WO2015016205A1 - 低エネルギｘ線画像形成装置及びその画像の形成方法 - Google Patents

低エネルギｘ線画像形成装置及びその画像の形成方法 Download PDF

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WO2015016205A1
WO2015016205A1 PCT/JP2014/069906 JP2014069906W WO2015016205A1 WO 2015016205 A1 WO2015016205 A1 WO 2015016205A1 JP 2014069906 W JP2014069906 W JP 2014069906W WO 2015016205 A1 WO2015016205 A1 WO 2015016205A1
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ray
energy
kev
image forming
soft tissue
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English (en)
French (fr)
Japanese (ja)
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吉衞 小寺
勉 山河
山本　修一郎
義治 小幡
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Nagoya University NUC
Job Corp
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Nagoya University NUC
Job Corp
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Priority to US14/908,178 priority Critical patent/US20160174922A1/en
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    • 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
    • A61B6/50—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment specially adapted for specific body parts; specially adapted for specific clinical applications
    • A61B6/502—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment specially adapted for specific body parts; specially adapted for specific clinical applications for diagnosis of breast, i.e. mammography
    • 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
    • A61B6/02—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/025—Tomosynthesis
    • 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
    • A61B6/02—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/03—Computed tomography [CT]
    • A61B6/032—Transmission computed tomography [CT]
    • 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
    • A61B6/40—Arrangements for generating radiation specially adapted for radiation diagnosis
    • A61B6/4035—Arrangements for generating radiation specially adapted for radiation diagnosis the source being combined with a filter or grating
    • 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
    • A61B6/4208—Arrangements for detecting radiation specially adapted for radiation diagnosis characterised by using a particular type of detector
    • A61B6/4233—Arrangements for detecting radiation specially adapted for radiation diagnosis characterised by using a particular type of detector using matrix detectors
    • 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
    • A61B6/4266—Arrangements for detecting radiation specially adapted for radiation diagnosis characterised by using a plurality of detector units
    • 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
    • A61B6/44—Constructional features of apparatus for radiation diagnosis
    • A61B6/4429—Constructional features of apparatus for radiation diagnosis related to the mounting of source units and detector units
    • A61B6/4435—Constructional features of apparatus for radiation diagnosis related to the mounting of source units and detector units the source unit and the detector unit being coupled by a rigid structure
    • A61B6/4441—Constructional features of apparatus for radiation diagnosis related to the mounting of source units and detector units the source unit and the detector unit being coupled by a rigid structure the rigid structure being a C-arm or U-arm
    • 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
    • A61B6/48—Diagnostic techniques
    • A61B6/482—Diagnostic techniques involving multiple energy imaging
    • 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
    • A61B6/52—Devices using data or image processing specially adapted for radiation diagnosis
    • A61B6/5205—Devices using data or image processing specially adapted for radiation diagnosis involving processing of raw data to produce diagnostic data
    • 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
    • A61B6/06—Diaphragms
    • 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
    • A61B6/4291—Arrangements for detecting radiation specially adapted for radiation diagnosis the detector being combined with a grid or grating
    • 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
    • A61B6/52—Devices using data or image processing specially adapted for radiation diagnosis
    • A61B6/5258—Devices using data or image processing specially adapted for radiation diagnosis involving detection or reduction of artifacts or noise

Definitions

  • the present invention relates to a low-energy X-ray image forming apparatus that uses X-rays to image a soft tissue (soft tissue) to be imaged and a substance having characteristics corresponding to the soft tissue with respect to X-rays, and formation of the image. More particularly, the present invention relates to a low-energy X-ray image forming apparatus using X-rays in which the X-ray energy range is optimized based on the X-ray transmission characteristics of the soft tissue, and an image forming method thereof.
  • the breast cancer screening mainly uses palpation and mammography, but an ultrasonic device is also effective. Further, as a more detailed diagnostic method, MRI, CT, biopsy (biopsy) or the like is used. Among these, X-ray mammography is considered to be the simplest and most effective method for early detection.
  • this X-ray mammography used a high-sharp and high-contrast intensifying screen film system to detect minute calcifications and low-contrast tumors.
  • digital image systems using CR (computed radiography) and FPD (flat panel detector) have become mainstream due to advances in digital technology.
  • the energy generated from the X-ray generator is 10 This is dealt with by devising to keep it in the range of ⁇ 20keV.
  • This device uses molybdenum (Mo), which has characteristic X-rays at 17.5 keV and 19.6 keV, as a target (anode) material for an X-ray tube, and has a low energy component around 10 keV that greatly affects skin exposure, A molybdenum filter is used to suppress a component of 20 keV or more that causes a decrease in contrast.
