WO2012165155A1 - Dispositif d'imagerie radiographique - Google Patents

Dispositif d'imagerie radiographique Download PDF

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Publication number
WO2012165155A1
WO2012165155A1 PCT/JP2012/062623 JP2012062623W WO2012165155A1 WO 2012165155 A1 WO2012165155 A1 WO 2012165155A1 JP 2012062623 W JP2012062623 W JP 2012062623W WO 2012165155 A1 WO2012165155 A1 WO 2012165155A1
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WO
WIPO (PCT)
Prior art keywords
low energy
radiation
scintillator
absorbing member
sensor panel
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Ceased
Application number
PCT/JP2012/062623
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English (en)
Japanese (ja)
Inventor
俊孝 阿賀野
美広 岡田
中津川　晴康
Current Assignee (The listed assignees may be inaccurate. Google has not performed a legal analysis and makes no representation or warranty as to the accuracy of the list.)
Fujifilm Corp
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Fujifilm Corp
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Filing date
Publication date
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Publication of WO2012165155A1 publication Critical patent/WO2012165155A1/fr
Anticipated expiration legal-status Critical
Ceased legal-status Critical Current

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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
    • 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 or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/42—Arrangements for detecting radiation specially adapted for radiation diagnosis
    • A61B6/4283—Arrangements for detecting radiation specially adapted for radiation diagnosis characterised by a detector unit being housed in a cassette
    • 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 radiographic image capturing apparatus that captures a radiographic image.
  • a radiographic image capturing system that captures a subject (a patient's imaging region) using radiation (for example, X-rays) is known.
  • the radiographic imaging system includes a radiation generation apparatus that irradiates radiation and a radiographic imaging apparatus that captures a radiographic image of an imaging region.
  • Radiographic imaging apparatuses include a stationary type incorporated in a standing position imaging stand and a standing position imaging stand, and a portable type (so-called electronic cassette) that can be carried.
  • a portable radiographic imaging device can be inserted and photographed under a patient sleeping on a bed in a hospital room or the like.
  • the FPD includes a direct conversion FPD that converts radiation directly into signal charge with a conversion layer made of amorphous selenium (a-Se), etc., and indirect conversion that converts radiation once into visible light and converts visible light into signal charge.
  • a type FPD is known.
  • the indirect type FPD includes a scintillator that converts radiation into visible light, a detection panel that is disposed to face the scintillator, and an electric control circuit.
  • the detection panel has a detection surface in which a photoelectric conversion unit that generates signal charges by photoelectric conversion is formed for each pixel on an insulating substrate such as a glass substrate, and converts visible light from the scintillator into signal charges. accumulate.
  • a TFT panel in which TFTs (thin-film transistors) and photoelectric conversion portions 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 portions and MOS transistors are formed in a matrix on a silicon (Si) single crystal semiconductor substrate by a semiconductor process.
  • the MOS transistor of the CMOS sensor is formed of a single crystal semiconductor, the carrier mobility is 3 to 4 digits higher than that of a TFT panel formed of an amorphous semiconductor, and high-speed signal charge reading is possible. is there.
  • the CMOS sensor has a small variation in characteristics at the time of manufacturing the photoelectric conversion unit and the MOS transistor (for example, the threshold voltage of the MOS transistor), it is possible to obtain a high S / N image.
  • the CMOS sensor is suitable for moving image shooting and high image quality shooting.
  • CMOS sensor can now be manufactured using a 12-inch wafer and having a square side of about 200 mm. For this reason, for example, an FPD having a side of 17 inches, which is generally used for medical purposes, can be configured using four CMOS sensors.
  • JP-A-2005-249039 a scattered X-ray absorption grid that absorbs X-rays scattered by the subject is disposed between the subject and the radiographic imaging apparatus, and the outer periphery of the scattered X-ray absorption grid is The thing provided with the absorption part which absorbs the X-ray irradiated to the element
  • region is disclosed.
  • Japanese Patent Application Laid-Open No. 2010-075553 discloses disposing a filter that increases the amount of X-rays absorbed in the radiation removal region in the radiation generator.
