[go: up one dir, main page]

WO2004051312A1 - Detecteur de rayonnement et dispositif de diagnostic par imagerie medicale - Google Patents

Detecteur de rayonnement et dispositif de diagnostic par imagerie medicale Download PDF

Info

Publication number
WO2004051312A1
WO2004051312A1 PCT/JP2003/015410 JP0315410W WO2004051312A1 WO 2004051312 A1 WO2004051312 A1 WO 2004051312A1 JP 0315410 W JP0315410 W JP 0315410W WO 2004051312 A1 WO2004051312 A1 WO 2004051312A1
Authority
WO
WIPO (PCT)
Prior art keywords
substrate
radiation detector
phosphor
radiation
angle
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/JP2003/015410
Other languages
English (en)
Japanese (ja)
Inventor
Makoto Sato
Tsuneyuki Kanai
Ichiro Miura
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.)
Hitachi Healthcare Manufacturing Ltd
Original Assignee
Hitachi Medical Corp
Priority date (The priority date 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 date listed.)
Filing date
Publication date
Application filed by Hitachi Medical Corp filed Critical Hitachi Medical Corp
Publication of WO2004051312A1 publication Critical patent/WO2004051312A1/fr
Anticipated expiration legal-status Critical
Ceased legal-status Critical Current

Links

Classifications

    • 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
    • G01T1/20183—Arrangements for preventing or correcting crosstalk, e.g. optical or electrical arrangements for correcting crosstalk
    • H—ELECTRICITY
    • H10—SEMICONDUCTOR DEVICES; ELECTRIC SOLID-STATE DEVICES NOT OTHERWISE PROVIDED FOR
    • H10F—INORGANIC SEMICONDUCTOR DEVICES SENSITIVE TO INFRARED RADIATION, LIGHT, ELECTROMAGNETIC RADIATION OF SHORTER WAVELENGTH OR CORPUSCULAR RADIATION
    • H10F30/00—Individual radiation-sensitive semiconductor devices in which radiation controls the flow of current through the devices, e.g. photodetectors
    • H10F30/20—Individual radiation-sensitive semiconductor devices in which radiation controls the flow of current through the devices, e.g. photodetectors the devices having potential barriers, e.g. phototransistors
    • H10F30/29—Individual radiation-sensitive semiconductor devices in which radiation controls the flow of current through the devices, e.g. photodetectors the devices having potential barriers, e.g. phototransistors the devices being sensitive to radiation having very short wavelengths, e.g. X-rays, gamma-rays or corpuscular radiation
    • 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]
    • H—ELECTRICITY
    • H04—ELECTRIC COMMUNICATION TECHNIQUE
    • H04N—PICTORIAL COMMUNICATION, e.g. TELEVISION
    • H04N23/00—Cameras or camera modules comprising electronic image sensors; Control thereof
    • H04N23/30—Cameras or camera modules comprising electronic image sensors; Control thereof for generating image signals from X-rays