  • Mo molybdenum
  • the current X-ray mammography can be said to be an imaging technique weighted to the characteristics of 17.5 keV and 19.6 keV, which are characteristic X-rays of molybdenum.
  • X-rays are taken of relatively light substances and small animals such as small fish and insects.
  • X-rays of about 10 keV to 25 keV are used, sufficient image quality cannot be obtained unless the X-ray tube current is increased or imaging time is secured to some extent.
  • contrast, sharpness, and noise characteristics which are the three elements of image quality, can be considered independently.
  • the main tissues of the breast, mammary gland and fat have an energy of 20 keV and linear attenuation coefficients of 0.8 cm-1 and 0.45 cm-1, respectively.
  • the linear attenuation coefficients of the mass and microcalcifications which are the main tissues of breast cancer, are 0.85 cm-1 and 1.45 cm-1, respectively, with the same energy, while the latter has a large contrast to the mammary tissue.
  • the contrast of the former is slight. Therefore, the X-ray energy used to keep this slight contrast large is low, and the film has a large contrast.
  • the tube voltage is set to 28 kV, Mo (molybdenum) is used as the target material, Mo is used as the filter material, while thick breasts are used.
  • Mo mobdenum
  • the tube voltage is set to 32kV
  • Mo is used as the target material
  • Rh rhodium
  • the remaining 99.5% is absorbed (exposed) in the body in the former and 98.6% in the latter, and does not contribute to the image at all. This means that most of the X-ray irradiation energy is exposed to the patient, and in particular, the side breast surface exposure to which X-rays are incident is very large.
  • the size of the quantum motor which is noise in general image systems, is 1 / ⁇ n of the average number n of X-ray quanta absorbed by the detector, X-rays absorbed by the detector The smaller the number of quanta, the greater the image noise. Therefore, the noise under these photographing conditions becomes very large, the effect of contrast obtained by lowering the energy is impaired, and the signal-to-noise ratio (SNR) or the contrast-to-noise ratio (CNR) is remarkably lowered.
  • SNR signal-to-noise ratio
  • CNR contrast-to-noise ratio
  • the X-ray tube current must be increased or the X-ray irradiation time must be increased.
  • the X-ray exposure dose to the breast increases. That is, a fine image and an X-ray exposure dose have a trade-off relationship.
  • the noise characteristics of the current detector system described above will be described. Regardless of whether the current digital type detector is a direct conversion method or an indirect conversion method, most of the methods obtain an output from the detector by integrating the X-ray dose for a certain period of time. In such an integral type signal detection method, integration is performed including electric noise generated when the output of the detector is converted into an electric signal. This electrical noise is generally generated without depending on the incident X-ray dose. For this reason, if the amount of signal (number of X-rays generated or energy) is small, the weight of the electrical noise component increases, and the X-ray transmission information tends to be buried in noise and cannot be seen. In mammography, this tendency is particularly noticeable because the X-ray energy handled is particularly low.
  • the present invention has been made in view of the problems in photographing soft tissue with X-rays, which are observed in the above-described conventional X-ray mammography and the like. Specifically, i) Significantly improve the signal-to-noise ratio due to electrical noise compared to a conventional imaging device equipped with an integral X-ray detector, ii) Circuits faced by conventional devices The contrast of the high-dose part and the low-dose part is insufficient due to the narrow dynamic length of the patient, so the X-ray energy must be lowered to ensure the contrast despite the large X-ray exposure dose of the patient. The objective is to improve the situation that cannot be obtained, or to prevent the X-ray dose contributing to imaging from changing depending on the size of the breast, resulting in a change in image quality.
  • a low energy X-ray image forming apparatus basically has an energy range higher than an effective energy in an energy region of 10 to 23 keV, and a lower limit energy value of 18 keV.
  • CNR contrast-to-noise ratio
  • the conventional integration type X-ray detector is faced with a poor SN ratio due to electrical noise and a narrow dynamic length of the circuit. Since the contrast discrimination ability is insufficient in the high-dose part and the low-dose part, the problem of having to secure the contrast by lowering the X-ray energy even though the patient's X-ray exposure dose is large is improved. can do.
  • FIG. 1 is a diagram illustrating an outline of a configuration of an X-ray mammography apparatus as a low energy X-ray image forming apparatus according to a first embodiment of the present invention.
  • FIG. 2 is a graph illustrating the energy spectrum of raw X-rays exposed from the anode of the X-ray tube.