  • the radiographic imaging device of the present invention includes a scintillator, a sensor panel, a housing, and a low energy absorbing member.
  • the scintillator converts radiation into light.
  • the sensor panel includes a photoelectric conversion layer that converts the light converted by the scintillator into electric charges and accumulates them, and a plurality of signal output circuits that output signals corresponding to the accumulated charges in units of pixels.
  • the signal output circuit is formed on a single crystal semiconductor substrate.
  • the casing includes a box-shaped casing main body that houses the scintillator and the sensor panel, and a top plate that seals the opening of the casing main body and is irradiated with radiation.
  • a low energy absorption member is arrange
  • the top plate, the low energy absorbing member, the sensor panel, and the scintillator are arranged in this order along the radiation incident direction.
  • the low energy absorbing member has a larger absorption amount of the low energy component in the peripheral portion than in the central portion.
  • the thickness of the low energy absorbing member is different between the central portion and the peripheral portion.
  • the low energy absorbing member has a recess at the center.
  • a cushioning material accommodated in the recess is provided.
  • the recess may have a shape whose depth gradually decreases from the central portion toward the peripheral portion.
  • the recess may have a shape in which the depth continuously decreases from the central portion toward the peripheral portion.
  • the low energy absorbing member may have a radiation absorbing layer that absorbs a low energy component of radiation at the peripheral portion.
  • the sensor panel is preferably composed of a CMOS image sensor.
  • the sensor panel is preferably a plurality of CMOS image sensors arranged in a rectangular shape as a whole.
  • the CMOS image sensor preferably includes a photoelectric conversion layer, a single crystal semiconductor substrate, an insulating layer, a first electrode, and a second electrode.
  • the insulating layer is formed on the surface of the single crystal semiconductor substrate.
  • the first electrodes are individually formed on the surface of the insulating layer in units of pixels.
  • the photoelectric conversion layer is provided in common for each pixel on the surface of the first electrode.
  • the second electrode is provided in common for each pixel on the surface of the photoelectric conversion layer.
  • the low energy absorbing member is preferably larger in size than the sensor panel.
  • the low energy absorbing member is preferably bonded to the top plate.
  • the low energy component of the radiation is preferably an energy component of 1/2 or less of the energy distribution of the radiation.
  • the low energy absorbing member is preferably formed of aluminum or glass.
  • the scintillator preferably has a columnar crystal structure.
  • the scintillator is preferably formed of CsI: Tl or CsI: Na.
  • the scintillator is deposited on a support substrate, and the support substrate is preferably disposed on the side opposite to the sensor panel with respect to the scintillator.
  • the low energy component is absorbed by the low energy absorbing member from the radiation applied to the sensor panel, so that the deterioration of the characteristics of the signal output circuit provided on the single crystal semiconductor substrate is suppressed. be able to.
  • the single crystal semiconductor substrate can be protected by the low energy absorbing member and can be prevented from being damaged.
  • the low energy absorbing member suppresses the deterioration of the characteristics of the signal output circuit in the peripheral portion without deteriorating the image quality of the radiographic image because the absorption amount of the low energy component in the central portion is smaller than the absorption amount in the peripheral portion. be able to.
  • a radiographic imaging system 5 includes a radiation generating device 6 that irradiates an imaging region of a subject (patient) H with X-rays as radiation, a radiographic imaging device 7 that captures a radiographic image of the subject H, and radiation generation.
  • a console 8 for controlling the apparatus 6 and the radiographic imaging apparatus 7 is provided.
  • the radiation generator 6 includes a radiation source 6a and a radiation source filter 6b.
  • the radiation source 6a includes an X-ray tube 6c that emits X-rays, and an irradiation field limiter (collimator) 6d that limits an irradiation field of the X-rays emitted by the X-ray tube 6c.
  • the X-ray tube 6c has a cathode made of a filament that emits thermoelectrons and an anode (target) that emits X-rays when the thermoelectrons emitted from the cathode collide.
