上する。 Since light emission of the scintillator 26 by X-rays mainly occurs on the surface side (X-ray incident side) of the scintillator 26, when the sensor panel 25 is disposed on the X-ray incident side of the scintillator 26 as in this embodiment Since the distance between the light emitting position of the scintillator 26 and the sensor panel 25 is short, the resolution of light detection and the amount of light received are high. As described above, the configuration in which the sensor panel 25 is disposed on the X-ray incident side of the scintillator 26 is referred to as an ISS (Irradiation Side Sampling) method. Conversely, a configuration in which the sensor panel is disposed on the opposite side to the X-ray incident side of the scintillator is called a PSS (Penetration Side Sampling) method. In the ISS method, the sharpness and sensitivity of the radiation image are improved compared to the PSS method.
　図４において、センサパネル２５は、４枚のＣＭＯＳ型イメージセンサ（以下、ＣＭＯＳセンサという）３３により構成されている。各ＣＭＯＳセンサ３３は、マトリクス状に配置された複数の画素３３ａ（図７参照）を有する。各ＣＭＯＳセンサ３３は、一辺の長さが２００ｍｍ程度の矩形状である。４枚のＣＭＯＳセンサ３３は、上下左右に互いに隣接するように並べられ、およそ一辺が１７インチの四角形を形成する。 In FIG. 4, the sensor panel 25 is configured by four CMOS type image sensors (hereinafter referred to as CMOS sensors) 33. Each CMOS sensor 33 has a plurality of pixels 33a (see FIG. 7) arranged in a matrix. Each CMOS sensor 33 has a rectangular shape with a side length of about 200 mm. The four CMOS sensors 33 are arranged adjacent to each other vertically and horizontally, and form a square of approximately 17 inches on a side.
　ＣＭＯＳセンサ３３は、米国公開２００９／０２２４１６２号公報に開示されたものと同様の構成である。具体的には、図５に示すように、ＣＭＯＳセンサ３３は、単結晶半導体基板３４と、絶縁層３５と、第１電極３６と、光電変換層３７と、第２電極３８とにより構成されている。 The CMOS sensor 33 has a configuration similar to that disclosed in U.S. Patent Publication 2009/0224162. Specifically, as shown in FIG. 5, the CMOS sensor 33 is constituted of a single crystal semiconductor substrate 34, an insulating layer 35, a first electrode 36, a photoelectric conversion layer 37, and a second electrode 38. There is.
　単結晶半導体基板３４は、単結晶Ｓｉで作られている。絶縁層３５は、単結晶半導体基板３４の表面上に形成され、この絶縁層３５には酸化シリコン等が用いられている。第１電極３６は、絶縁層３５の表面上に、画素３３ａ毎に個別に形成されている。光電変換層３７は、各第１電極３６の表面上に、各画素３３ａに共通に設けられている。第２電極３８は、光電変換層３７の表面上に、各画素３３ａに共通に設けられている。第２電極３８の表面上には、前述のシンチレータ２６が接着剤（図示せず）により貼り合わされている。 The single crystal semiconductor substrate 34 is made of single crystal Si. The insulating layer 35 is formed on the surface of the single crystal semiconductor substrate 34, and silicon oxide or the like is used for the insulating layer 35. The first electrode 36 is individually formed for each pixel 33 a on the surface of the insulating layer 35. The photoelectric conversion layer 37 is provided on the surface of each first electrode 36 in common to each pixel 33 a. The second electrode 38 is provided on the surface of the photoelectric conversion layer 37 in common to each pixel 33 a. The aforementioned scintillator 26 is bonded onto the surface of the second electrode 38 by an adhesive (not shown).
　第２電極３８は、シンチレータ２６で発生した可視光が光電変換層３７に入射するように、可視光に対して透明な導電性材料（例えば、酸化インジウムスズ（ＩＴＯ））で形成されている。本実施形態では、第２電極３８を各画素３３ａに共通に設けているが、画素３３ａ毎に個別に設けてもよい。 The second electrode 38 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 26 is incident on the photoelectric conversion layer 37. In the present embodiment, the second electrode 38 is provided in common to each pixel 33a, but may be provided individually for each pixel 33a.
　光電変換層３７は、シンチレータ２６との組み合わせにより、Ｘ線の入射量に応じた信号電荷を発生する。光電変換層３７は、シンチレータ２６により発生された可視光を吸収して、吸収した光量に応じた信号電荷を発生するものであり、有機又は無機の光電変換材料で形成されている。無機の光電変換材料としては、例えば、アモルファスシリコン（ａ－Ｓｉ）がある。有機の光電変換材料としては、例えば、キナクリドンがある。 The photoelectric conversion layer 37 generates a signal charge according to the amount of incident X-rays, in combination with the scintillator 26. The photoelectric conversion layer 37 absorbs visible light generated by the scintillator 26, generates a signal charge according to the amount of absorbed light, and is formed of an organic or inorganic photoelectric conversion material. Examples of the inorganic photoelectric conversion material include amorphous silicon (a-Si). An example of the organic photoelectric conversion material is quinacridone.
　図６に示すように、キナクリドンからなる有機光電変換材料（ＯＰＣ）の感度は、ＣｓＩ：Ｎａや、単結晶Ｓｉ（ｃ－Ｓｉ）よりも、ＣｓＩ：Ｔｌが発生する可視光の波長域に近い。このため、シンチレータ２６としてＣｓＩ：Ｔｌを用いた本実施形態では、光電変換層３７をキナクリドンで形成することが好ましく、高い検出効率を得ることができる。 As shown in FIG. 6, the sensitivity of the organic photoelectric conversion material (OPC) made of quinacridone is closer to the wavelength range of visible light generated by CsI: Tl than by CsI: Na or single crystal Si (c-Si). . For this reason, in the present embodiment using CsI: Tl as the scintillator 26, it is preferable to form the photoelectric conversion layer 37 of quinacridone, and high detection efficiency can be obtained.
　単結晶半導体基板３４には、画素３３ａ毎に信号出力回路４１が設けられている。信号出力回路４１は、ＣＭＯＳ回路により形成されている。信号出力回路４１と第１電極３６との間は、コンタクト配線４２によって電気的に接続されている。第２電極３８には、バイアス電圧が印加されており（図７参照）、光電変換層３７により発生された信号電荷を各画素３３ａの第１電極３６により収集する。信号出力回路４１は、第１電極３６により収集された信号電荷を、その信号電荷量に応じた電圧信号に変換して出力する。 In the single crystal semiconductor substrate 34, a signal output circuit 41 is provided for each pixel 33a. The signal output circuit 41 is formed of a CMOS circuit. The signal output circuit 41 and the first electrode 36 are electrically connected by the contact wiring 42. A bias voltage is applied to the second electrode 38 (see FIG. 7), and the signal charge generated by the photoelectric conversion layer 37 is collected by the first electrode 36 of each pixel 33 a. The signal output circuit 41 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.