  • FIG. 3 is a graph illustrating the energy spectrum of X-rays irradiated from an X-ray tube and after passing through an aluminum filter.
  • FIG. 4 is a graph illustrating the difference in energy spectrum between the X-ray according to the present invention and the X-ray used in the conventional mammography.
  • FIG. 1 is a diagram illustrating an outline of a configuration of an X-ray mammography apparatus as a low energy X-ray image forming apparatus according to a first embodiment of the present invention.
  • FIG. 2 is a graph illustrating the energy spectrum of raw X-rays exposed from the anode of the X-ray tube.
  • FIG. 5 is another graph (with characteristic X-rays) showing the energy spectrum of the X-rays irradiated from the X-ray tube and after passing through the aluminum filter.
  • FIG. 6 is a diagram schematically illustrating a front view of the apparatus illustrated in FIG. 1.
  • FIG. 7 is a partially broken plan view showing an outline of the X-ray detector.
  • FIG. 8 is a perspective view and a cross-sectional view illustrating an outline of a detection module.
  • FIG. 9 is a block diagram showing a data acquisition circuit individually connected to a semiconductor cell constituting each pixel.
  • FIG. 10 is a diagram for explaining a relationship between an electric pulse generated in response to incidence of X-ray photons (photons) and a threshold value for discriminating the intensity.
  • FIG. 10 is a diagram for explaining a relationship between an electric pulse generated in response to incidence of X-ray photons (photons) and a threshold value for discriminating the intensity.
  • FIG. 11 is a diagram illustrating a plurality of energy ranges (BIN) and X-ray photon collection / reconstruction for each energy range.
  • FIG. 12 is a block diagram showing a configuration of an electric system including a console.
  • FIG. 13 is a diagram for explaining an outline of the configuration of an X-ray foreign matter detection apparatus as a low energy X-ray image forming apparatus according to the second embodiment of the present invention.
  • FIG. 14 is a partially broken plan view showing an outline of the X-ray detector employed in the second embodiment.
  • the low energy X-ray image forming apparatus is intended for imaging a soft tissue (soft tissue) part of a human body or the like and a substance made of soft tissue.
  • low energy of “low energy X-ray image forming apparatus” means the lower X of the X-ray energy range used in general X-ray medical diagnostic equipment, excluding conventional mammography. This name is used in the sense of using linear energy.
  • the “formation” of “image formation” goes beyond the concept of taking an X-ray image, and creates an image by applying various processes to the X-ray signal transmitted through the object received by the detector. "Formation” in the sense that it also includes.
  • CNR contrast-to-noise ratio
  • soft tissue is a term for hard tissue and is defined as a collective term for connective tissue excluding bone tissue.
  • soft tissue is defined in terms of tube voltage and CNR in a form that includes the general concept of the medical field. For this reason, the soft tissue referred to in the present invention includes not only human breasts but also non-destructive inspection objects such as food (for example, vegetables such as peppers).
  • the low-energy X-ray image forming apparatus is also called an X-ray mammography apparatus or a mammography apparatus if it is used for imaging a human breast.
  • this low energy X-ray image forming apparatus has been in the spotlight as an X-ray foreign matter detection device as a non-destructive inspection device that detects foreign matters such as hair in food.
  • an X-ray mammography apparatus will be described in the first embodiment, and an X-ray foreign object detection apparatus will be described in the second embodiment.
  • This X-ray mammography apparatus images a breast of a subject.
  • X-ray detection is performed by a method called photon counting (photon counting), and the detected value is processed by a tomosynthesis method (tomosynthesis) method.
  • photon counting photon counting
  • tomosynthesis tomosynthesis
  • the process for obtaining this image may be to obtain a transmission image called a scanogram, or to obtain a CT (Computed Tomography) image.
  • the X-ray mammography apparatus 1 includes an upright gantry 11 and an arm unit 12 that is rotatably held by the gantry 11 in the lateral direction.
  • an orthogonal coordinate system having the longitudinal direction of the gantry 11 as the Y-axis direction is set as shown in FIG.
  • the arm portion 12 has a substantially C-shaped side shape, and the beam portions 12A and 12B extending in the upper and lower two lateral directions, and one end portion of each of the beam portions 12A and 12B in the vertical direction (Y-axis direction). 12C of link parts connected to.
  • one beam portion 12A includes an X-ray generator 21 that generates X-rays.
  • the other beam portion 12B includes an X-ray detection device 31 that detects X-rays and a photon counting method.
  • the apparatus 1 includes compression plates 32A and 32B that compress the breast BR of the subject P in a plate shape so that the position in the height direction (that is, the Y-axis direction) can be adjusted.