  • the irradiation field limiter 6d is formed, for example, by arranging a plurality of lead plates that shield X-rays on each side of a quadrangle and forming an irradiation opening that transmits X-rays in the center, and moves the position of the lead plate By changing the size of the irradiation aperture, the irradiation field is limited.
  • the radiation source filter 6b removes low energy components that cause scattering and deterioration of the radiation image from the X-rays emitted from the radiation source 6a when passing through the imaging region.
  • a material having a property of absorbing only a low energy component of X-rays is used for the source filter 6b.
  • aluminum is suitable.
  • the high energy component of the X-ray that has passed through the radiation source filter 6 b is used for photographing the subject H.
  • the energy distribution of X-rays radiated from the X-ray tube 6c is approximately when, for example, the tube voltage of the X-ray tube 6c is 70 kV and the maximum energy of X-rays radiated 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 set as a low energy component, and 1/2 or more (40 to 70 KeV) is set as a high energy component.
  • the source filter 6b absorbs a low energy component of 15 to 40 KeV.
  • the radiographic image capturing device 7 is composed of an FPD 19, a low energy absorbing member 20, an electric circuit unit 23, and a portable housing 12 for housing them.
  • the housing 12 includes a top plate 13 and a flat box-shaped housing body 14.
  • the top plate 13 seals the opening 14 a at the top of the housing body 14.
  • the top 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 for the top board 13 on which the weight of the subject H is applied.
  • the housing body 14 is made of ABS resin or the like.
  • the housing 12 has the same size (for example, 17 inch square) as a conventional radiation film cassette that records a radiation image on a photosensitive material.
  • the radiographic imaging device 7 is portable like the radiographic film cassette, and can be used in place of the radiographic film cassette.
  • the top plate 13 of the radiation image capturing apparatus 7 is provided with a display unit 16 composed of a plurality of LEDs.
  • a display unit 16 composed of a plurality of LEDs.
  • an operation mode such as an operation mode (for example, “ready state” or “data transmitting”) of the radiographic imaging device 7 and a remaining capacity of the battery 54 is displayed.
  • you may comprise the display part 16 by light emitting elements other than LED, a liquid crystal display, an organic EL display, etc.
  • the display unit 16 may be provided in the housing body 14.
  • a low energy absorbing member 20 and an FPD 19 are laminated in order from the top plate 13 side.
  • the low energy absorbing member 20 faces the top plate 13, absorbs a low energy component of X-rays transmitted through the top plate 13, and protects the FPD 19 from a load or an impact applied to the top plate 13.
  • 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 of which are shown in FIG. 9).
  • the electric circuit portion 23 is disposed on one end side along the short direction inside the housing 12.
  • the FPD 19 is operated by electric power supplied from the battery 54.
  • a radiation shielding member such as a lead plate is provided on the top plate 13 side of the electric circuit unit 23 in order to prevent the electric circuit unit 23 from being damaged by X-rays.
  • the low energy absorbing member 20 is bonded to the entire inner surface of the top plate 13 with an adhesive (not shown).
  • the FPD 19 is bonded to the lower surface of the low energy absorbing member 20 with an adhesive (not shown).
  • the low energy absorbing member 20 and the FPD 19 are sequentially bonded to the top plate 13, whereby the radiographic imaging device 7 is thinned and the FPD 19 is reinforced by the low-energy absorbing member 20.
  • the low energy absorbing member 20 is preferably larger than the size of the FPD 19 so that the FPD 19 can be completely bonded.
  • the FPD 19 includes a sensor panel 25 and a scintillator 26, and the sensor panel 25 and the scintillator 26 are sequentially stacked from the top plate 13 side.
  • a support substrate 27 that supports the scintillator 2 is provided on the lower surface of the scintillator 26.
  • a sealant 28 is provided on the outer periphery of the FPD 19 to protect the scintillator 26 from moisture and the like.
  • a drive circuit board 29 of the FPD 19 is disposed on the bottom surface in the housing 12. The drive circuit board 29 and the sensor panel 25 are electrically connected via a flexible cable 30.