　図７において、信号出力回路４１は、出力トランジスタＴ１、行選択トランジスタＴ２、リセットトランジスタＴ３、行選択線Ｌ１、信号出力線Ｌ２、リセット線Ｌ３により構成されている。出力トランジスタＴ１、行選択トランジスタＴ２、リセットトランジスタＴ３は、それぞれＭＯＳトランジスタである。行選択線Ｌ１、信号出力線Ｌ２、リセット線Ｌ３は、前述の絶縁層３５内にアルミニウム等の金属で形成されている。 In FIG. 7, the signal output circuit 41 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 35 described above.
　出力トランジスタＴ１は、第１電極３６に接続されており、第１電極３６により収集された信号電荷に応じた電圧がゲートに印加される。行選択トランジスタＴ２は、行選択線Ｌ１に印加される選択信号よりオンとなり、出力トランジスタＴ１のゲート電圧に応じて制御された電圧信号が信号出力線Ｌ２に印加される。リセットトランジスタＴ３は、リセット線Ｌ３に印加される選択信号よりオンとなり、第１電極３６により収集された信号電荷を電源配線Ｖｄｄに破棄する。 The output transistor T1 is connected to the first electrode 36, and a voltage corresponding to the signal charge collected by the first electrode 36 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 36 to the power supply wiring Vdd.
　以上のように、ＣＭＯＳセンサ３３の単結晶半導体基板３４には、シリコンが用いられているため、各トランジスタＴ１～Ｔ３のキャリア移動度は、ａ－Ｓｉ等の非晶質半導体からなるＴＦＴに比べて、３～４桁以上高く、高速読み出しが可能である。また、単結晶半導体基板３４には、ＦＰＤ１９の制御部等の周辺回路を混載することも可能である。 As described above, since silicon is used for the single crystal semiconductor substrate 34 of the CMOS sensor 33, 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. In addition, peripheral circuits such as a control unit of the FPD 19 can be mixedly mounted on the single crystal semiconductor substrate 34.
　信号出力回路４１のうち、行選択線Ｌ１、信号出力線Ｌ２、リセット線Ｌ３は、アルミニウム等の金属で形成されているためＸ線による劣化は少ないが、出力トランジスタＴ１、行選択トランジスタＴ２、リセットトランジスタＴ３は、単結晶Ｓｉで形成されているため、Ｘ線により特性が劣化（閾値電圧の変化や暗電流の増加）する恐れがある。これは、単結晶Ｓｉを用いたＭＯＳ構造では、Ｘ線の吸収によって単結晶半導体基板３４と絶縁層３５との界面に電荷（以下、界面電荷という）が生じて蓄積されるためである。 Of the signal output circuit 41, the row selection line L1, the signal output line L2, and the reset line L3 are formed of a metal such as aluminum, so there is little deterioration due to X-rays, 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 may be degraded (change in threshold voltage or increase in dark current) due to X-rays. This is because in the MOS structure using single crystal Si, charges (hereinafter referred to as interface charge) are generated and accumulated at the interface between the single crystal semiconductor substrate 34 and the insulating layer 35 by absorption of X-rays.
　放射線吸収部２０は、天板１３を透過したＸ線の一部（ＭＯＳトランジスタの特性劣化に影響する低エネルギ成分）を吸収する。Ｘ線の高エネルギ成分は、ＣＭＯＳセンサ３３を透過するが、Ｘ線の低エネルギ成分は、ＣＭＯＳセンサ３３を透過するだけのエネルギが無く、ＣＭＯＳセンサ３３に吸収されてしまうため、各トランジスタＴ１～Ｔ３の界面電荷を増加させる可能性がある。 The radiation absorbing portion 20 absorbs a part of the X-rays transmitted through the top plate 13 (a low energy component affecting the characteristic deterioration of the MOS transistor). The high energy component of the X-ray passes through the CMOS sensor 33, but the low energy component of the X-ray does not have enough energy to pass through the CMOS sensor 33, and is absorbed by the CMOS sensor 33. There is a possibility of increasing the interface charge of T3.
　放射線吸収部２０は、袋体４７と放射線吸収性流動体４８とで構成されている。袋体４７は、被写体Ｈによる荷重により変形可能な弾力性を有する。放射線吸収性流動体４８は、袋体４７内に封入されている。袋体４７は、例えば厚みが数ｍｍ～１ｃｍ程度のゴムやプラスチック材料等の弾力性を有する材質により袋状に形成されている。放射線吸収性流動体４８は、液状またはジェル状であり、Ｘ線の一部（低エネルギ成分）を吸収する性質を有する。本実施形態では、放射線吸収性流動体４８として、Ｘ線を吸収して光に変換する液状シンチレータを用いている。 The radiation absorbing portion 20 is composed of a bag 47 and a radiation absorbing fluid 48. The bag body 47 has elasticity that can be deformed by the load of the subject H. The radiation absorbing fluid 48 is enclosed in a bag 47. The bag body 47 is formed in a bag shape, for example, of an elastic material such as rubber or plastic material having a thickness of about several mm to 1 cm. The radiation absorbing fluid 48 is liquid or gel and has a property of absorbing a part of X-rays (low energy component). In the present embodiment, a liquid scintillator that absorbs X-rays and converts them into light is used as the radiation absorbing fluid 48.
　液状シンチレータは、蛍光体からなる溶質、この溶質が混合される溶媒、界面活性剤などにより形成されている。液状シンチレータの特性は、溶媒、溶質、界面活性剤などの種類及び使用量に依存する。本実施形態では、溶媒として、例えば混合キシレン（異性体）、トルエン、プソイドクメン、ジオキサン等を用いる。また、溶質としては、例えばＰＰＯ、ｂｕｌｔｙ－ＰＢＤ、ＤＭＰＯＰＯＰ、Ｂｉｓ－ＭＳＢ等を用いる。これらの溶媒及び溶質を用い、溶質量を例えば０．５～５ｇ／リットルにして液状シンチレータを構成する。例えば、溶質としてＰＰＯを用いた場合には、４４０ｎｍ以下の発光スペクトルが得られる。なお、溶質内に波長変換材を混ぜることにより、発光スペクトルを長波側にシフトすることも可能である。 The liquid scintillator is formed of a solute composed of a phosphor, a solvent in which the solute is mixed, a surfactant, and the like. The properties of the liquid scintillator depend on the type and amount of solvent, solute, surfactant and the like. In the present embodiment, for example, mixed xylene (isomer), toluene, pseudocumene, dioxane or the like is used as the solvent. Further, as the solute, for example, PPO, bulty-PBD, DMPOPOP, Bis-MSB or the like is used. Using these solvents and solutes, the solution mass is, for example, 0.5 to 5 g / liter to construct a liquid scintillator. For example, when PPO is used as a solute, an emission spectrum of 440 nm or less is obtained. It is also possible to shift the emission spectrum to the long wave side by mixing a wavelength conversion material in the solute.
　光吸収層２１は、放射線吸収部２０によってＸ線から変換された光を吸収することにより、放射線吸収部２０の光がセンサパネル２５に入射することを防止する。これにより、放射線画像の画質が防止される。光吸収層２１は、液状シンチレータで発光した光を吸収可能な材質によって形成されている。 The light absorbing layer 21 absorbs the light converted from the X-rays by the radiation absorbing unit 20 to prevent the light of the radiation absorbing unit 20 from being incident on the sensor panel 25. This prevents the image quality of the radiation image. The light absorption layer 21 is formed of a material capable of absorbing light emitted by the liquid scintillator.