Definitions

  • the present invention relates to a radiation detector that detects X-rays, ⁇ -rays, and the like, and particularly relates to a radiation detector suitable for a medical image diagnostic device such as an X-ray CT device or PET.
  • the present invention also relates to a medical image diagnostic device using the radiation detector. Background technology
  • Solid-state detectors which combine a phosphor element and a photodiode, are widely used as radiation detectors for medical image diagnostic devices such as X-ray CT and PET.
  • the structure of a solid state detector is generally a combination of a scintillator that emits visible light by absorbing X-rays and a photodiode that converts the emitted light into an electric signal.
  • Element arrays arranged in one or two dimensions are used. In such a detection element array, the scintillator needs to be channel-separated in order to prevent crosstalk between adjacent channels, and a reflection layer is usually formed between adjacent scintillator elements.
  • the scintillator elements and the reflection layer have a rectangular parallelepiped shape for a channel-separated scintillator element, and the reflection layer provided between the elements has a rectangular parallelepiped shape.
  • a scintillator element and a reflection layer alternately arranged are adhered on a plane of a photodiode array substrate to form a detection element array block.
  • Fig. 12 shows a conventional detector array. Focusing on one detector array block 90, the scintillator element 91 at the center of the block faces the X-ray tube focal point, while the scintillator element 92 at the end of the block faces the X-ray tube focal point. It will be. For this reason, the scintillator element 92 and the scattered ray removing collimator 93 are shifted from each other toward the end of the element array block, and the irradiation area seen from the focal point of the X-ray tube becomes wider. As a result, the amount of light emitted during X-ray irradiation increases, and the output changes. I do.
  • the path of the X-ray passing through the element differs between the scintillator element 91 at the center of the block and the scintillator element 92 at the end, the X-ray energy characteristics also change. For these reasons, variations occur in the light emission characteristics within the element array block, and artifacts tend to occur in the CT image.
  • Japanese Patent Application Laid-Open No. 2000-237178 discloses that a plurality of detector element array blocks each composed of a plurality of elements are arranged in a chord shape by observing the circle around the focal point of the X-ray tube from the circumferential axis direction. This constitutes a channel detection element array.
  • a solid-state detector is provided with a scattered radiation removing collimator in order to prevent scattered X-rays scattered by the subject from being incident on the detector element.
  • a plurality of scattered radiation removing collimators are arranged in an arc shape so as to cover the gap between the detector elements and are accurately aligned with the detector element array.
  • angles of the plurality of reflective layers are particularly determined such that the plurality of phosphors fractionated by the plurality of reflective layers on the substrate are directed to the radiation source (focal point). .
  • the radiation detector according to the present invention includes a substrate, a plurality of phosphor elements arranged on the substrate, and a reflection layer disposed between the adjacent phosphor elements to prevent crosstalk between the adjacent phosphor elements.
  • a radiation detector comprising: a plurality of photoelectric conversion elements provided at positions corresponding to the phosphor elements on the substrate; and a contact surface between the phosphor element and the reflective layer; and a substrate at a central portion of the substrate. The angle between the substrate and the surface changes from the center of the substrate toward the end along the arrangement direction of the phosphor elements.
  • the orientation of the phosphor element defined by the phosphor element and the reflection layers on both sides of the phosphor element is substantially changed with respect to all the phosphor elements arranged in a plane. Since the incident direction can be the same, the apparent irradiation area of the radially incident radiation and the passage in the element can be made uniform, and the Variations in the output of the element can be eliminated.
  • the angle formed between the contact surface between the phosphor element and the reflective layer and the substrate surface at the central portion of the substrate is, for example, with respect to the contact surface between the phosphor element and the reflective layer located at the central portion of the substrate or the position closest to the central portion.
  • the distance from the radiation source to the top surface of the phosphor element is calculated from 90 ° for each phosphor array pitch, where Ld is the distance from the radiation source to the top of the phosphor element. Lp should be varied by the arrangement pitch of the phosphor elements.) Considering the processing accuracy of the radiation detector, an error of ⁇ 0.2 ° can be tolerated.
  • the angle between the contact surface between the base phosphor element and the reflection layer located at the end of the substrate and the substrate surface at the center of the substrate (the angle at the center of the substrate).
  • the width changes along the depth direction by changing the direction of the phosphor elements in the arrangement direction.
  • This change in width can be provided in the phosphor element and / or the reflective layer.
  • the length in the direction in which the phosphor elements and the reflection layer are arranged is referred to as the width of the phosphor elements and the reflection layer, and the distance from the radiation incident surface to the substrate surface is referred to as the height or the depth.
  • a medical image diagnostic apparatus includes a radiation source, a radiation detector having the above-described features disposed to face the radiation source, and holding the radiation source and the radiation detector, and driving the radiation source around a subject. Rotating disk, and radiation detected by the radiation detector. Image reconstruction means for reconstructing an image of the tomographic image of the subject based on the intensity of the line.