  • the scintillator 26 transmits the subject H and is irradiated onto the irradiation surface 11 of the housing 12, and absorbs X-rays that have passed through the top plate 13, the low energy absorbing member 20, and the sensor panel 25 to generate visible light.
  • CsI: Tl cesium iodide added with thallium
  • CsI: Na cesium iodide added with sodium
  • GOS Gd 2 O 2 S: Tb
  • CsI: Tl is used as the scintillator 26.
  • the scintillator 26 is formed by evaporating 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 in the scintillator 26 propagates through the columnar crystal by 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 since the scintillator 26 has a columnar crystal structure, the diffusion of visible light emitted from the scintillator 26 to the sensor panel 25 side is suppressed. Therefore, the sharpness of the radiographic image captured by the radiographic image capturing device 7 is reduced. Will improve.
  • a reflective layer (not shown) is provided on the surface of the support substrate 27 on the scintillator 26 side. Visible light emitted from the scintillator 26 and propagated to the support substrate 27 side is reflected again by the reflective layer to the sensor panel 25 side through the scintillator 26, so that the incident light amount to the sensor panel 25 (light emitted from the scintillator 26 is emitted). (Light detection efficiency) is improved.
  • the configuration in which the sensor panel is arranged on the side opposite to the X-ray incident side of the scintillator is called a PSS (PenetrationeneSide Sampling) method.
  • PSS PulenetrationeneSide Sampling
  • the sensor panel 25 includes four CMOS image sensors (hereinafter referred to as “CMOS sensors”) 33.
  • CMOS sensors CMOS 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 so as to be adjacent to each other vertically and horizontally, and form a quadrangle having a side of approximately 17 inches.
  • the CMOS sensor 33 has the same configuration as that disclosed in US Publication No. 2009/0224162. Specifically, as shown in FIG. 5, the CMOS sensor 33 includes a single crystal semiconductor substrate 34, an insulating layer 35, a first electrode 36, a photoelectric conversion layer 37, and a second electrode 38. Yes.
  • the single crystal semiconductor substrate 34 is made of single crystal Si.
  • the insulating layer 35 is formed of silicon oxide or the like on the surface of the single crystal semiconductor substrate 34.
  • the first electrode 36 is individually formed on the surface of the insulating layer 35 for each pixel 33a.
  • the photoelectric conversion layer 37 is provided in common to each pixel 33 a on the surface of each first electrode 36.
  • the second electrode 38 is provided on the surface of the photoelectric conversion layer 37 in common for each pixel 33a. On the surface of the second electrode 38, the above-mentioned scintillator 26 is bonded with an adhesive (not shown).
  • the second electrode 38 is formed of a conductive material (for example, indium tin oxide (ITO)) that is transparent to visible light so that visible light generated by the scintillator 26 enters the photoelectric conversion layer 37.
  • ITO indium tin oxide
  • the second electrode 38 is provided in common for each pixel 33a, but may be provided for each pixel 33a.
  • the photoelectric conversion layer 37 generates signal charges corresponding 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 and generates a signal charge corresponding to the amount of light, and is formed of an organic or inorganic photoelectric conversion material.
  • An example of the inorganic photoelectric conversion material is amorphous silicon (a-Si).
  • An example of an organic photoelectric conversion material is quinacridone.
  • the sensitivity of an organic photoelectric conversion material (OPC) made of quinacridone is closer to the wavelength range of visible light where CsI: Tl is generated than CsI: Na or single crystal Si (c-Si). .
  • CsI: Tl as the scintillator 26
  • an organic photoelectric conversion material having a high carrier mobility and a small manufacturing variation as the material of the photoelectric conversion layer 37.
  • a signal output circuit 41 is provided for each pixel 33a.
  • the signal output circuit 41 is formed by a CMOS circuit.
  • the signal output circuit 41 and the first electrode 36 are electrically connected by a contact wiring 42.
  • a bias voltage is applied to the second electrode 38 (see FIG. 7), and the signal charges generated by the photoelectric conversion layer 37 are collected by the first electrode 36 of each pixel 33a.