　図８に示すように、被写体Ｈの撮影部位が手の場合には、撮影部位がＦＰＤ１９の撮影範囲よりも小さい。また、このような場合には、撮影部位はＦＰＤ１９の撮影範囲の中央に配置された状態で撮影されることがほとんどである。このため、撮影部位が配置される中央部以外の周縁部は、Ｘ線が直接ＦＰＤ１９に照射される素抜け領域となり、ＣＭＯＳセンサ３３において特性劣化が発生しやすい。したがって、ＦＰＤ１９に入射するＸ線は、放射線画像の画質に影響する中央部では放射線吸収部２０によるＸ線の吸収量が少なく、一方の素抜け領域にあたる周縁部では放射線吸収部２０によるＸ線の吸収量が多いことが好ましい。 As shown in FIG. 8, when the imaging region of the subject H is a hand, the imaging region is smaller than the imaging range of the FPD 19. Moreover, in such a case, the imaging region is mostly imaged in a state of being disposed at the center of the imaging range of the FPD 19. For this reason, the peripheral portion other than the central portion where the imaging region is arranged becomes a blank area where the X-rays are directly irradiated to the FPD 19, and characteristic deterioration easily occurs in the CMOS sensor 33. Therefore, the X-ray incident on the FPD 19 has a small amount of X-ray absorption by the radiation absorbing portion 20 in the central portion that affects the image quality of the radiation image, and in the peripheral portion corresponding to one missing region Preferably, the amount of absorption is large.
　放射線吸収部２０は、被写体Ｈの体重によって変形可能である。具体的には、図９に示すように、放射線吸収部２０の中央部に被写体Ｈが載置された場合には、袋体４７がその重さによって変形し、放射線吸収性流動体４８が袋体４７の変形に合わせて袋体４７内で流動する。これにより、放射線吸収部２０は、被写体Ｈの直下である中央部の厚みが薄くなり、周縁部の厚みが厚くなる。 The radiation absorbing unit 20 can be deformed by the weight of the subject H. Specifically, as shown in FIG. 9, when the subject H is placed at the central portion of the radiation absorbing unit 20, the bag 47 is deformed by its weight, and the radiation absorbing fluid 48 is a bag. It flows in the bag 47 in accordance with the deformation of the body 47. As a result, in the radiation absorbing unit 20, the thickness of the central portion immediately below the subject H is reduced, and the thickness of the peripheral portion is increased.
　一般に、物体を透過した後のＸ線強度Ｉは、下記式（１）で示され、減弱係数μが一定であるときには、物体の厚みＴが増加すると指数的に低下する。したがって、放射線吸収部２０は、図９に示すように変形した場合には、周縁部のＸ線吸収量が高くなるため、素抜け領域におけるＣＭＯＳセンサ３３の特性劣化が抑制される。また、放射線吸収部２０の中央部におけるＸ線の吸収量は低下し、被写体Ｈが天板１３と近接した状態となるので、放射線吸収部２０によるＸ線散乱もほとんど生じることがなく、放射線画像の画質低下が抑制される。 In general, the X-ray intensity I after passing through the object is represented by the following equation (1), and decreases exponentially as the thickness T of the object increases when the attenuation coefficient μ is constant. Therefore, when the radiation absorbing portion 20 is deformed as shown in FIG. 9, the X-ray absorption amount at the peripheral portion becomes high, so that the characteristic deterioration of the CMOS sensor 33 in the blank area is suppressed. Further, the amount of absorption of X-rays at the central portion of the radiation absorbing portion 20 is reduced, and the subject H is in a state of being close to the top 13, so X-ray scattering by the radiation absorbing portion 20 hardly occurs. Image quality deterioration is suppressed.
　Ｉ＝Ｉ’ｅｘｐ（－μＴ）・・・（１）
　Ｉ：透過後のＸ線強度
　Ｉ’：透過前のＸ線強度
　μ：減弱係数（物質の種類、密度で異なる）
　Ｔ：透過厚み I = I'exp (-μT) (1)
I: X-ray intensity after transmission I ': X-ray intensity before transmission μ: Attenuation coefficient (different according to the type of substance and density)
T: Transmission thickness
　上述したように、被写体Ｈの撮影を行う際には、天板１３に被写体Ｈ自身の荷重が加わる。ＣＭＯＳセンサ３３の単結晶半導体基板３４は、材質的に割れやすく、かつ厚みが数十μｍ程度と薄いため、天板１３に加わった被写体Ｈの荷重による破損を防止する保護構造を有することが好ましい。特に、４枚のＣＭＯＳセンサ３３で構成されたセンサパネル２５では、センサパネル２５の中央部に衝撃や荷重が加わると、４枚全てのＣＭＯＳセンサ３３が破損することがあり、修理コストが非常に大きい。 As described above, when the subject H is photographed, the load of the subject H itself is applied to the top 13. Since the single crystal semiconductor substrate 34 of the CMOS sensor 33 is easily broken in material and thin as about several tens of μm, it is preferable that the single crystal semiconductor substrate 34 have a protective structure that prevents damage due to the load of the object H applied to the top plate 13 . In particular, in the sensor panel 25 configured of four CMOS sensors 33, when an impact or load is applied to the central portion of the sensor panel 25, all four CMOS sensors 33 may be damaged, and the repair cost is very high. large.
　本実施形態では、放射線吸収部２０が被写体Ｈからの荷重によって変形することにより、天板１３の中央部に加わった荷重や衝撃を吸収するので、ＣＭＯＳセンサ３３の破損を防止することができる。また、放射線吸収部２０により、被写体Ｈの熱がセンサパネル２５に伝達されるのを抑制することができ、センサパネル２５の温度むらによる放射線画像の画質劣化も抑制することができる。 In the present embodiment, the radiation absorbing portion 20 is deformed by the load from the subject H to absorb the load and the impact applied to the central portion of the top 13, so that the CMOS sensor 33 can be prevented from being damaged. In addition, the radiation absorbing unit 20 can suppress the heat of the subject H from being transmitted to the sensor panel 25, and can also suppress the image quality deterioration of the radiation image due to the temperature unevenness of the sensor panel 25.
　図１０において、放射線画像撮影装置７の電気回路は、センサパネル２５、信号処理部５０、画像メモリ５１、制御部５２、無線通信部５３、バッテリ５４等で構成されている。信号処理部５０は、センサパネル２５の各画素３３ａから出力された画素信号を増幅するアンプや、Ａ／Ｄ（アナログ／デジタル）変換器等を備えており、センサパネル２５から出力された画素信号をデジタルの画像データに変換する。 In FIG. 10, the electric circuit of the radiation image capturing apparatus 7 includes a sensor panel 25, a signal processing unit 50, an image memory 51, a control unit 52, a wireless communication unit 53, a battery 54 and the like. The signal processing unit 50 includes an amplifier for amplifying a pixel signal output from each pixel 33 a of the sensor panel 25, an A / D (analog / digital) converter, and the like, and the pixel signal output from the sensor panel 25 Into digital image data.