  • a substrate having a plurality of photoelectric conversion elements arranged in at least one predetermined direction is disposed on the substrate so as to correspond one-to-one to the plurality of photoelectric conversion elements.
  • a radiation detector in which an angle between a contact surface between a phosphor element and the reflective layer and the substrate surface changes from the center to the end of the substrate along the arrangement direction.
  • the angle between the contact surface between the phosphor element and the reflective layer and the substrate surface becomes smaller from the center to the end of the substrate.
  • the width of the phosphor element in the predetermined direction is constant irrespective of the position in the depth direction of the phosphor element orthogonal to the predetermined direction.
  • the width of the reflective layer in the predetermined direction is constant irrespective of the position in the depth direction of the reflective layer orthogonal to the predetermined direction.
  • the plurality of phosphor elements are formed of a single member, and the upper surface and the lower surface thereof are respectively aligned and parallel, and the reflection layer is formed of the plurality of phosphor elements. It is formed in a groove for partitioning the phosphor element.
  • the width of the groove for partitioning the plurality of phosphor elements is substantially constant irrespective of the position in the thickness direction, and the reflecting material is cured after filling with a curable material. Then, a metal plate or a metal plate was adhered in the groove.
  • the width of the groove for partitioning the plurality of phosphor elements becomes thicker as the position in the thickness direction becomes deeper, and the reflecting material is cured after filling and shrinking a material that cures and contracts. did.
  • the material to be cured is epoxy resin, acrylic resin It is at least one of inorganic resin powders such as titanium oxide, aluminum oxide, and barium sulfate contained in an organic resin such as lunar phenol resin.
  • the material that cures and shrinks is an organic resin such as an epoxy resin, an acryl resin, or a phenol resin, and an inorganic compound powder such as titanium oxide, silicon oxide, or barium sulfate. It is at least one of those contained.
  • a plurality of the above-described radiation detectors are prepared and juxtaposed in a one-dimensional or 2′-dimensional direction such that their long-axis directions intersect at a predetermined angle of 0 ° to 90 °.
  • a radiation detector unit configured as follows.
  • a radiation detector arranged to face a radiation source, a rotating disk that is driven to rotate while holding the radiation detector, and detected by the radiation detector 3.
  • a medical image diagnostic apparatus comprising: image reconstruction means for reconstructing a tomographic image of an object based on the intensity of radiation, wherein the medical image diagnostic apparatus uses the radiation detector according to claim 2 as the radiation detector.
  • a radiation detector arranged to face a radiation source, a rotating disk that is driven to rotate while holding the radiation detector, and detected by the radiation detector
  • a medical image diagnostic apparatus comprising: image reconstruction means for reconstructing a tomographic image of an object based on the intensity of radiation, wherein the medical image diagnosis uses the radiation detector unit according to claim 10 as the radiation detector.
  • the shortest distance from the radiation source to the upper surface of the phosphor element is Ld
  • the arrangement pitch of the phosphor elements in the predetermined direction is Ld.
  • the shortest distance from the radiation source to the upper surface of the phosphor element is Ld,
  • FIG. 1 is a cross-sectional view of a radiation detector according to one embodiment of the present invention.
  • 2A and 2B are cross-sectional views of a main part of the radiation detection element array block in FIG. 1, wherein FIG. 2A shows a case where the number of detection elements is even, and FIG. 2B shows a case where the number of detection elements is odd.
  • Figure 3 is a diagram showing the positional relationship between the radiation source and the radiation detection element array.
  • FIG. 4 is a sectional view of a main part of a radiation detector according to another embodiment of the present invention.
  • FIG. 5 is an example in which a plurality of detectors are arranged in the rotation axis direction according to the embodiment.
  • FIG. 1 is a cross-sectional view of a radiation detector according to one embodiment of the present invention.
  • 2A and 2B are cross-sectional views of a main part of the radiation detection element array block in FIG. 1, wherein FIG. 2A shows a case where the number of detection elements is even, and FIG
  • FIG. 6 is a diagram illustrating an example of a method for manufacturing a radiation detector according to the present invention.
  • FIG. 1 is a configuration diagram of an X-ray CT apparatus equipped with a radiation detector according to the present invention.
  • FIG. 8 (a) is a diagram showing the X-ray energy characteristics of the radiation detector of the present invention
  • FIG. 8 (b) is a diagram showing the X-ray energy characteristics of the conventional radiation detector.
  • FIG. 9 shows a flow of manufacturing the detector according to the first embodiment of the present invention.
  • FIG. 10 shows a flow of manufacturing a detector according to the second embodiment of the present invention.
  • FIG. 11 shows a flow of manufacturing the detector according to the third embodiment.
  • FIG. 12 is a diagram showing the relationship between a conventional radiation detecting element array and a scattered radiation removing collimator.
  • FIG. 1 is a cross-sectional view of a radiation detector according to one embodiment of the present invention.
  • the radiation detector includes a substrate 11, a photodiode 14 (photoelectric conversion element) formed on the substrate, a plurality of scintillator elements 12 (phosphor elements), and a reflection layer 13.