  • 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 a 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 a 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. Therefore, it is 3 to 4 digits higher and can be read at high speed.
  • peripheral circuits such as 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 that deterioration due to X-rays is small.
  • the transistor T3 is formed of single crystal Si, the characteristics may be deteriorated (threshold voltage change or dark current increase) by X-rays. This is because in a MOS structure using single crystal Si, charges (hereinafter referred to as interface charges) are generated and accumulated at the interface between the single crystal semiconductor substrate 34 and the insulating layer 35 due to the absorption of X-rays.
  • 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 low energy absorbing member 20 disposed under the top plate 13 absorbs the low energy component that affects the characteristic deterioration of the MOS transistor.
  • the low energy absorbing member 20 is a plate-like body made of a material (for example, aluminum or glass) that does not absorb much high energy components of X-rays contributing to radiography and absorbs many low energy components. is there.
  • a material for example, aluminum or glass
  • the same material is used for the radiation source filter 6b and the low energy absorbing member 20 since the X-ray absorption characteristics of both are equal, the high energy component of the X-rays transmitted through the radiation source filter 6b is reduced. Absorption of X-rays used in radiography is reduced.
  • the imaging region is a hand
  • the hand is smaller than the imaging range of the FPD 19.
  • the hand is often photographed with the hand placed in the center of the photographing range of the FPD 19.
  • the peripheral portion other than the central portion where the hand is placed becomes a blank region where X-rays are directly irradiated without passing through the imaging region, and characteristic deterioration is likely to occur in the CMOS sensor 33. Therefore, the X-rays incident on the FPD 19 have a small amount of X-ray absorption by the low energy absorbing member 20 in the central portion that affects the image quality of the radiation image, and X by the low energy absorbing member 20 in the peripheral portion corresponding to one of the blank areas. It is preferable that the amount of absorption of the wire is large.
  • the low energy absorbing member 20 has a recess 20a at the center of the surface on the top plate 13 side, and the thickness of the center is thinner than the thickness of the peripheral region. For this reason, the low energy absorbing member 20 has a larger amount of X-ray absorption in the peripheral region than in the central portion, and the characteristic deterioration of the CMOS sensor 33 in the peripheral region corresponding to the blank region is prevented. As a result, it is possible to simultaneously suppress the deterioration of the image quality of the radiation image and the characteristic deterioration of the CMOS sensor 33.
  • the subject H when performing radiography, the subject H is placed on the top board 13, and thus the load is applied. Since the single crystal semiconductor substrate 34 of the CMOS sensor 33 is easily broken in material and has a thin thickness of about several tens of ⁇ m, the single crystal semiconductor substrate 34 has a protective structure that prevents damage due to a load from the subject H applied to the top plate 13. preferable. In particular, in the sensor panel 25 composed of four CMOS sensors 33, if 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 sensor panel 25 is reinforced by the low energy absorbing member 20 disposed between the top plate 13 and the sensor panel 25, and the buffer material 20 b is accommodated in the recess 20 a of the low energy absorbing member 20. Yes.
  • the buffer material 20b absorbs the load and impact applied to the central portion of the top plate 13. Sponge, rubber, or the like is used as the buffer material 20b.
  • the radiographic imaging device 7 includes an FPD 19, 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 that amplifies the voltage signal output from each pixel 33a of the sensor panel 25, an A / D (analog / digital) converter, and the like, and the voltage signal output from the sensor panel 25. Is converted to digital image data.
  • the signal processing unit 50 includes an image correction unit 50a that corrects a radiation image according to the absorption distribution of the low energy component of the X-rays by the low energy absorbing member 20. Since the low energy absorbing member 20 has a different thickness at the central portion and the peripheral portion and has a different X absorption amount, the radiation image based on the X-rays transmitted through the low energy absorbing member 20 is affected (density difference).
  • the image correction unit 50a corrects the image data according to the X-ray absorption amount of the low energy absorption member 20, thereby removing the influence due to the region difference of the X-ray absorption amount of the low energy absorption member 20.