　また、信号処理部５０は、放射線吸収部２０によるＸ線の低エネルギ成分の吸収分布に応じて放射線画像を補正する画像補正部５０ａを備えている。撮影部位は、中央部に載せられることが殆どであるから、撮影時には、放射線吸収部２０の中央部が薄くなり、周縁部と比べてＸ吸収量が少ない。中央部の凹み状態は、撮影部位によって決まる。そこで、撮影部位に対する各部の厚みを実験的に求める。そして、この厚み量に応じた各画素または領域毎のＸ線吸収量を求めるための補正係数を画像補正部５０ａにセットしておく。こうすることにより、コンソール８で撮影部位を指定することによって、画像補正部５０ａは、補正係数を用いて画像データを補正することにより、放射線吸収部２０のＸ線吸収量による影響を取り除く。なお、性別または体重と身長を入力することにより、補正の精度を向上させることができる。 The signal processing unit 50 further includes an image correction unit 50 a that corrects a radiation image according to the absorption distribution of low energy components of X-rays by the radiation absorption unit 20. Since the region to be imaged is mostly placed at the central portion, the central portion of the radiation absorbing unit 20 becomes thinner at the time of imaging, and the X absorption amount is smaller than that at the peripheral portion. The state of depression at the center depends on the imaging site. Therefore, the thickness of each part with respect to the imaging site is determined experimentally. Then, a correction coefficient for obtaining an X-ray absorption amount for each pixel or area corresponding to the thickness amount is set in the image correction unit 50a. By doing this, the image correction unit 50a corrects the image data using the correction coefficient by designating the imaging region with the console 8, thereby removing the influence of the X-ray absorption amount of the radiation absorption unit 20. The accuracy of the correction can be improved by inputting the gender or the weight and the height.
　信号処理部５０には画像メモリ５１が接続されており、信号処理部５０の画像補正部５０ａから出力された画像データは画像メモリ５１に記憶される。画像メモリ５１は複数フレーム分の画像データを記憶可能な記憶容量を有している。放射線画像の撮影が行われる毎に、撮影によって得られた画像データが画像メモリ５１に順次記憶される。 An image memory 51 is connected to the signal processing unit 50, and the image data output from the image correction unit 50 a of the signal processing unit 50 is stored in the image memory 51. The image memory 51 has a storage capacity capable of storing image data of a plurality of frames. Image data obtained by imaging is sequentially stored in the image memory 51 each time a radiographic image is captured.
　制御部５２は、ＣＰＵ５２ａ、ＲＡＭ５２ｂ、ＲＯＭ５２を有し、放射線画像撮影装置７の全体の動作を制御する。ＲＡＭ５２ｂは、ＤＲＡＭ等からなる一時記憶メモリである。ＲＯＭ５２は、フラッシュメモリ等からなる不揮発性メモリである。 The control unit 52 includes a CPU 52a, a RAM 52b, and a ROM 52, and controls the overall operation of the radiation imaging device 7. The RAM 52 b is a temporary storage memory made of a DRAM or the like. The ROM 52 is a non-volatile memory including a flash memory or the like.
　無線通信部５３は、ＩＥＥＥ（Institute of Electrical and Electronics
Engineers）８０２．１１ａ／ｂ／ｇ／ｎ等に代表される無線ＬＡＮ（Local Area Network）規格に対応しており、外部機器との間での各種情報の無線通信を可能とする。制御部５２は、無線通信部５３を介してコンソール８と無線通信を行い、コンソール８との間で各種情報の送受信を行う。 The wireless communication unit 53 is an IEEE (Institute of Electrical and Electronics).
Engineers) correspond to wireless LAN (Local Area Network) standards represented by 802.11a / b / g / n etc., and enable wireless communication of various information with external devices. The control unit 52 performs wireless communication with the console 8 via the wireless communication unit 53, and transmits and receives various types of information to and from the console 8.
　バッテリ５４は、放射線画像撮影装置７内の各部に電力を供給する。バッテリ５４は、充電可能な二次電池であり、また、放射線画像撮影装置７から着脱可能である。なお、信号処理部５０、画像メモリ５１、制御部５２、無線通信部５３を、駆動回路基板２９内に設けることも可能である。これらはバスを介して互いに接続されている。 The battery 54 supplies power to each part in the radiation imaging device 7. The battery 54 is a rechargeable secondary battery, and is detachable from the radiation imaging device 7. The signal processing unit 50, the image memory 51, the control unit 52, and the wireless communication unit 53 may be provided in the drive circuit board 29. These are connected to one another via a bus.
　図１１において、コンソール８は、ＣＰＵ５７、ＲＯＭ５８、ＲＡＭ５９、ＨＤＤ６０を有しており、これらはバス６７を介して互いに接続されている。ＣＰＵ５７は、コンソール８の各部を制御する。ＲＯＭ５８は、制御プログラムを含む各種プログラムを記憶している。ＲＡＭ５９は、各種データを一時的に記憶する。ＨＤＤ６０は、各種データを記憶する。 In FIG. 11, the console 8 includes a CPU 57, a ROM 58, a RAM 59, and an HDD 60, which are connected to one another via a bus 67. The CPU 57 controls each part of the console 8. The ROM 58 stores various programs including a control program. The RAM 59 temporarily stores various data. The HDD 60 stores various data.
　また、バス６７には、通信Ｉ／Ｆ６１、無線通信部６２、ディスプレイドライバ６４、操作入力検出部６６が接続されている。ディスプレイドライバ６４には、ディスプレイ６３が接続されている。操作入力検出部６６には、操作パネル６５が接続されている。 A communication I / F 61, a wireless communication unit 62, a display driver 64, and an operation input detection unit 66 are connected to the bus 67. A display 63 is connected to the display driver 64. An operation panel 65 is connected to the operation input detection unit 66.
　通信Ｉ／Ｆ６１は、接続端子６１ａ、通信ケーブル６９及び放射線発生装置６の接続端子７０ａを介して、放射線発生装置６の通信Ｉ／Ｆ７０と接続されている。コンソール８のＣＰＵ５７は、放射線発生装置６との間での曝射条件等の各種情報の送受信を通信Ｉ／Ｆ６１経由で行う。無線通信部６２は、放射線画像撮影装置７の無線通信部５３と無線通信を行う機能を備えている。コンソール８のＣＰＵ５７は、放射線画像撮影装置７との間で、画像データ等の各種情報の送受信を、無線通信部６２を介して行う。 The communication I / F 61 is connected to the communication I / F 70 of the radiation generation device 6 via the connection terminal 61 a, the communication cable 69, and the connection terminal 70 a of the radiation generation device 6. The CPU 57 of the console 8 performs transmission and reception of various information such as the irradiation condition with the radiation generator 6 via the communication I / F 61. The wireless communication unit 62 has a function of performing wireless communication with the wireless communication unit 53 of the radiation image capturing apparatus 7. The CPU 57 of the console 8 exchanges various information such as image data with the radiation imaging device 7 via the wireless communication unit 62.