  • the compression plates 32A and 32B are made of an X-ray transmissive material.
  • the X-ray mammography apparatus 1 includes a high voltage generator 3 that supplies a high voltage for driving to an X-ray tube, which will be described later, and a console 4 for control and image processing.
  • the high voltage generator 3 is disposed inside the beam portion 12A described above.
  • the console 4 is provided separately from the gantry 11.
  • the console 4 includes an input device 5 and a display device 6 that are used as an interface by the operator.
  • the console 4 controls the drive unit (not shown) of the gantry 11, the arm unit 12, the X-ray detection device 31, and the compression plates 32 ⁇ / b> A and 32 ⁇ / b> B, and the electrical elements in the gantry 11 and the high voltage generator 3.
  • the drive is electrically controlled. For this reason, the console 4 is communicably connected to necessary parts of the gantry 11.
  • the X-ray generator 21 includes an X-ray tube 22 and a filter 23 disposed in order on the X-ray irradiation side of the X-ray tube 22.
  • the filter 23 is a filter in which an aluminum (Al) material is formed into a plate having a desired thickness, and is hereinafter referred to as an aluminum filter.
  • the high voltage is supplied to the X-ray tube 22 from the high voltage generator 3 that generates a high voltage by inverter control.
  • This X-ray tube 22 uses tungsten (W) as its anode material 22A.
  • pulsed X-rays are exposed from the X-ray tube 22 described above.
  • the X-ray is irradiated as a pulsed X-ray beam or continuous X-ray beam collimated toward the breast BR of the subject P through the aluminum filter 23 and the collimator (or slit) 24 (see the dotted line BM1 in FIG. 1). ).
  • the collimator 24 has a contour of the X-ray beam BM1 that is substantially perpendicular to the contour of the sternum side of the subject P, and the contour of the X-ray beam BM1 on the opposite side spreads in a fan shape.
  • X-rays are collimated as follows. This is because the breast BR is imaged as closely as possible to the edge of the sternum side, and an excessive X-ray exposure of the part on the sternum side is prevented.
  • the enlargement ratio 2 times, and the phase contrast effect can be obtained.
  • the following documents can be referred to.
  • the voltage applied to the X-ray tube 22 is, for example, 30 kV. In the present invention, this voltage is set to a value between 30 and 37 kV.
  • the energy of the X-ray generated by the X-ray tube 22 itself (that is, the X-ray before passing through the filter 23) has a spectrum as schematically shown in FIG.
  • energy [keV] is taken on the horizontal axis
  • X-ray photon (photon) count is taken on the vertical axis.
  • the amount of the vertical axis of the distribution is assigned to the photon count (number of photons).
  • the tube voltage is set to 30 kV, so that 30 keV is the upper limit of energy, and there is a spectrum peak in the vicinity of 25 keV, and the energy is lower than that. Distribution extends to the band. In other words, the distribution is broad continuously from the energy on the low band side almost 0 to 30 keV and has a peak near 25 keV.
  • the intensity and energy of the generated X-rays are raised or lowered accordingly. That is, the height (equivalent to the photon count) and the width (energy value) of the energy spectrum become large (wide) according to the rise and fall of the tube voltage.
  • This energy spectrum distribution is not suitable for X-ray mammography.
  • the energy spectrum distribution of the raw X-rays exposed from the X-ray tube 22 is corrected by the aluminum filter 23. That is, the aluminum filter 23 cuts or suppresses the energy spectrum on the low frequency side, that is, the energy component of about 18 keV or less in this example.
  • the plate thickness of the aluminum filter 23 is selected so that such energy components can be cut or suppressed.
  • the X-rays exposed from the X-ray tube 22 pass through the aluminum filter 23 and have an energy spectrum as shown in FIG. According to the figure, the spectrum distribution on the low band side is cut by both filters 23, and the wide band side is suppressed by the tube voltage 30 kV.
  • the tube voltage is 37 kV
  • the energy spectrum extends to 37 keV.
  • the tube voltage can be arbitrarily set between 30 and 37 keV according to the intention of the operator. Therefore, as shown in FIG.
  • CNR contrast-to-noise ratio
  • the condition of an energy range higher than the effective energy in the energy range of 10 to 23 keV was also considered from the viewpoint of improving the problems faced by conventional mammography.
  • the central band of 18 keV to 30 (-37) keV as the X-ray band to be used may be shifted.
  • the point in creating this desired X-ray spectrum is that the energy band used for mammography in the present invention is sufficiently higher than the energy band used in conventional mammography (approximately 10 keV to 23 keV).