  • the scintillator elements 12 are arranged directly above the photodiodes 14 in a one-to-one relationship, and the scintillator elements 12 and the reflective layers 13 are alternately arranged in the width direction. Therefore, immediately below the reflective layer 13 is a portion of the substrate 11 where the photodiode 14 does not substantially exist.
  • One scintillator element is constituted by the scintillator element 12 which is channel-separated by sandwiching both sides in the width direction with the reflective layer 13 and the photodiode 14 immediately below the scintillator element.
  • the size of the scintillator element and the reflective layer and the number of scintillator elements arranged depend on the application of the radiation detector and are not particularly limited.For example, in the case of a detector array for an X-ray CT device, a substrate having a width of about 20 to 40 mm is used.
  • the scintillator elements 12 are provided at an arrangement pitch Lp of about 1 mm.
  • the width of the reflective layer 13 in the arrangement direction is about 1/10 to 210 of the normal arrangement pitch Lp.
  • the radiation detector according to the present embodiment is configured so that the angle formed between the contact surface between the scintillator element 12 and the reflective layer 13 and the substrate surface gradually changes while keeping the arrangement pitch of the scintillator elements constant.
  • the change in the angle is such that the direction of the normal to the contact surface between the scintillator element 12 and the reflective layer 13 is almost perpendicular to the line connecting the scintillator element 12 and the radiation source (or the focal point, if it is an X-ray).
  • the arrangement of the scintillator element 12 and the reflective layer 13 of the radiation detector is bilaterally symmetric about the reflective layer 13 in FIG. 2A, and bilaterally symmetric about the scintillator element 120 in FIG. 2B.
  • the center of the reflective layer 13 is located at the center of the substrate 11, as shown in FIG.
  • the angle ⁇ no between the contact surface between the reflective layer 13 and the scintillator elements 121 on both sides and the substrate plane is 90 °.
  • Scintillator elements 121, 121 of the angle theta .kappa.1 the contact surface of the reflective layer and the scintillator elements in the end portion side on the outside makes with the substrate surface (angle of the substrate center side) as will be hereinafter both 90 ° I do. Thereafter, the angle between the contact surface and the substrate surface gradually decreases toward the end. I do.
  • the center of the scintillator element 120 is located at the center of the substrate 11, as shown in FIG.
  • the angle 0so between the contact surface of the scintillator element 120 and the reflective layers 131 on both sides with the substrate plane is 90 °.
  • the angle 0 S1 (the angle at the center of the substrate) formed by the contact surface between the scintillator elements 121 and 121 and the reflection layers 131 at the end further to be 90 ° or less. You. Thereafter, the angle between the contact surface and the substrate surface is gradually reduced toward the end.
  • the angle is changed such that the direction of the scintillator element matches the direction of the radiation source.
  • the number of phosphor elements in the arrangement direction is an even number
  • the contact surface between the reflective layer 13 and the scintillator element 12 located at the edge of the substrate and the substrate at the center of the substrate are described.
  • Ld is the distance from the radiation source to the top of the scintillator element
  • Lp is the The arrangement pitch
  • a in the above equations 1 and 2 is a coefficient.
  • ⁇ Otan-1 (Lp / Ld)) is equivalent to the elevation angle per pitch of the scintillator element, and thus the shape of the scintillator element for each elevation pitch By inclining the shape, the shape of the scintillator element always faces the radiation source. This can prevent radiation entering the adjacent scintillator element from entering, thereby suppressing variations in the amount of emitted light.
  • Ld 1037 mm
  • the detection element array block 32 is the conventional detection element array block 80
  • the reflective layer at the center of the substrate is configured to be parallel to the normal direction of the substrate.
  • the angle between the normal to the side of the reflective layer at the center of the substrate (the surface that intersects the substrate surface) and the direction of the focal point of the X-ray tube is 90 degrees.
  • the shape of the scintillator element is not Will be. The same applies to the case where an odd number of scintillator elements are arranged.
  • FIGS. 1 and 2 show a case where the width of the reflective layer is constant at any position in the depth direction, but as shown in FIG. 4, the width of the reflective layer is changed in the depth direction.
  • the width of the scintillator element may be constant at any position in the depth direction by changing the position.
  • only one-dimensional array is shown in the figure, a two-dimensional element array can have a similar structure in each two-dimensional array direction.
  • a two-dimensional element array As shown in FIG. 5, a plurality of the above-described two-dimensional element arrays are rotated so that their major axes intersect at a predetermined angle of 0 ° or more and 90 ° or less.
  • the two-dimensional element arrays can be arranged at different angles so that each two-dimensional element array faces the direction of the X-ray tube focal point. In this case, the variation in the detector array direction can be suppressed over a wider area, and the artifact caused by this can be reduced and eliminated.
  • any of the following conventional methods (1) to (3) for manufacturing a channel-separated scintillator element block can be applied.
  • the method (3) is desirable in terms of ease of production and performance of the obtained element.
  • a plurality of plate-shaped scintillator materials are laminated at predetermined Then, the laminated block is cut to a predetermined thickness and bonded to the substrate on which the photodiode array is formed.
  • Cut grooves are formed at a predetermined pitch in the plate-shaped scintillator material, and the material constituting the reflective layer is filled in this groove by a specific method. A method of bonding on a substrate on which is formed.