　ディスプレイドライバ６４は、ディスプレイ６３への各種情報を表示させるための信号を生成して出力する。コンソール８のＣＰＵ５７は、操作メニューや撮影された放射線画像等を、ディスプレイドライバ６４を介してディスプレイ６３に表示させる。また、操作パネル６５は、複数のキーを有しており、各種の情報や操作指示を入力可能とする。操作入力検出部６６は、操作パネル６５に対してなされた操作を検出し、検出結果をＣＰＵ５７に送信する。 The display driver 64 generates and outputs a signal for displaying various information on the display 63. The CPU 57 of the console 8 causes the display 63 to display the operation menu, the captured radiographic image, and the like via the display driver 64. Further, the operation panel 65 has a plurality of keys, and can input various information and operation instructions. The operation input detection unit 66 detects an operation performed on the operation panel 65, and transmits the detection result to the CPU 57.
　放射線発生装置６は、通信Ｉ／Ｆ７０と、線源制御部７２と、放射線源６ａとで構成されている。通信Ｉ／Ｆ７０は、コンソール８との間で曝射条件等の各種情報の送受信を行う。線源制御部７２は、コンソール８から受信した曝射条件（管電圧、管電流の情報を含む）に基づいて放射線源６ａを制御する。 The radiation generating apparatus 6 is configured of a communication I / F 70, a radiation source control unit 72, and a radiation source 6a. The communication I / F 70 transmits and receives various information such as an irradiation condition to and from the console 8. The radiation source control unit 72 controls the radiation source 6 a based on the irradiation conditions (including the information of the tube voltage and the tube current) received from the console 8.
　次に、放射線撮影システム５の作用を説明する。まず、撮影者（例えば放射線技師等）は、被写体Ｈの撮影部位と撮影台との間に、照射面１１側を撮影部位に向けた状態で放射線画像撮影装置７を挿入する。次に、放射線画像撮影装置７の向きや位置等を細かく調整する。 Next, the operation of the radiation imaging system 5 will be described. First, the photographer (for example, a radiographer or the like) inserts the radiation imaging device 7 between the imaging region of the subject H and the imaging table with the irradiation surface 11 facing the imaging region. Next, the orientation, position, and the like of the radiation imaging device 7 are finely adjusted.
　図８に示すように、放射線画像撮影装置７の上に撮影部位が直接載置される場合には、放射線画像撮影装置７には撮影部位を載置する際の衝撃や荷重が加わるが、この衝撃や荷重は、放射線吸収部２０が変形することによって吸収される。本実施形態の放射線画像撮影装置７は、ＩＳＳ方式であり、センサパネル２５がシンチレータ２６の天板１３側に配置されているため、センサパネル２５の破損が効果的に防止される。 As shown in FIG. 8, when the imaging region is directly placed on the radiation imaging device 7, the radiation imaging device 7 is subjected to an impact or a load when mounting the imaging region. The impact and load are absorbed by deformation of the radiation absorbing portion 20. The radiation image capturing apparatus 7 of the present embodiment is an ISS method, and the sensor panel 25 is disposed on the top plate 13 side of the scintillator 26, so that breakage of the sensor panel 25 is effectively prevented.
　撮影者は、準備作業が完了すると、操作パネル６５を操作して撮影開始を指示する。これにより、コンソール８では、曝射開始を指示する指示信号を放射線発生装置６へ送信し、放射線発生装置６は放射線源６ａからＸ線を射出させる。放射線源６ａから射出されたＸ線は、線源フィルタ６ｂにより低エネルギ成分が吸収され、被写体Ｈの撮影部位を透過し、放射線吸収部２０及び光吸収層２１を介して照射面１１に照射される。そして、Ｘ線は、天板１３及びセンサパネル２５を透過して、シンチレータ２６に入射する。 When the preparation work is completed, the photographer operates the operation panel 65 to instruct start of imaging. Thus, the console 8 transmits an instruction signal instructing the start of exposure to the radiation generator 6, and the radiation generator 6 causes the radiation source 6a to emit X-rays. The low energy component is absorbed by the radiation source filter 6b, the X-ray emitted from the radiation source 6a is transmitted through the imaging region of the subject H, and is irradiated to the irradiation surface 11 through the radiation absorbing portion 20 and the light absorbing layer 21. Ru. Then, the X-rays pass through the top plate 13 and the sensor panel 25 and enter the scintillator 26.
被写体Ｈを透過したＸ線は、放射線吸収部２０により低エネルギ成分が吸収されるので、センサパネル２５がＸ線を吸収することによる各ＣＭＯＳセンサ３３の特性劣化が抑制される。特に、放射線吸収部２０は、被写体Ｈの重さによって変形することにより、中央部よりも周縁部の厚みが厚いため、周縁部（素抜け領域）にＸ線が直接照射されることによる特性劣化が効果的に抑制される。また、放射線吸収部２０は、中央部の厚みが周縁部よりも薄いため、被写体Ｈの撮影部位を透過したＸ線が必要以上に吸収されることがなく、放射線画像の画質を大きく劣化させることはない。 The low energy component of the X-rays transmitted through the object H is absorbed by the radiation absorbing unit 20, so that the characteristic deterioration of each CMOS sensor 33 due to the sensor panel 25 absorbing the X-rays is suppressed. In particular, the radiation absorbing portion 20 is deformed by the weight of the object H, and the thickness of the peripheral portion is thicker than that of the central portion. Therefore, the characteristic deterioration due to direct irradiation of the X-ray to the peripheral portion Is effectively suppressed. In addition, since the radiation absorbing portion 20 has a central portion thinner than the peripheral portion, X-rays transmitted through the imaging region of the subject H are not absorbed more than necessary, and the image quality of the radiation image is greatly degraded. There is no.
　シンチレータ２６に入射したＸ線は、シンチレータ２６のＸ線の入射面近傍、すなわちセンサパネル２５側で大部分が可視光に変換される。シンチレータ２６で発生した可視光のうち、センサパネル２５側に向かって伝搬する可視光は、センサパネル２５に入射する。また、シンチレータ２６で発生した可視光のうち、支持基板２７側に向かって伝搬した可視光は、支持基板２７の反射層により反射された後、再びシンチレータ２６を通ってセンサパネル２５に向かい、センサパネル２５に入射する。シンチレータ２６で発光した可視光は、ＣｓＩ：Ｔｌからなる柱状結晶によってガイドされるので、画像ボケが抑制される。 Most of the X-rays incident on the scintillator 26 are converted into visible light in the vicinity of the X-ray incident surface of the scintillator 26, that is, on the sensor panel 25 side. Among the visible light generated by the scintillator 26, the visible light propagating toward the sensor panel 25 is incident on the sensor panel 25. Further, among the visible light generated by the scintillator 26, the visible light propagated toward the support substrate 27 is reflected by the reflective layer of the support substrate 27 and then passes through the scintillator 26 again toward the sensor panel 25, and the sensor The light is incident on the panel 25. Since the visible light emitted by the scintillator 26 is guided by the columnar crystal made of CsI: Tl, image blurring is suppressed.