  • An image memory 51 is connected to the signal processing unit 50, and 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 for a plurality of frames. Each time a radiographic image is captured, the image data obtained by the imaging is sequentially stored in the image memory 51.
  • the control unit 52 includes a CPU 52a, a RAM 52b, and a ROM 52c, and controls the overall operation of the radiation image capturing apparatus 7.
  • the RAM 52b is a temporary storage memory composed of a DRAM or the like.
  • the ROM 52 is a non-volatile memory composed of a flash memory or the like.
  • the wireless communication unit 53 is an IEEE (Institute of Electrical and Electronics). Engineers) It supports wireless LAN (Local Area Network) standards represented by 802.11a / b / g / n, etc., and enables 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 / receives various information to / from the console 8.
  • the battery 54 supplies power to each part in the radiation image capturing apparatus 7.
  • the battery 54 is a rechargeable secondary battery and is detachable from the radiation image capturing apparatus 7.
  • the signal processing unit 50, the image memory 51, the control unit 52, and the wireless communication unit 53 can be provided in the drive circuit board 29. These are connected to each other via a bus.
  • the console 8 has a CPU 57, ROM 58, RAM 59, and HDD 60, which are connected to each other 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 generating apparatus 6 via the connection terminal 61a, the communication cable 69, and the connection terminal 70a of the radiation generating apparatus 6.
  • the CPU 57 of the console 8 transmits / receives various information such as the exposure conditions to / from 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 transmits and receives various kinds of information such as image data to and from the radiographic image capturing 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 displays an operation menu, a captured radiographic image, and the like on the display 63 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 a detection result to the CPU 57.
  • the radiation generator 6 includes a radiation source 6a, a communication I / F 70, and a radiation source controller 72.
  • the communication I / F 70 transmits and receives various information such as exposure conditions to and from the console 8.
  • the radiation source control unit 72 controls the radiation source 6 a based on the exposure conditions (including information on tube voltage and tube current) received from the console 8.
  • a radiographer inserts a radiographic imaging device 7 with the irradiation surface 11 facing upward between the imaging region of the subject H and the imaging table, and adjusts the orientation, position, and the like. Perform preparatory work.
  • the radiographic image capturing apparatus 7 of the present embodiment is an ISS system, and the sensor panel 25 is disposed on the top plate 13 side of the scintillator 26, so that damage to the sensor panel 25 is effectively prevented.
  • the photographer When the preparatory work is completed, the photographer operates the operation panel 65 to instruct the start of photographing. As a result, the console 8 transmits an instruction signal instructing the start of exposure to the radiation generation apparatus 6, and the radiation generation apparatus 6 emits X-rays from the radiation source 6a.
  • the X-rays emitted from the radiation source 6a are absorbed by the radiation source filter 6b and the low energy component is transmitted through the imaging region of the subject H and applied to the irradiation surface 11 of the radiographic imaging device 7. Then, the X-rays pass through the top plate 13, the low energy absorbing member 20 and the sensor panel 25 and enter the scintillator 26.
  • the low energy component of the X-rays transmitted through the top plate 13 is absorbed by the low energy absorbing member 20, the characteristic deterioration of each CMOS sensor 33 due to the sensor panel 25 absorbing the X-rays is suppressed.
  • the low energy absorbing member 20 has a thicker peripheral portion than the central portion, deterioration of characteristics due to direct irradiation of the peripheral portion (element missing region) with X-rays is effectively suppressed.
  • the thickness of the central portion of the low energy absorbing member 20 is thinner than that of the peripheral portion, X-rays transmitted through the imaging region are not absorbed more than necessary, and the image quality of the radiographic image is not greatly deteriorated. .
  • Visible light incident on the scintillator 26 passes through the second electrode 38 and enters the photoelectric conversion layer 37, and is converted into signal charges by the photoelectric conversion layer 37.
  • the signal charges generated in the photoelectric conversion layer 37 after the end of the X-ray exposure are collected by the first electrode 36 and are extracted as a voltage signal from the signal output circuit 41 after the end of the X-ray exposure.
