JP6672621B2 - Radiation image conversion screen and radiation detection device - Google Patents

Description

  The present invention relates to a radiation image conversion screen and a radiation detection device.

  Image diagnosis by X-ray transmission imaging and image diagnosis by X-ray CT (Computed Tomography) imaging are used as effective diagnostic methods because they can detect lesions such as cancer and ulcers inside a subject.

  In these image diagnoses, an X-ray image is generated by measuring the dose of X-rays transmitted through the subject (or the dose of X-rays absorbed by the subject) and expressing an increase or decrease in the dose in an image. . The elements constituting the living body are relatively light elements such as hydrogen, nitrogen, oxygen, calcium and the like, and their X-ray absorption coefficients are small. Therefore, in order to generate a clear X-ray image when a living body is used as a subject, it is necessary to increase the amount of X-rays applied to the subject, and exposure of the subject becomes a problem.

  In particular, when an X-ray CT image of a cancer tissue is taken, the X-ray absorption amount of the cancer tissue and the X-ray absorption amount of the normal tissue are almost the same. Is difficult to identify. Therefore, when identifying a cancer tissue, an iodine contrast agent is injected into a human body and X-ray CT imaging is performed. This is based on the fact that cancer tissues have more blood vessels than normal tissues. The ratio of the injected iodine contrast agent staying in the cancer tissue having many blood vessels becomes larger than that in the normal tissue. Therefore, a tissue having a large X-ray absorption amount can be identified as a cancer tissue.

  However, even in the above-described method in which the difference in the amount of X-ray absorption is increased, it may be difficult to identify, for example, calcium attached to blood vessels.

  In order to solve this, usually, X-ray irradiation is performed twice while changing the X-ray tube voltage to perform CT measurement. However, in this method, the amount of exposure of the subject becomes a problem.

  This is either a radiation detector that converts X-rays directly into electric signals (direct method) or a radiation detector that converts X-rays into light with a scintillator layer and converts the light into electric signals (indirect method). Even when used, there is a problem.

  In order to solve such a problem, for example, Patent Document 1 discloses that in a direct method, four element detectors are arranged in a line in the X-ray incident direction, and the second and third elements from the X-ray incident side of the element detectors are arranged. Discloses a method of obtaining energy information by installing an X-ray absorber between them and changing the response function of a third element detector.

  However, in the method of Patent Document 1, for example, the current output from an element detector installed at the rear is reduced to be easily affected by noise, or backscattered X-rays are incident on a front element detector. As a result, it may be difficult to accurately obtain energy information.

  In addition, in order to cope with fan beam X-rays and cone beam X-rays used in the third generation CT, it is necessary to arrange a detector group having a plurality of element detectors in a depth direction in a two-dimensional array. Accompany. In addition, it is difficult to arrange the element detectors in the depth direction with, for example, 1 mm square or less, and the spatial resolution as a CT detector is insufficient.

  In order to solve such a problem, Non-Patent Document 1 discloses that two types of absorbers are orthogonally arranged on the X-ray incident side of a two-dimensional detector such as a flat panel detector and X-rays having four different energy spectra. A technique has been proposed in which X-ray energy information is acquired with high accuracy by making a line incident on a two-dimensional detector.

JP 2013-57554 A

Workshop "Radiation detectors and their applications" (28th) Abstracts, High Energy Accelerator Research Organization Radiation Science Center, January 28, 2014, pp. 13-14

  However, in the method described in Non-Patent Document 1, it is assumed that the size of a region defined by using two types of absorbers is larger than the size of one pixel (pixel) of the two-dimensional detector. Therefore, the high spatial resolution of the two-dimensional detector cannot be effectively used.

  The present invention has been made in view of the above-described conventional situation, and provides a radiation image conversion screen and a radiation detection device capable of accurately acquiring X-ray energy information and improving the spatial resolution of a CT image. The task is to

  As a result of intensive studies, the present inventor can solve the above-described problem by arranging two or more types of radiation absorption filters one-dimensionally and periodically on the radiation incident side in a radiation detection device or a radiation image conversion screen. And arrived at the present invention.

  That is, a first gist of the present invention is to provide a radiation detection apparatus having a radiation detector array in which a plurality of radiation detectors for converting radiation energy into electric charges are arranged vertically and horizontally, and two or more radiation detectors are provided on a radiation incident side. The present invention relates to a radiation detection device in which absorption filters are arranged one-dimensionally and periodically.

  A second gist of the present invention is a radiation image conversion screen having a layer containing a phosphor, wherein two or more radiation absorbing filters are arranged one-dimensionally and periodically on the layer containing the phosphor. And a radiation image conversion screen.