  • Figure 6 illustrates the method (3).
  • a notch groove 51 having a predetermined pitch, a predetermined width, and a predetermined depth is formed in a rectangular parallelepiped scintillator raw material 50 without leaving a bottom thereof.
  • a phosphor having a high radiation-to-light conversion efficiency such as a rare-earth phosphor can be used.
  • the phosphor is, for example, a rare-earth acid composed of a host crystal having a garnet structure containing at least Gd, Al, Ga, and O with Ce as a light-emitting element as described in Japanese Patent Application Publication No. JP-A-2001-004753. It is a scintillator of arsenal.
  • the cut groove 51 is filled with a material 52 for forming a reflective layer.
  • each cut groove 51 is formed to have a predetermined angle with respect to the surface of the scintillator material. This angle is such that the angle formed by the contact surface between the reflective layer and the scintillator element and the substrate surface, when formed as a detection element array as described above, is inclined from the center to the end of the substrate.
  • the processing can be achieved by sequentially changing the angle between the material and the cutting edge in a processing machine such as a multi-wire saw-slicer every time it approaches the end.
  • the material 52 is a material that cures and shrinks
  • the cut grooves 51 may be formed in parallel.
  • the reflective layer tries to shrink due to curing shrinkage, but the dimensions of the bottom surface do not change because the bottom surface of the scintillator material is connected. Therefore, the reflective layer near the bottom of the scintillator does not shrink much, but shrinks greatly toward the top. As a result, a substantially wedge-shaped cross-sectional shape whose width changes in the depth direction as described above can be provided to the reflective layer. After the cut grooves 51 are formed in this way, a material 52 for forming a reflective layer is filled in the grooves.
  • epoxy resin acrylic resin
  • acrylic resin An organic resin such as an enol resin, a powder containing an inorganic compound powder such as titanium oxide, aluminum oxide, barium sulfate or the like as a reflective material, a metal plate having a high light reflectance, or a combination thereof.
  • the material 52 when a material containing a reflecting material in a resin is used as the material 52, it is centrifugally injected using a centrifuge. As a result, the material 52 can be injected to the bottom of the groove, and air bubbles in the resin can be substantially eliminated.
  • the reflecting layer is formed by hardening the organic resin in the material 52.
  • the adhesive is cured as necessary. Thereafter, the slice is sliced to a desired thickness in a direction orthogonal to the depth direction of the groove.
  • the resin shrinks when the organic resin hardens, but does not shrink because the bottom of the groove is fixed to the bottom of the scintillator material. Therefore, a slope is formed from the bottom to the top of the groove.
  • the degree of cure shrinkage varies depending on the type and viscosity of the resin, and by appropriately selecting these conditions, a desired inclination can be imparted to the groove. For example, when an epoxy resin having a viscosity of 10000 20000 cPs is used as an organic resin, a groove having a width of 0.15 mm and a depth of 10 mm can form a slope of about 0.0400 °.
  • the width in the depth direction as described above is substantially changed.
  • a wedge-shaped cross-sectional shape can be imparted to the reflective layer.
  • a reflection layer having a desired angle can be formed by previously inclining the cut groove 51 itself.
  • the radiation detector of the present invention is obtained by bonding the scintillator manufactured in this manner on the substrate 53 on which the photodiode array is formed in advance in the same arrangement pitch as the scintillator elements.
  • a known photodiode such as a PIN photodiode can be used as the photodiode.
  • the outline of an X-ray CT apparatus is shown as an example of medical image diagnosis to which the present invention is applied.
  • the X-ray CT apparatus includes a scan gantry unit 610 and an image reconstruction unit 620.
  • the scan gantry unit 610 includes a rotating disk 611 having an opening 614 into which a subject is loaded, and the rotating disk 611.
  • X-ray tube 612 mounted on the X-ray tube
  • collimator 613 mounted on the X-ray tube 612 to control the radiation direction of the X-ray flux
  • an X-ray detector mounted on a rotating disk 611 facing the X-ray tube 612 615
  • a detector circuit 616 for converting the X-rays detected by the X-ray detector 615 into a predetermined signal
  • a scan control circuit 617 for controlling the rotation of the rotating disk 611 and the width of the X-ray flux.
  • the image reconstruction unit 620 performs image processing for reconstructing a CT image by processing the measurement data S1 transmitted from the input device 621 and the detector circuit 616 for inputting the subject's name, examination date and time, examination conditions, and the like.
  • Circuit 622 an image information adding unit 623 for adding information such as the subject name, examination date and time, and examination conditions input from the input device 621 to the CT image created by the image arithmetic circuit 622, and image information added.
  • a display circuit 624 that adjusts the display gain of the CT image signal S2 and outputs the result to the display monitor 630.
  • X-rays are emitted from an X-ray tube 612 in a state where a subject is placed on a bed (not shown) installed in an opening 614 of a scan gantry section 610.
  • This X-ray is given directivity by a collimator 613 and detected by an X-ray detector 615.
  • X-rays transmitted through the subject are detected while changing the direction of X-ray irradiation.
  • the tomographic image created by the image reconstruction unit 620 based on the measurement data is displayed on the display monitor 630.
  • the X-ray detector 615 is composed of a large number of detection element array blocks, each of which has a plurality of detection elements that combine a scintillator element (phosphor element) and a photodiode (photoelectric conversion element), arranged in a polygonal shape close to an arc. Consists of It is assumed that the radiation detector has a detection element array block according to the present embodiment as shown in FIG. 1, FIG. 2, or FIG.