  • the present inventors in this embodiment have an average X-ray energy that is at least higher than the energy band in which the conventional mammography apparatus is used, and overlap with the conventional energy region ( This is why the use of an energy band of 20% or less is proposed (see the hatched portion in FIG. 4 described later).
  • FIG. 4 shows a comparison between the energy spectrum of the X-rays exposed from the X-ray generator 21 toward the breast BR of the subject P and the mainstream of conventional X-ray mammography.
  • the energy spectrum for conventional X-ray mammography is an example in which molybdenum (Mo) is used for the anode of the X-ray tube and a rhodium (Rh) filter is used as the filter. This energy spectrum is illustrated as Mo / Rh.
  • FIG. 5 shows another energy spectrum applicable in the present invention.
  • This energy spectrum uses a material other than tungsten, for example, molybdenum or copper, as the anode material 22A of the X-ray tube 22.
  • the number of photons of energy of the characteristic X-ray can be increased.
  • the image contrast due to energy in the vicinity of 26 keV is the highest, the amount of information necessary for imaging can be optimized on the X-ray generation side.
  • the compression plates 32A and 32B are configured to sandwich the breast BR of the subject P between the upper surface of the X-ray detection apparatus 31 and compress the breast BR. This is because the lesioned part can be depicted more precisely by capturing an image in a state where the breast BR is deformed to be as thin as possible.
  • FIG. 6 shows an X-ray tube 22, a collimator (slit) 24, a breast BR, and a detector 42 (described later) when the gantry 11 shown in FIG. 1 is viewed from the front direction (the direction of the arrow FR). The geometric positional relationship with the center is shown.
  • the X-ray detection device 31 includes a grid 41 for preventing scattered X-rays, an X-ray detector (hereinafter simply referred to as a detector) 42 for detecting X-rays, and a high voltage applied to the detector 42. And a bias power source 43 for supplying a bias voltage of.
  • this detector 42 includes a substrate BD and three elongated rectangular shapes that are mounted on the substrate BD at a predetermined distance from each other and parallel to each other and in which X-ray imaging elements are two-dimensionally arranged.
  • Detectors 42A to 42C Each of the three detectors 42A to 42C provides a detection surface 42F.
  • the three detectors 42A to 42C are formed as blocks independent from each other, and are mounted on the substrate BD. In this way, by arranging the three detectors 42A to 42C in a discrete manner, the X-ray imaging elements are arranged in the entire region including the space between the detectors 42A to 42C, and the detector components are compared with the detector configuration. Cost can be reduced and the incidence of scattered radiation can be suppressed.
  • Each detector 42A (to 42C) is configured as a photon counting type X-ray detector of a direct conversion method using a semiconductor.
  • each detector 42A (to 42C) is configured by cascading a plurality of detection modules M 1 to M m with a gap of a predetermined width in one direction, and on the substrate BD in the scanning direction. It is inclined by ⁇ ° (for example, 16.5 °) with respect to the orthogonal direction.
  • Each detection module M 1 ( ⁇ M m ) has collection pixels C (for example, 12 ⁇ 80 pixels) arranged in a two-dimensional manner as shown in FIG. Accordingly, the collection pixel C is also arranged obliquely by ⁇ ° with respect to the scanning direction. Therefore, even if there is a gap between the detection modules M 1 to M m , the collection pixels C are arranged over the entire desired imaging range in the direction orthogonal to the scan direction. That is, signals are reliably collected from the portion corresponding to the gap.
  • the collimator 24 is formed so that X-rays are irradiated only to the detection surfaces 42F located obliquely of the three detectors 42A to 42C.
  • Each detection module M 1 ( ⁇ M m ) includes an ASIC (Application Specific Integrated Circuit) layer A1 mounted on the substrate BD and a detection layer A2 bonded and bonded between the ASIC layer A1.
  • ASIC Application Specific Integrated Circuit
  • Each detector 42A ( ⁇ 42C) has, for example, ten detection modules M arranged in a straight line, so each detector has a collection pixel C (for example, 12 ⁇ 800 pixels).
  • the size of each collection pixel C is, for example, 200 ⁇ m ⁇ 200 ⁇ m.
  • the size of the X-ray detection surface of each detector 42A ( ⁇ 42C) is, for example, 4 mm wide ⁇ 160 mm long).
  • the detector 42 individually counts the photons (photons) corresponding to the incident X-rays by the N pixels constituting the incident surface 42F of each detector 42A ( ⁇ 42C) and calculates the count value.
  • the reflected electric quantity data is output at a high frame rate of 300 to 3,300 fps, for example. This data is also called frame data.