  According to the present invention, it is possible to provide a radiation image conversion screen and a radiation detection device that acquire X-ray energy information with high accuracy and improve the spatial resolution of an obtained CT image. Also, since the phosphors that can be used for X-ray (CT) are limited, the range of energy discrimination may be limited. On the other hand, the method of using a filter involves selection of a substance that can be used as a filter. The width (type, price, thickness, K absorption edge, etc.) is extremely wide, and the degree of freedom in device design and the range of energy discrimination are wide and high precision.

It is a schematic structure figure of an X-ray inspection device concerning an embodiment of the present invention. It is the front view which looked at the X-ray detection apparatus from the X-ray incidence side. FIG. 2 is an exploded perspective view of the X-ray detection device. FIG. 2 is a partially enlarged view of the X-ray detection device. It is the front view which looked at the X-ray detection device concerning another embodiment from the X-ray incidence side. It is an exploded perspective view of the X-ray detection device concerning another embodiment. It is the front view which looked at the X-ray detector concerning yet another embodiment from the X-ray incidence side. FIG. 11 is an exploded perspective view of an X-ray detection device according to still another embodiment. It is the schematic of the X-ray detection apparatus used in the experiment. It is a figure showing an experimental system. FIG. 5 is a profile diagram of a signal amount in a lateral direction of a subject. (A) is a diagram showing a CT image reconstructed using the number of X-rays in the energy range E2 when the absorber according to the experimental example is used, and (b) is a case using the absorber according to the comparative example. It is a figure showing a radiation absorption filter by a comparative example. 13 is a graph showing CT value profiles of the energy-resolved CT images of FIGS. 12A and 12B and CT value profiles obtained by calculation.

  Hereinafter, embodiments of the present invention will be described with reference to the drawings.

  FIG. 1 is a schematic configuration diagram of an X-ray inspection apparatus according to an embodiment of the present invention, FIG. 2 is a front view of the X-ray detection apparatus, and FIG. 3 is an exploded perspective view of the X-ray detection apparatus.

  As shown in FIG. 1, an X-ray inspection apparatus 1 includes an X-ray tube 2 that irradiates X-rays toward a subject S, an X-ray detection device 3 that detects X-rays transmitted through the subject S, and an X-ray detector. The apparatus includes a current measuring device 6, an arithmetic device 7, and an imaging device 8 for measuring an output current from the device, and performs energy-resolved CT measurement.

  X-rays emitted from the X-ray tube 2 are not particularly limited, but white X-rays emitted by, for example, colliding electrons accelerated to 120 kVp with a tungsten target and emitting are preferable. When X-rays are emitted from the X-ray tube 2 toward the subject S, the X-rays transmitted through the subject S enter the X-ray detection device 3.

  The X-ray detection device 3 includes a radiation detector array 5 in which a number of radiation detectors (X-ray detectors) 50 that generate electric charges by energy given from incident X-rays are arranged vertically and horizontally. On the radiation incident side of the array 5, two or more types of radiation absorbing filters are arranged one-dimensionally and periodically.

  For example, as shown in FIGS. 2 and 3, the radiation absorption filter includes a plurality of absorbers 4A and 4B arranged along a horizontal direction (the left-right direction in FIG. 2). That is, two types of striped absorbers 4A and 4B are arranged in parallel. The absorber 4A and the absorber 4B have different radiation absorptivity. The width in the short direction of the absorber 4A and the absorber 4B is larger than the width of the radiation detector 50. The longitudinal length of the absorber 4A and the absorber 4B is larger than the horizontal length of the radiation detector array 5. The longitudinal length of the absorber 4A and the longitudinal length of the absorber 4B may be the same or different.

  A plurality of absorbers 4A and absorbers 4B are arranged at intervals. The arrangement interval between the absorber 4A and the absorber 4B is larger than the width of the radiation detector 50. The absorber 4A and the absorber 4B are fixed, for example, surrounded by a frame (not shown).

  In the example shown in FIGS. 2 and 3, the width W in the short direction of the absorber 4A and the absorber 4B is equal. The absorber 4A and the absorber 4B are arranged at intervals W in the vertical direction (the vertical direction in the figure perpendicular to the horizontal direction). The absorber 4A is arranged on the X-ray incident side of the absorber 4B. Further, the absorber 4A and the absorber 4B are arranged to be shifted by W / 2 in the vertical direction.

  Thus, the radiation absorbing filter has a region R1 having neither the absorber 4A nor the absorber 4B, a region R2 having only the absorber 4A, a region R3 having only the absorber 4B, and a region R4 having the absorber 4A and the absorber 4B overlapping. These four regions R1 to R4 have different radiation absorptances. In the radiation absorption filter, regions R1 to R4 having different radiation absorptances are arranged one-dimensionally and periodically.