　シンチレータ２６に入射した可視光は、第２電極３８を透過して光電変換層３７に入射し、光電変換層３７で信号電荷に変換される。Ｘ線の曝射終了後、光電変換層３７で発生した信号電荷は、第１電極３６により収集され、Ｘ線の曝射終了後、信号出力回路４１によって電圧信号に変換される。この電圧信号は、画素信号として各画素３３ａから順次出力される。出力された各画素信号は、信号処理部５０により画像データに変換される。 The visible light incident on the scintillator 26 is transmitted through the second electrode 38 and incident on the photoelectric conversion layer 37, and is converted into signal charge by the photoelectric conversion layer 37. After the end of the X-ray emission, the signal charge generated in the photoelectric conversion layer 37 is collected by the first electrode 36 and converted to a voltage signal by the signal output circuit 41 after the end of the X-ray emission. This voltage signal is sequentially output from each pixel 33a as a pixel signal. The output pixel signals are converted into image data by the signal processing unit 50.
　画像データは、画像補正部５０ａにより、放射線吸収部２０の各部分の厚みによって変化するＸ線吸収量に応じて補正される。補正後の画像データは、画像メモリ５１に記憶される。ＣＰＵ５２ａは、画像メモリ５１に記憶された画像データを、無線通信部５３を介してコンソール８に送信する。コンソール８のＣＰＵ５７は、放射線画像撮影装置７から受信した画像データを、ＲＡＭ５９を介してＨＤＤ６０に記憶する。また、ＣＰＵ５７は、ディスプレイドライバ６４を介して、ＨＤＤ６０に記憶されている画像データに基づく放射線画像を、ディスプレイ６３に表示させる。 The image data is corrected by the image correction unit 50a according to the amount of X-ray absorption which changes with the thickness of each portion of the radiation absorption unit 20. The image data after correction is stored in the image memory 51. The CPU 52 a transmits the image data stored in the image memory 51 to the console 8 via the wireless communication unit 53. The CPU 57 of the console 8 stores the image data received from the radiation imaging device 7 in the HDD 60 via the RAM 59. Further, the CPU 57 causes the display 63 to display a radiation image based on the image data stored in the HDD 60 via the display driver 64.
　上述したように、放射線吸収部２０により、センサパネル２５に入射するＸ線から低エネルギ成分が吸収されるので、ＣＭＯＳセンサ３３の特性劣化が抑制される。また、放射線画像撮影装置７上に載置される被写体Ｈの位置に応じて放射線吸収部２０が変形し、低エネルギ成分の吸収量が変化するので、放射線画像の画質を劣化させることなく、素抜け領域におけるＣＭＯＳセンサ３３の特性劣化が抑制される。これらの効果は、センサパネル２５がシンチレータ２６の天板１３側に配置され、Ｘ線入射量の大きいＩＳＳ方式において、特に顕著な効果が得られる。 As described above, since the low energy components from the X-rays incident on the sensor panel 25 are absorbed by the radiation absorbing unit 20, the characteristic deterioration of the CMOS sensor 33 is suppressed. In addition, the radiation absorbing unit 20 is deformed according to the position of the subject H placed on the radiation imaging device 7, and the amount of absorption of the low energy component changes, so that the image quality of the radiation image is not deteriorated. Characteristic deterioration of the CMOS sensor 33 in the missing region is suppressed. These effects are particularly remarkable in the ISS method in which the sensor panel 25 is disposed on the top plate 13 side of the scintillator 26 and the X-ray incident amount is large.
　ＴＦＴパネルを用いた従来のＩＳＳ方式の放射線画像撮影装置は、ＴＦＴパネルの基板として用いられる無アルカリガラスのＸ線吸収が大きいため、管電圧が低いマンモグラフィには適用が容易ではなかったが、本発明によればＩＳＳ方式の放射線画像撮影装置にＸ線吸収が小さいＣＭＯＳセンサを用いることができるので、マンモグラフィへの適用が容易である。 The conventional ISS-type radiation imaging apparatus using a TFT panel is not easy to apply to mammography with a low tube voltage because the X-ray absorption of alkali-free glass used as a substrate of the TFT panel is large. According to the invention, since a CMOS sensor with low X-ray absorption can be used for a radiation imaging apparatus of ISS method, application to mammography is easy.
（その他の実施形態）
　上記実施形態では、放射線吸収部２０から放射される光を吸収するために光吸収層２１を別途設けているが、袋体４７の天板１３との貼り合せ面、または袋体４７全体に光吸収性を持たせてもよい。また、溶質の種類あるいは波長変換材を適宜選択することにより、放射線吸収性流動体４８から放射される光の波長域と、センサパネル２５が検出可能な波長域とを異ならせてもよい。これにより、光吸収層２１を省略することができる。 (Other embodiments)
In the above embodiment, the light absorbing layer 21 is separately provided to absorb the light emitted from the radiation absorbing unit 20, but the light is applied to the bonding surface of the bag 47 with the top plate 13 or the entire bag 47. It may have absorbency. The wavelength range of the light emitted from the radiation absorbing fluid 48 may be different from the wavelength range detectable by the sensor panel 25 by appropriately selecting the type of solute or the wavelength conversion material. Thereby, the light absorption layer 21 can be omitted.
　また、上記実施形態では、放射線吸収性流動体４８として液状シンチレータを用いているが、パラフィンシンチレータを用いてもよい。パラフィンシンチレータは、低温でワックス状であり、４０°Ｃ程度で流動性が生じる。パラフィンシンチレータとしては、特開平０７－３０１６７６号公報に詳しい記載があるため、詳細な説明は省略する。さらに、放射線吸収性流動体４８として、ゲル化した水ガラス（ケイ酸ナトリウム）を用いてもよい。水ガラスのゲル化は、例えば、水ガラスに酸を加えて攪拌することにより生じる。 Further, in the above embodiment, a liquid scintillator is used as the radiation absorbing fluid 48, but a paraffin scintillator may be used. The paraffin scintillator is waxy at low temperature and has fluidity at about 40 ° C. The paraffin scintillator is described in detail in Japanese Patent Application Laid-Open No. 07-301676, so the detailed description is omitted. Furthermore, as the radiation absorbing fluid 48, gelled water glass (sodium silicate) may be used. The gelation of water glass occurs, for example, by adding an acid to water glass and stirring.
　また、放射線吸収性流動体４８を、液状またはジェル状の流動体中にＸ線吸収材を分散させて構成してもよい。このＸ線吸収材としては、Ｘ線の低エネルギ成分を吸収する効果が高く後方散乱が少ない、原子番号２０～３１の金属（例えば、銅）が好ましい。 Alternatively, the radiation absorbing fluid 48 may be configured by dispersing an X-ray absorbing material in a liquid or gel fluid. As this X-ray absorbing material, metals of atomic numbers 20 to 31 (for example, copper) having a high effect of absorbing low energy components of X-rays and little back scattering are preferable.