  • Each of the plurality of collection pixels C includes a scintillator such as a cadmium telluride semiconductor (CdTe semiconductor), a cadmium zinc telluride semiconductor (CdZnTe semiconductor), a silicon semiconductor (Si semiconductor), CsI, and a photoelectric converter such as a C-MOS.
  • a scintillator such as a cadmium telluride semiconductor (CdTe semiconductor), a cadmium zinc telluride semiconductor (CdZnTe semiconductor), a silicon semiconductor (Si semiconductor), CsI, and a photoelectric converter such as a C-MOS.
  • Semiconductor cell (sensor) Sn (n 1 to N). Each of the semiconductor cells Sn detects incident X-rays and outputs a pulse electric signal corresponding to the energy value.
  • the X-ray detection material forming each collection pixel C is a scintillator with a fast decay time using a crystal such as Pr: LuAG (praseodymium-added lutetium, aluminum, garnet) or Ce: GAGG (gadolinium aluminum gallium garnet). It may be an element combining photoelectric conversion elements such as SiPM (silicon photomultiplier).
  • the structure of the group of semiconductor cells Sn is also known from Japanese Patent Application Laid-Open Nos. 2000-69369, 2004-325183, and 2006-101926.
  • the size (200 ⁇ m ⁇ 200 ⁇ m) of each collection pixel C described above is a sufficiently small value that can detect the number of X-rays as particles (X-ray photons).
  • the size capable of detecting X-rays as the particles means “an electric pulse signal in response to each incident when a plurality of radiation (for example, X-ray) particles are successively incident at or near the same position.
  • the occurrence of a superposition phenomenon (also called pile-up) is defined as “a size that can be substantially ignored or whose amount is predictable”.
  • each detector 42A (to 42C) is set to such a size that it can be assumed that this counting-out does not occur or substantially does not occur, or the counting-off amount can be estimated.
  • the feature of each detector 42A ( ⁇ 42C) is that the number of X-ray pulses can be accurately measured while accurately performing energy discrimination.
  • a waveform shaping circuit, a multistage circuit is provided at the subsequent stage of the charge amplifier. Comparator, multi-stage counter, multi-stage D / A converter, latch circuit, and serial converter. The circuit configuration of these is known from Japanese Patent Application Laid-Open No. 2006-101926.
  • one pulse signal can be individually compared with different analog amount threshold values th 1 to th 3 .
  • the reason for this comparison is that the energy value of the incident X-ray particle is in any of the energy regions ER EX , ER 1 to ER 3 (also referred to as BIN: see FIG. 11) set in advance in three.
  • the lowest analog amount threshold th 1 is usually used to prevent detection of disturbance, noise caused by circuits such as the semiconductor cell Sn and the charge amplifier, or low-energy radiation that is not necessary for imaging.
  • the band ER EX whose energy is lower than the lowest analog amount threshold th 1 is treated as “a non-measurable (non-measurement) region” because there is much information that depends on noise and disturbance.
  • the number of photons in the highest energy region ER 3 is counted, but is treated as a value not used for image reconstruction.
  • the counters 56 1 to 56 3 arranged in each data collection circuit 51 n enter the first (to third) energy region ER 1 (to ER 3 ) that they should be responsible for counting. Count the number of photons with energy and energy exceeding it. Therefore, if the number of X-ray photons having energy belonging to the first to third energy regions ER 1 to ER 3 , that is, the number of X-ray photons to be obtained for each energy region is W 1 , W 2 , and W 3.
  • the meaning of “acquisition” for each energy region of the number of X-ray photons according to the present application is the meaning of “determining by calculation” from the actual count value as described above, and the energy region as in a modification described later. Both meanings of directly “counting” the number of X-ray photons for each are included.
  • the counter 56 1-56 4 described above is given a signal to start and stop the controller to be described later of the console 4. Counting for a fixed time is managed from the outside using a reset circuit included in the counter itself.
  • the number of thresholds that is, the number of comparators is not necessarily limited to three, and may be two including the analog amount threshold th 1 or any number of three or more. Also good.
  • analog amount threshold values th 1 to th 3 described above are given digital values from the console 4 as calibrated values for each collection pixel C, that is, for each collection channel.
  • the number of X-ray particles incident on each detector 42A ( ⁇ 42C) is obtained by the three counters 56 1 to 563 during the fixed collection time reset at a fixed cycle, so that the collection pixel C It is measured every time and every energy region.
  • Count of the X-ray particle number count data W 1 of the digital value from each of the first to third counters 56 1 ⁇ 56 3 ', W 2', after being outputted in parallel as W 3 ', illustrated Not converted to serial format by serial converter.