  • the radiation detector may be a one-dimensional detector in which elements are arranged in a row in the circumferential direction, or a two-dimensional detector in which elements are arranged in a row. In the case of a two-dimensional detector, the array direction parallel to the rotation axis of the rotating disk 611 (the body axis direction of the subject) Also, each element is arranged facing the direction of the X-ray tube focal point.
  • the X-ray detector 615 is arranged on an arc centered on the focal point of the X-ray tube 612 so that the side where the scintillator element is arranged faces the radiation source.
  • the X-ray detector 615 is configured to have a shape directed in the X-ray irradiation direction from the focal point because the scintillator element is divided by the scintillator element and the reflective layers on both sides thereof. Even in the case of a scintillator element, sufficient radiation is incident as in the case of the central scintillator element, and a uniform output can be obtained.
  • Fig. 8 (a) shows the X-ray energy characteristics when the X-ray detector according to the embodiment of the present invention as shown in Fig. 1 is used
  • Fig. 8 (b) shows the results using the conventional X-ray detector.
  • the X-ray energy characteristics in each case are schematically shown.
  • the X-ray energy characteristics in the element array block vary widely. This causes variations in the light emission characteristics, and artifacts are likely to occur in CT images.
  • the dispersion of the X-ray energy characteristics inside the block can be suppressed, and the dispersion of the emission characteristics can be reduced. For this reason, a high-quality CT image with few artifacts can be obtained.
  • Gds (Al, Ga) 5 0 12 Ce ceramic scintillator from configured scintillator plate X- Y plane of the (X, Y and Z directions of dimensions each 26 X 30 X 1.8mm), 24 at lmm pitch
  • the channel was attached to a substrate on which a one-dimensional photodiode array was formed. Then, using a 0.13 mm thick diamond grindstone, grooves parallel to the Y direction were machined at lmm pitch in the X direction of the scintillator plate. The depth of the groove at this time is 1.8 mm.
  • the setting was such that the grindstone entered at an angle of 0.66 ° with respect to the Z direction (the normal direction of the photodiode array substrate).
  • the grindstone was set to enter the groove at the angle of 0.605 ° in the second and subsequent grooves from the end.
  • the angle of the grindstone was gradually reduced in steps of 0.055 ° toward the center. In this way, the center groove (13th groove) of the scintillator plate
  • the angle of the grindstone becomes 0 °
  • the groove on the other end side gradually increases the angle of the grindstone in the opposite direction, and the grindstone becomes 0.66 ° in the groove closest to the other end (25th).
  • the angle was set for machining.
  • the reflective layer has a shape in which both side surfaces are parallel as shown in FIG.
  • the X-ray energy characteristics E were defined as follows by irradiating X-rays at a tube voltage of 100 kV and 120 kV and a tube current of 100 mA, and by the output ratio of each channel at that time.
  • X 10 o (X120), R100 (Ri2o) is sump lambda ⁇ when the respective tube voltages 100 kV (120 kV), which is the output of the reference.
  • the variation in X-ray energy characteristics was defined as the difference between the maximum value and the minimum value of the characteristic ⁇ of each channel in the detector element block. If this variation is 5 or more, ring artifacts are likely to occur in CT images.
  • the grooves of the scintillator plate subjected to groove processing cured by pouring Epoki sheet resin mixed with Ti0 2 powder to form a reflective layer.
  • the scintillator elements were connected at the bottom, so that the shape became inclined toward the center.
  • the upper and lower surfaces were cut and polished to a thickness of 1.8 mm to produce a 24 ⁇ 16 channel scintillator element array.
  • the scintillator element was parallelized on both sides because it was burned by a wire saw, but the reflective layer had a shape in which the upper surface in the Z direction became thin due to curing shrinkage.
  • the angle formed between the side surface of the scintillator element and the substrate surface was 89.43 ° for the end element in the X direction and 89.14 ° for the end element in the Y direction.
  • the scintillator element array produced as described above was adhered to a 24 ⁇ 16 channel photodiode array substrate to obtain a detector element block.
  • This detector element block was mounted on an X-ray CT apparatus so that the X direction was the circumferential direction, and the light emission characteristics were measured as in Example 1.
  • the variation in the conventional element block in both the X and Y directions is 1, the value is about 0.4, indicating that the X-ray CT system using this detector element block can suppress artifacts in CT images. .
  • a phosphor defined by a phosphor element (scintillator element) and reflection layers disposed on both sides thereof
  • a plurality of scintillator element arrays may be prepared, and the radiation detectors may be configured by juxtaposing them in a one-dimensional or two-dimensional direction such that their major axes intersect at a predetermined angle of 0 ° or more and 90 ° or less. .
  • a 24 x 8 channel scintillator element array was produced in the same process as in Example 2, A radiation detector was obtained by adhering to a 24 ⁇ 16 channel photodiode array substrate in a state where these two long axes were arranged in the channel direction at a predetermined angle. At this time, the photodiode of 24 ⁇ 16 channels was divided into regions of 24 ⁇ 8 channels, and each region was inclined at 0.50 ° with respect to the substrate surface.
  • each detector element can be a detection element array oriented in the direction of the X-ray tube focal point, and since it is juxtaposed at a predetermined angle, Variations in the detector array direction can be suppressed over a wider area, and artifacts due to this can be reduced and eliminated.