　また、上記実施形態では、放射線吸収部２０を、袋体４７と放射線吸収性流動体４８とで構成しているが、ゴム等の弾性体で形成してもよい。さらに、この弾性体中にＸ線吸収材を分散させてもよい。このＸ線吸収材としては、同様に原子番号２０～３１の金属が好ましい。 Moreover, in the said embodiment, although the radiation absorption part 20 is comprised with the bag body 47 and the radiation absorptive fluid body 48, you may form with elastic bodies, such as rubber | gum. Furthermore, an X-ray absorber may be dispersed in this elastic body. As this X-ray absorbing material, metals of atomic numbers 20 to 31 are similarly preferable.
　また、上記実施形態では、放射線画像撮影装置７に１つの放射線吸収部２０を設けているが、複数の放射線吸収部を設けてもよい。例えば、図１２に示すように、放射線吸収部２０に加えて、天板１３とセンサパネル２５との間に、Ｘ線の低エネルギ成分を吸収する板状の低エネルギ吸収板７５を配置する。これにより、天板１３上の放射線吸収部２０が被写体Ｈの重さで変形した場合でも、ＦＰＤ１９の検出範囲全域においてＸ線の低エネルギ成分を少なくとも一定量吸収することができる。なお、低エネルギ吸収板７５は、天板１３と放射線吸収部２０との間に配置してもよいし、天板１３の代わりに低エネルギ吸収板７５を用いてもよい。 Moreover, in the said embodiment, although the one radiation absorption part 20 is provided in the radiographic imaging apparatus 7, you may provide several radiation absorption parts. For example, as shown in FIG. 12, in addition to the radiation absorbing unit 20, a plate-like low energy absorbing plate 75 for absorbing low energy components of X-rays is disposed between the top plate 13 and the sensor panel 25. Thereby, even when the radiation absorbing portion 20 on the top 13 is deformed by the weight of the subject H, the low energy component of the X-ray can be absorbed at least a constant amount in the entire detection range of the FPD 19. The low energy absorbing plate 75 may be disposed between the top plate 13 and the radiation absorbing unit 20, or the low energy absorbing plate 75 may be used instead of the top plate 13.
　低エネルギ吸収板７５の材料には、例えば、アルミニウムやガラス等を用いることができるが、上述した線源フィルタ６ｂと同じ材料を含むようにしてもよい。線源フィルタ６ｂと放射線吸収部２０とに同じ材料を用いることにより、透過するＸ線エネルギ分布が同じになる。こうすると、線源フィルタ６ｂを透過し得るＸ線が放射線吸収部２０によっても吸収されにくいので、被写体Ｈの情報を含むＸ線が放射線吸収部２０によって吸収されるのを抑制することができる。結果的にＸ線の利用効率が向上する。 As a material of the low energy absorbing plate 75, for example, aluminum, glass or the like can be used, but the same material as the above-mentioned source filter 6b may be included. By using the same material for the radiation source filter 6b and the radiation absorbing portion 20, the transmitted X-ray energy distribution becomes the same. In this case, X-rays that can pass through the radiation source filter 6b are difficult to be absorbed by the radiation absorbing unit 20 as well, so that it is possible to suppress X-rays including information of the object H from being absorbed by the radiation absorbing unit 20. As a result, the utilization efficiency of X-rays is improved.
　また、ＣＭＯＳセンサにフレキシブル性を付与するように、ＣＭＯＳセンサを、プラスチックフイルム上に形成された有機薄膜トランジスタによって構成してもよい。有機薄膜トランジスタについては、「Tsuyoshi Sekitani、Flexible organic
transistors and circuits with extreme bending stability、Nature Materials 9、平成２２年１１月７日、p.1015-1022」において詳細に説明されているので、詳しい説明は省略する。 Further, the CMOS sensor may be configured by an organic thin film transistor formed on a plastic film so as to impart flexibility to the CMOS sensor. For organic thin film transistors, see “Tsuyoshi Sekitani, Flexible organic”.
Since this is described in detail in “Transistors and circuits with extreme bending stability, Nature Materials 9, November 7, 2010, pp. 1015-1022”, the detailed description will be omitted.
　また、ＣＭＯＳセンサにフレキシブル性を付与するには、フレキシブル性を有するプラスチック基板上に、単結晶Ｓｉによって形成されたフォトダイオード及びトランジスタを配置すればよい。プラスチック基板上へのフォトダイオード及びトランジスタの配置には、例えば、数十ミクロン程度の大きさのデバイスブロックを溶液中で散布し、基板上の必要な位置に配置する技術であるＦＡＳ（Fluidic Self-Assembly）法を用いることができる。なお、ＦＳＡ法については、「前澤宏一、「Fluidic Self-Assemblyのための共鳴トンネルデバイスブロック作製技術」、電子情報通信学会技術研究報告 ED,電子デバイス、社団法人電子情報通信学会、平成２０年６月６日、１０８巻、８７号、ｐ．６７－７１」において詳細に説明されているので、詳しい説明は省略する。 Further, in order to provide the CMOS sensor with flexibility, a photodiode and a transistor formed of single crystal Si may be provided over a flexible plastic substrate. For arrangement of photodiodes and transistors on a plastic substrate, for example, a device block of about several tens of microns in size is dispersed in a solution, and FAS (Fluidic Self- Assembly) 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 May 6, 108, 87, p. Detailed description will be omitted as it is described in detail in 67-71.
　また、上記実施形態では、センサパネルをＣＭＯＳセンサで構成しているが、本発明は、単結晶半導体基板により形成されたＣＣＤイメージセンサでセンサパネルを構成した放射線画像撮影装置にも適用可能である。また、本発明は、ＩＳＳ方式に限られず、ＰＳＳ方式の放射線画像撮影装置にも適用可能である。さらに、本発明は、放射線を直接電荷に変換する直接変換型の放射線画像撮影装置にも適用が可能である。 Further, in the above embodiment, the sensor panel is constituted by a CMOS sensor, but the present invention is also applicable to a radiation image photographing apparatus in which the sensor panel is constituted by a CCD image sensor formed by a single crystal semiconductor substrate. . Further, the present invention is not limited to the ISS method, and is also applicable to a PSS type radiation imaging apparatus. Furthermore, the present invention is also applicable to a direct conversion type radiation imaging apparatus that converts radiation directly into electric charges.
　また、上記実施形態では、ＦＰＤをカセッテサイズの筐体に組み込む例について説明したが、立位型、臥位型の撮影装置や、マンモグラフィ装置に組み込むことも可能である。また、本発明は、γ線など、Ｘ線以外の放射線を使用する放射線画像撮影装置にも適用可能である。上記実施形態で説明した本発明に係る放射線画像撮影装置の構成は一例であり、本発明の主旨を逸脱しない範囲内において適宜変更可能であることは言うまでもない。 Further, in the above embodiment, an example in which the FPD is incorporated into a cassette-sized casing has been described, but it is also possible to be incorporated into a standing-type or lying-down type imaging device or a mammography device. The present invention is also applicable to a radiation imaging apparatus using radiation other than X-rays, such as γ-rays. The configuration of the radiation image capturing apparatus according to the present invention described in the above embodiment is an example, and it goes without saying that the configuration can be appropriately changed without departing from the scope of the present invention.