  • This serial converter is serially connected to the serial converters of all remaining acquisition channels. For this reason, the count data of all digital quantities are serially output from the serial converter of the last channel and sent to the console 4.
  • the console 4 includes an interface (I / F) 61 that performs input and output of signals, and a controller (CPU) 63, a RAM ( A storage unit) 64, an image processor 65, and a ROM 70.
  • the interface 61 is connected to the input device 5 and the display device 6 and can communicate with the controller 63.
  • the controller 63 controls the driving of the gantry 11 in accordance with a program given in advance to the ROM 70. This control includes a command value sending command to the high voltage generator 3.
  • the RAM 64 temporarily stores the frame data sent from the gantry 11 via the interface 61.
  • the image processor 65 executes various processes under the management of the controller 63 based on a program given in advance to the ROM 70.
  • This process includes a process for executing a known CT reconstruction method or a process for executing a tomosynthesis method called “shift and add”.
  • a tomographic image of a desired cross section of the breast BR of the subject P is obtained using frame data based on the count value of the number of X-ray photons collected for each energy region output from each detector 42A ( ⁇ 42C). Created.
  • the display 6 displays the image created by the image processor 65.
  • the display device 6 is also responsible for displaying information indicating the operating status of the gantry 11 and operator operation information provided via the input device 5.
  • the input device 5 is used by an operator to give information necessary for imaging to the system.
  • the controller 63 and the image processor 65 include a CPU (central processing unit) that operates according to a given program. Those programs are stored in the ROM 70 in advance.
  • the arm portion 12 of the gantry 11 is rotated or rotated around the breast BR of the subject P under the control of the controller 63. During this rotation, X-rays are emitted from the X-ray generator 21 toward the breast BR to be imaged.
  • the energy spectrum of this X-ray is corrected by the aluminum filter 23 as described above. That is, the spectrum is corrected as shown in FIG. According to this corrected spectrum, it has a broad energy in a band of about 18-30 (or -35) keV. That is, energy is almost cut by the aluminum filter 23 in a band lower than about 18 keV. X-rays having main energy in the band of about 18-30 (or -37) keV pass through the breast BR, which is soft tissue.
  • the above-described frame data of the data directly converted from the X-rays into the digital electricity is output from the detectors 42A to 42C.
  • This frame data is data reflecting the integrated value of the number of X-ray photons for each energy band ER in each collection pixel C.
  • This frame data is collected for each frame at a constant frame rate while the arm unit 12 rotates around the center of rotation (see FIG. 6) or rotates or moves within a certain range.
  • the frame data is sequentially sent to the console 4 and stored in the RAM 64.
  • the image processor 65 reads out the frame data stored in the RAM 64 in accordance with an operator command from the input device 5, and uses this frame data to produce an image such as a breast.
  • An X-ray transmission image of a cross section with BR is reconstructed, for example, under the tomosynthesis method. From each collected pixel C, frame data of two energy regions ER 1 and ER 2 are obtained.
  • the image processor 65 for example, subjected to a weighting of low or zero frame data higher energy band ER 2, subjected to a high weighted by frame data of lower energy bands ER 2, they Each collection pixel C is mutually added. Thereby, collected data is created for each collection pixel C. As a result, data associated with the X-ray scan collected from all the collected pixels C is prepared, so that these collected data are processed by an appropriate reconstruction method to reconstruct an image of the breast BR (FIG. 11, step S1). This panoramic image is displayed on the display 36, for example (FIG. 11, step S12). Of course, the image may be reconstructed without weighting.
  • Photon counting detection technology is a technology in which X-rays are regarded as particles, the energy of the particles can be regarded as the height of the pulse, and the pulse signal is shaped and an energy threshold value is provided, so that only the pulses exceeding the threshold value are measured with a counter.
  • a system capable of independently measuring with pixels of about 200 ⁇ m ⁇ 200 ⁇ m and distinguishing a plurality of energy thresholds has been commercialized. This technique is characterized in that electric noise is not applied in order to set a threshold value for energy that is at least higher than electric noise.
  • the linear absorption coefficient of the material differs depending on the X-ray energy. Therefore, in a material having a high linear absorption coefficient, it is easy to obtain contrast at a high energy and the linear absorption coefficient is low. Substances tend to provide contrast with low energy. From this, it is possible to display both the mass and calcification optimally, or to perform a processing method that obtains the maximum contrast in either one by appropriately weighting and adding the images for each BIN. Become.