Landscapes

  • 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)
  • Solid State Image Pick-Up Elements (AREA)
  • Transforming Light Signals Into Electric Signals (AREA)
  • Light Receiving Elements (AREA)
  • Apparatus For Radiation Diagnosis (AREA)

Abstract

Un détecteur de rayonnement comprend une pluralité d'éléments phosphore agencés sur un substrat de réseau de photodiodes en une ou deux dimensions, dans lequel un angle formé par la surface de contact entre un élément phosphore et une couche de réflexion et une surface du substrat peut varier vers l'extrémité de ce substrat de sorte que la direction d'un élément phosphore définie par un élément phosphore (élément scintillateur) et des couches de réflexion situées à l'opposé de celui-ci soit sensiblement identique à la direction incidente du rayonnement. Par conséquent, des variations de caractéristiques d'émission lumineuse, dans un bloc de détection, d'un rayonnement dont le faisceau est appliqué radialement sont limitées de façon à fournir un détecteur de rayonnement de haute précision. Ce détecteur de rayonnement peut fournir un dispositif de diagnostic par imagerie médicale tel qu'un dispositif CT à rayon X avec un minimum d'artefact.
PCT/JP2003/015410 2002-12-02 2003-12-02 Detecteur de rayonnement et dispositif de diagnostic par imagerie medicale Ceased WO2004051312A1 (fr)

Applications Claiming Priority (2)

Application Number Priority Date Filing Date Title
JP2002-349891 2002-12-02
JP2002349891A JP2004184163A (ja) 2002-12-02 2002-12-02 放射線検出器及び医用画像診断装置

Publications (1)

Publication Number Publication Date
WO2004051312A1 true WO2004051312A1 (fr) 2004-06-17

Family

ID=32463056

Family Applications (1)

Application Number Title Priority Date Filing Date
PCT/JP2003/015410 Ceased WO2004051312A1 (fr) 2002-12-02 2003-12-02 Detecteur de rayonnement et dispositif de diagnostic par imagerie medicale

Country Status (2)

Country Link
JP (1) JP2004184163A (fr)
WO (1) WO2004051312A1 (fr)

Cited By (1)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
NL1032782C2 (nl) * 2005-11-01 2013-01-29 Ge Med Sys Global Tech Co Llc Rontgendetector en rontgen-ct-apparatuur.