  • This voltage signal is sequentially output from each pixel 33a.
  • the output voltage signal is converted into image data by a signal processing unit 50 having an A / D converter and a multiplexer.
  • the image data is corrected by the image correction unit 50a according to the region difference of the X-ray absorption amount of the low energy absorbing member 20.
  • the corrected image data 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 image capturing apparatus 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. Further, since the X-ray absorption amount is different between the central portion and the peripheral portion of the low energy absorbing member 20, the deterioration of the characteristics of the CMOS sensor 33 in the blank region is suppressed without deteriorating the image quality of the radiation image. . These effects are particularly remarkable in the ISS system in which the sensor panel 25 is disposed on the top plate 13 side of the scintillator 26 and the amount of incident X-rays is large.
  • the low energy absorbing member 20 is provided with the concave portion 20a having a constant depth, but the shape of the concave portion may be appropriately modified.
  • the recess 20d in the vicinity of the peripheral portion, the recess 20d has a shape that gradually decreases in depth as it approaches the peripheral portion, and as shown in FIG. 12, the depth becomes closer to the peripheral portion. You may use the recessed part 20e of the shape which becomes shallow continuously.
  • an X-ray absorption layer 20f may be provided on one surface of the low energy absorption member 20 so as to cover the peripheral portion.
  • a metal layer excellent in absorption of low energy components of X-rays is preferable.
  • a metal layer containing a metal having an atomic number of about 20 to 31 for example, copper may be used.
  • the recess 20a of the low energy absorbing member 20 is provided on the top plate 13 side, but the recess 20a may be provided on the sensor panel 25 side. Furthermore, when the image quality degradation of the radiation image due to the X-ray absorption of the low energy absorbing member 20 does not become a problem, a flat low energy absorbing member may be used without providing the recess 20a. In this case, no cushioning material is provided, but only the low energy absorbing member has an effect of protecting the sensor panel 25.
  • the low energy absorption member 20 absorbs the low energy component of the X-ray.
  • High energy components may also be absorbed before entering the CMOS sensor 33.
  • the CMOS sensor may be composed of an organic thin film transistor formed on a plastic film so as to give flexibility to the CMOS sensor.
  • organic thin film transistors see “Tsuyoshi Sekitani, Flexible organic Transistors and circuits with extreme bending stability, Nature Materials 9, November 7, 2010, p.1015-1022 ”, detailed description will be omitted.
  • a photodiode and a transistor formed of single crystal Si may be arranged on a plastic substrate having flexibility.
  • a plastic substrate for example, FAS (Fluidic Self-Self-), which is a technology that disperses device blocks of a size of several tens of microns in a solution and places them at necessary positions on the substrate. Assembly) method can be used.
  • the sensor panel was comprised with the CMOS sensor, this invention is applicable also to the radiographic imaging apparatus which comprised the sensor panel with the CCD image sensor formed of the single crystal semiconductor substrate. . Further, the present invention is not limited to the ISS system, and can also be applied to a PSS system radiographic apparatus.
  • the FPD is incorporated into a cassette-size housing.
  • the FPD may be incorporated into a standing type or a standing type photographing apparatus or a mammography apparatus.
  • the present invention is also applicable to a radiographic imaging apparatus that uses radiation other than X-rays, such as ⁇ rays. It goes without saying that the configuration of the radiographic imaging apparatus according to the present invention described in the above embodiment is an example, and can be appropriately changed without departing from the gist of the present invention.