  • the SN ratio is poor due to electrical noise, which is confronted by the conventional integral X-ray detector, and the dynamic length of the circuit is low. Due to the narrowness, the contrast discrimination ability is insufficient in the high-dose part and the low-dose part, so the X-ray energy has to be lowered to ensure contrast even though the patient's X-ray exposure dose is large. The point can be improved.
  • the problem that the pixel size of 100 um or less is relatively difficult due to the large circuit mounting amount of the photon counting type detector is realized by using a small focus X-ray tube to realize an enlargement effect and a phase contrast effect. Resolve the resolution issue required for visualization. It is possible to optimize the imaging of a tumor with a relatively excellent linear absorption coefficient and a high calcification and a low linear absorption coefficient, which is a characteristic matter in X-ray mammography.
  • the X-ray detection method employs a photon counting type detector that can output at least two energy bands, and the resolution is less than twice that of the subject or object to be obtained. Resolution.
  • the X-ray generator has, for example, a filter disposed in an X-ray tube having an anode, which is a filter that suppresses the passage of X-ray particles having energy on the higher frequency side than the energy spectrum.
  • the tube focus size is 0.056 mm or less
  • the subject or test object is separated from the X-ray tube focus position by 0.5 m or more
  • the distance from the subject or test object to the detector is 0.5 m or more
  • the resolution is secured by the enlargement effect.
  • phase contrast effect for example, “Konica Minolta's phase contrast technology: 1406264584500_0.html” can be referred to.
  • the energy band in which the contrast is optimally obtained differs depending on the tumor or calcification. For this reason, an optimum image can be obtained by weighting and adding the obtained image for each energy band.
  • the CNR Contrast-to-Noise Ratio
  • the CNR can be optimized in order to optimize the depiction of the tumor in the energy band in the energy range from 18 keV to 30 keV to 35 keV. From this viewpoint, a technique of optimizing characteristic X-rays is also possible.
  • this X-ray foreign object detection device 80 is a food FD (substance corresponding to soft tissue of the human body from the viewpoint of contrast-to-noise ratio (CNR)) that is carried on a belt conveyor 81A, 81B, 82C. ) Is a device that detects hair HR as foreign matter that may be present in or around the X-ray. For this reason, this X-ray foreign material detection device 80 is provided on the belt conveyor 81B in the middle and periodically forms an X-ray image at regular intervals in a non-contact manner without stopping the food FD being conveyed. The hair HR is detected and appropriate processing such as notification is performed.
  • CNR contrast-to-noise ratio
  • the foreign object detection device 80 has a box-shaped casing 90, the X-ray generator 21 described above is provided inside the casing 90, and the collimator 24 is provided on the emission side of the X-ray generator 21. ing.
  • a flexible X-ray shield 90 ⁇ / b> A is provided at the food inlet and the food outlet of the casing 90.
  • an X-ray detector 83 that receives transmitted X-rays is provided below the belt conveyor 81B.
  • L1 L2.
  • the detector 83 may be positioned in the space 81S between the belts BL that move in opposite directions on the upper side and the lower side in the height direction (Y-axis direction) of the belt conveyor 81B.
  • the belt BL is made of a material having X-ray transparency.
  • the detector 83 is configured by arranging the 29 detection modules M described above in tandem in one direction.
  • the detector 83 is disposed on the substrate BD so as to be inclined by ⁇ ° (for example, 16.5 °) with respect to the scanning direction, that is, the direction in which the food FD is conveyed.
  • the detector 83 according to the second embodiment has a single number and a large number of modules arranged in tandem, that is, , Except that it is longer than that of the first embodiment.
  • the console 4 applies the frame data detected by the detector 83 at a high frame rate to a shift & add process in accordance with the moving speed of the belt conveyor 81B, for example.
  • a tomographic image along a virtual surface assumed at the same height as the detection surface 83F of the detector 83 or a virtual surface assumed at a desired height is formed at a constant period.
  • the food FD is reflected, and if there is a foreign object such as hair HR, it is reflected together in a state where it is superimposed on the food FD.
  • the console 4 recognizes this foreign substance by a known image recognition method, and performs a process of notifying the operator and giving an instruction to remove the corresponding food FD from the line.
  • the X-ray foreign matter detection device 80 in addition to the same operational effects as described above, the presence of foreign matter that is difficult to be imaged by conventional X-ray photography, such as hair and fine and fine foreign matter, is high-resolution. It can be detected through image formation. Further, the apparatus can be miniaturized by shortening the time for detecting the foreign matter or reducing the tube current. Furthermore, the manufacturing cost of the apparatus can be reduced.

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