Families Citing this family (8)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
ITBZ20110002A1 (it) * 2011-01-17 2012-07-18 Microtec Srl Scintillatore per la rilevazione di raggi x e metodo per realizzare un dispositivo di conversione di raggi x in impulsi di luce da utilizzare in uno scintillatore
JP2014510274A (ja) * 2011-03-03 2014-04-24 サン−ゴバン セラミックス アンド プラスティクス，インコーポレイティド 不均一な隔壁を使用した撮像アレイ用のシステム、方法、及び装置
CN105190775B (zh) * 2013-04-01 2018-02-23 株式会社东芝 闪烁器阵列、x射线检测器、及x射线检查装置
CN104345070B (zh) * 2013-07-29 2018-03-23 同方威视技术股份有限公司 探测器模块、安装探测器模块的方法及射线检测系统
CN108428706B (zh) * 2017-02-15 2020-11-20 奕瑞影像科技（太仓）有限公司 一种图像传感器
JP6968419B2 (ja) * 2018-04-27 2021-11-17 国立大学法人東北大学 波面制御素子の製造方法
KR102092889B1 (ko) * 2018-05-17 2020-04-23 국립암센터 방사선 측정기, 이를 포함하는 방사선 측정 시스템 및 방사선 측정방법
JP7467178B2 (ja) * 2020-03-16 2024-04-15 キヤノンメディカルシステムズ株式会社 コリメータ及びコリメータモジュール

Citations (2)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JP2000237178A (ja) * 1999-02-23 2000-09-05 Shimadzu Corp Ｘ線ｃｔ装置
JP2002071815A (ja) * 2000-08-24 2002-03-12 Canon Inc Ｘ線画像撮影装置

Patent Citations (2)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JP2000237178A (ja) * 1999-02-23 2000-09-05 Shimadzu Corp Ｘ線ｃｔ装置
JP2002071815A (ja) * 2000-08-24 2002-03-12 Canon Inc Ｘ線画像撮影装置

Cited By (1)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
NL1032782C2 (nl) * 2005-11-01 2013-01-29 Ge Med Sys Global Tech Co Llc Rontgendetector en rontgen-ct-apparatuur.

Also Published As

Publication number Publication date
JP2004184163A (ja) 2004-07-02

Similar Documents

Publication Publication Date Title
JP4215318B2 (ja) コンピュータ断層撮影装置用シンチレータの製造方法
JP4215320B2 (ja) コンピュータ断層撮影装置用シンチレータの製造方法
US7355180B2 (en) Radiation detector and a method of manufacturing the detector
US9217794B2 (en) Scintillator array, and X-ray detector and X-ray examination device using scintillator array
US7067814B2 (en) Radiation detector and a method of manufacturing the detector
JP4274839B2 (ja) Ｘ線検出器用シンチレータ、その製造方法並びにそれを用いたｘ線検出器及びｘ線ｃｔ装置
WO2007105288A1 (fr) Detecteur de rayonnement et son procede de fabrication
WO2004051312A1 (fr) Detecteur de rayonnement et dispositif de diagnostic par imagerie medicale
US7019298B2 (en) Radiation detector and method of producing the same
US20080093559A1 (en) Arrangement of a Scintillator and an Anti-Scatter-Grid
JPWO2013146167A1 (ja) シンチレータデュアルアレイの製造方法
JP6052595B2 (ja) シンチレータアレイの製造方法
JP2007514143A (ja) Ｐｅｔ検出器用のシンチレーション層
US20250180761A1 (en) Scintillation crystal array, detector, medical imaging device, and manufacturing method
JP2002236182A (ja) 一次元又は多次元検出器アレイの製造方法
EP0112475A1 (fr) Détecteur de rayonnement à canaux multiples et sa méthode de fabrication
JP4817524B2 (ja) Ｘ線固体検出器の製造方法
WO2012147747A1 (fr) Procédé de production d'un réseau de scintillateur
JP2005087366A (ja) マルチスライスｘ線ｃｔ装置用ｘ線検出器およびマルチスライスｘ線ｃｔ装置
JP6052594B2 (ja) シンチレータアレイの製造方法
US20250072843A1 (en) Self contained ct scintillators within a 3-d printed collimator field
JPH06160538A (ja) Ｃｔセンサーの形成方法及びそのｃｔセンサー
JP2000088965A (ja) Ｘ線ｃｔ装置用固体検出器

Legal Events

Date Code Title Description
AK Designated states

Kind code of ref document: A1

Designated state(s): CN US

DFPE Request for preliminary examination filed prior to expiration of 19th month from priority date (pct application filed before 20040101)