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  • Physics & Mathematics (AREA)
  • Health & Medical Sciences (AREA)
  • Life Sciences & Earth Sciences (AREA)
  • General Physics & Mathematics (AREA)
  • High Energy & Nuclear Physics (AREA)
  • Molecular Biology (AREA)
  • Spectroscopy & Molecular Physics (AREA)
  • Measurement Of Radiation (AREA)
  • Apparatus For Radiation Diagnosis (AREA)

Abstract

L'invention concerne un dispositif d'imagerie radiographique (7) dans lequel un élément absorbeur de basse énergie (20), un panneau de capteurs (25) et un scintillateur (26) sont disposés, dans l'ordre listé, sur le côté interne d'un plateau supérieur (13) exposé à des rayons X. Le panneau de capteurs (25) est conçu à partir d'une pluralité de capteurs CMOS (33) dans lesquels des circuits de sortie de signal (41) sont formés sur des substrats composés de Si monocristallin. A partir des rayons X qui ont traversé la couche isolante (13), l'élément absorbeur de basse énergie (20) absorbe une composante à basse énergie qui cause la détérioration des caractéristiques des circuits de sortie de signal (41). L'élément absorbeur de basse énergie (20) comprend une partie concave (20a) en son centre. Un élément d'amortissement (20b) est monté dans la partie concave (20a).
PCT/JP2012/062623 2011-05-31 2012-05-17 Dispositif d'imagerie radiographique Ceased WO2012165155A1 (fr)

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JP2011-121966 2011-05-31
JP2011121966A JP2012247401A (ja) 2011-05-31 2011-05-31 放射線撮影装置

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Cited By (3)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
CN109567836A (zh) * 2017-09-29 2019-04-05 通用电气公司 X射线检测器结构
EP3502749A1 (fr) * 2017-12-22 2019-06-26 Fujifilm Corporation Dispositif de détection de rayonnements
US12414746B2 (en) * 2021-10-25 2025-09-16 Canon Kabushiki Kaisha Radiographing apparatus

Families Citing this family (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JP6264723B2 (ja) * 2013-01-23 2018-01-24 コニカミノルタ株式会社 放射線画像撮影装置
JP6478538B2 (ja) 2014-09-10 2019-03-06 キヤノン株式会社 放射線撮像装置および放射線撮像システム
JP2016104116A (ja) * 2014-11-19 2016-06-09 富士フイルム株式会社 放射線画像撮影装置
JP6502731B2 (ja) * 2015-04-13 2019-04-17 キヤノン株式会社 放射線撮像装置及び放射線撮像システム
JP7370950B2 (ja) 2020-09-28 2023-10-30 富士フイルム株式会社 放射線画像撮影装置

Citations (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JP2002214352A (ja) * 2001-01-19 2002-07-31 Canon Inc 放射線画像撮影装置
JP2003035781A (ja) * 2001-07-24 2003-02-07 Canon Inc 放射線画像撮影装置
JP2005506552A (ja) * 2001-10-26 2005-03-03 トリクセル　エス．アー．エス． 固体ｘ線検出器
JP2005114518A (ja) * 2003-10-07 2005-04-28 Canon Inc 放射線検出装置及びその製造方法
JP2010262134A (ja) * 2009-05-07 2010-11-18 Fujifilm Corp 放射線検出装置及び放射線画像撮影システム

Patent Citations (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JP2002214352A (ja) * 2001-01-19 2002-07-31 Canon Inc 放射線画像撮影装置
JP2003035781A (ja) * 2001-07-24 2003-02-07 Canon Inc 放射線画像撮影装置
JP2005506552A (ja) * 2001-10-26 2005-03-03 トリクセル　エス．アー．エス． 固体ｘ線検出器
JP2005114518A (ja) * 2003-10-07 2005-04-28 Canon Inc 放射線検出装置及びその製造方法
JP2010262134A (ja) * 2009-05-07 2010-11-18 Fujifilm Corp 放射線検出装置及び放射線画像撮影システム

Cited By (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
CN109567836A (zh) * 2017-09-29 2019-04-05 通用电气公司 X射线检测器结构
EP3502749A1 (fr) * 2017-12-22 2019-06-26 Fujifilm Corporation Dispositif de détection de rayonnements
CN109959959A (zh) * 2017-12-22 2019-07-02 富士胶片株式会社 放射线检测装置
US10732308B2 (en) 2017-12-22 2020-08-04 Fujifilm Corporation Radiation detection device
US12414746B2 (en) * 2021-10-25 2025-09-16 Canon Kabushiki Kaisha Radiographing apparatus

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