Claims (20)

  1. 　被写体を透過した放射線を検出して画像データを生成する放射線検出器と、
    　前記放射線検出器の放射線入射側に配置され、前記放射線検出器に入射する前記放射線の一部を吸収するとともに、被写体の荷重によって変形可能な放射線吸収部と、
    　を備えたことを特徴とする放射線画像撮影装置。     A radiation detector that detects radiation transmitted through an object and generates image data;
    A radiation absorbing portion disposed on the radiation incident side of the radiation detector, absorbing a part of the radiation incident on the radiation detector, and being deformable by a load of an object;
    A radiation image capturing apparatus comprising:
  2. 　前記放射線検出器は、単結晶半導体基板を有することを特徴とする請求の範囲第１項に記載の放射線画像撮影装置。 The radiation imaging device according to claim 1, wherein the radiation detector has a single crystal semiconductor substrate.
  3. 　前記放射線吸収部は、袋体と、この袋体に封入された液状またはジェル状の放射線吸収性流動体とを有することを特徴とする請求の範囲第２項に記載の放射線画像撮影装置。 The radiation imaging apparatus according to claim 2, wherein the radiation absorbing portion includes a bag and a liquid or gel radiation absorbing fluid sealed in the bag.
  4. 　前記放射線吸収性流動体は、前記放射線を吸収して光に変換する液状シンチレータであることを特徴とする請求の範囲第３項に記載の放射線画像撮影装置。 The radiation imaging apparatus according to claim 3, wherein the radiation absorbing fluid is a liquid scintillator which absorbs the radiation and converts it into light.
  5. 　前記放射線検出器と前記放射線吸収部との間に、前記液状シンチレータにより変換された光を吸収する光吸収層を備えることを特徴とする請求の範囲第４項に記載の放射線画像撮影装置。 5. The radiation image capturing apparatus according to claim 4, further comprising a light absorbing layer for absorbing the light converted by the liquid scintillator, between the radiation detector and the radiation absorbing unit.
  6. 　前記放射線吸収性流動体は、ゲル化された水ガラスであることを特徴とする請求の範囲第３項に記載の放射線画像撮影装置。 The radiation imaging apparatus according to claim 3, wherein the radiation absorbing fluid is gelled water glass.
  7. 　前記放射線吸収性流動体は、液状またはジェル状の流動体と、前記流動体中に分散された放射線吸収材とを有することを特徴とする請求の範囲第３項に記載の放射線画像撮影装置。 The radiation imaging apparatus according to claim 3, wherein the radiation absorbing fluid has a liquid or gel fluid and a radiation absorbing material dispersed in the fluid.
  8. 　前記放射線吸収部は、被写体からの荷重で変形する弾性体と、この弾性体中に分散された放射線吸収材とを有することを特徴とする請求の範囲第２項に記載の放射線画像撮影装置。 The radiation image capturing apparatus according to claim 2, wherein the radiation absorbing unit includes an elastic body that is deformed by a load from a subject, and a radiation absorbing material dispersed in the elastic body.
  9. 　前記放射線吸収部を複数備えることを特徴とする請求の範囲第２項に記載の放射線画像撮影装置。 The radiation imaging device according to claim 2, comprising a plurality of the radiation absorption units.
  10. 　前記放射線検出器は、前記放射線を吸収して光に変換するシンチレータと、前記シンチレータの放射線照射側に配置され、前記シンチレータにより変換された光を検出するセンサパネルとを有することとを特徴とする請求の範囲第２項に記載の放射線画像撮影装置。 The radiation detector includes a scintillator that absorbs the radiation and converts it into light, and a sensor panel that is disposed on the radiation irradiation side of the scintillator and detects light converted by the scintillator. The radiation imaging device according to claim 2.
  11. 　前記シンチレータ及び前記センサパネルを収容する筐体を有し、この筐体上に前記放射線吸収部が取り付けられていることを特徴とする請求の範囲第１０項に記載の放射線画像撮影装置。 11. The radiation image capturing apparatus according to claim 10, further comprising a case for accommodating the scintillator and the sensor panel, wherein the radiation absorbing unit is mounted on the case.
  12. 　前記放射線吸収部は、袋体と、この袋体に封入された液状シンチレータとを有することを特徴とする請求の範囲第１１項に記載の放射線画像撮影装置。 12. The radiation image capturing apparatus according to claim 11, wherein the radiation absorbing portion has a bag and a liquid scintillator enclosed in the bag.
  13. 　前記センサパネルは、前記シンチレータにより変換された光を電荷に変換する光電変換層と、前記光電変換層で変換された電荷に応じた信号を出力する信号出力回路とを含む画素を複数有し、前記信号出力回路は、前記単結晶半導体基板に設けられていることを特徴とする請求の範囲第１２項に記載の放射線画像撮影装置。 The sensor panel has a plurality of pixels including a photoelectric conversion layer that converts light converted by the scintillator into charge, and a signal output circuit that outputs a signal according to the charge converted by the photoelectric conversion layer. The radiation image photographing apparatus according to claim 12, wherein the signal output circuit is provided on the single crystal semiconductor substrate.
  14. 　前記センサパネルは、ＣＭＯＳ型のイメージセンサにより構成されていることを特徴とする請求の範囲第１３項に記載の放射線画像撮影装置。 The radiation image capturing apparatus according to claim 13, wherein the sensor panel is configured of a CMOS type image sensor.
  15. 　前記光電変換層は、アモルファスシリコンまたはキナクリドンで形成されていることを特徴とする請求の範囲第１４項に記載の放射線画像撮影装置。 The radiation image capturing apparatus according to claim 14, wherein the photoelectric conversion layer is formed of amorphous silicon or quinacridone.
  16. 　前記液状シンチレータが発生する光の波長域は、前記センサパネルが検出可能な光の波長域と異なることを特徴とする請求の範囲第４項に記載の放射線画像撮影装置。 5. The radiation image capturing apparatus according to claim 4, wherein a wavelength range of light generated by the liquid scintillator is different from a wavelength range of light detectable by the sensor panel.
  17. 　前記放射線吸収部は、前記放射線検出器よりもサイズが大きいことを特徴とする請求の範囲第２項に記載の放射線画像撮影装置。 The radiation imaging device according to claim 2, wherein the radiation absorbing unit is larger in size than the radiation detector.
  18. 　前記放射線吸収部の放射線吸収量の領域差に応じて前記画像データを補正する画像補正部を備えることを特徴とする請求の範囲第２項に記載の放射線画像撮影装置。 The radiation image photographing apparatus according to claim 2, further comprising: an image correction unit that corrects the image data according to a region difference of the radiation absorption amount of the radiation absorption unit.
  19. 　前記放射線吸収部は、前記放射線の低エネルギ成分を吸収することを特徴とする請求の範囲第２項に記載の放射線画像撮影装置。 The radiation imaging device according to claim 2, wherein the radiation absorbing unit absorbs a low energy component of the radiation.
  20. 　前記放射線の低エネルギ成分は、前記放射線のエネルギ分布の１／２以下のエネルギ成分であることを特徴とする請求の範囲第１９項に記載の放射線画像撮影装置。 20. The radiation imaging device according to claim 19, wherein the low energy component of the radiation is an energy component of 1/2 or less of the energy distribution of the radiation.
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