  As the absorber 4A and the absorber 4B, a material containing an element from Al having an atomic number of 13 to Bi having an atomic number of 83 can be used. It is conceivable to use a material made of an element having an atomic number smaller than that of aluminum, but in that case, the thickness of the absorber increases, which is not practical. Further, when a material containing Tc having an atomic number of 43, Pm having an atomic number of 61, and an element having an atomic number of 84 or more has a radioactive isotope, it is not suitable as an absorber used in the present invention.

  When measuring the iodine contrast agent in the subject S, tin is typically used as a material of the absorber 4A or the absorber 4B, but is not limited to tin, and may be aluminum, iron, iodine, tantalum, or the like. Can also be used. For example, when tin is used as the absorber 4A or the absorber 4B, it is typically preferable to use a tin foil (metal foil), but not limited to the metal foil, and a powder of an oxide or the like is filled in a thin container made of an acrylic resin. It is also possible to use a cured product or a product cured using an acrylic resin.

  In addition, for example, when measuring a barium contrast agent in the subject S, aluminum, tin, silver, or the like is preferable as the absorber 4A or the absorber 4B. When measuring the gold contrast agent in the subject S, aluminum, gold, tantalum, tungsten, or the like is preferable as the absorber 4A or the absorber 4B.

  FIG. 4 is a partially enlarged view of the X-ray detection device 3 when viewed from the radiation incident side, and shows boundaries of the radiation detector 50 for explanation. By arranging the absorber 4A and the absorber 4B in parallel in the horizontal direction by shifting the position in the vertical direction, as shown in FIG. 4, the radiation detector 50 has neither the absorber 4A nor the absorber 4B. Detector α corresponding to the region, detector β corresponding to the region of only absorber 4A, detector γ corresponding to the region of only absorber 4B, detector corresponding to the overlapped region of absorber 4A and absorber 4B δ. The detectors α, β, γ, and δ have four different incident X-ray energy spectra, and thus are four types of detectors having different response functions.

  The radiation detector array 5 is composed of a direct conversion flat panel detector (Flat Panel Detector) using Si, CdTe or CdZnTe as a base material, and acquires X-rays as direct charges (current). The radiation detectors 50 (detectors α, β, γ, and δ) are arranged at the same distance from the X-ray tube 2, and X-rays having passed through filters having different radiation absorptivity are incident thereon. Alternatively, the radiation detector array 5 includes a radiation detector (detectors α, β, γ, and δ) having an X-ray conversion layer made of an amorphous selenium semiconductor or the like, and the X-rays incident on the X-ray conversion layer are electrons. And the holes charge the storage capacitor. The detected charges stored in the storage capacitor are read out corresponding to the pixel (each detector) by turning on a thin film transistor switch made of amorphous silicon or the like.

  The current measuring device 6 amplifies and measures the electric charge output from the radiation detector 50, and outputs the measurement result to the arithmetic device 7. The arithmetic unit 7 uses the current values detected by the four types of radiation detectors 50 (detectors α, β, γ, and δ) and the response function, and performs the unfolding method to detect the detectors α, β, γ, An X-ray energy distribution for every four pixels corresponding to δ is obtained. The response function of each detector is obtained in advance by using a substance having a known thickness as the subject S, irradiating it with X-rays, and measuring. Further, the response function may be obtained by calculation using a deterministic method. In the deterministic method, the X-ray intensity after passing through a substance is calculated as in the following equation (1) based on the mass attenuation coefficient, density, and thickness of the substance.

Here, Y ₀ (E) is the number of X-rays having energy E at the exit of the X-ray tube 2, Y (E) is the number of X-rays having energy E incident on the radiation detector 50, and μ _i (E ) / Ρ _i , ρ _i , and t _i are the mass attenuation coefficient, density, and thickness of the substance i, respectively.

  The arithmetic unit 7 obtains an X-ray intensity distribution in the radiation detector array 5 for each set energy range. The energy range is set, for example, as shown in Table 1. It is also possible to set the energy range again for the same measurement and analyze the same. The X-ray tube 2 and the X-ray detection device 3 are arranged to face each other, and irradiate the subject S with X-rays from various directions. The arithmetic unit 7 obtains an X-ray intensity distribution for each energy range of the radiation detector array 5 in each direction (rotation angle). The imaging device 8 constructs (reconstructs) a CT image for each energy range using the calculation result of the calculation device 7.

  As described above, according to the present embodiment, the radiation detectors 50 are two-dimensionally arranged so as to be located on a plane at the same distance from the X-ray tube 2. Therefore, as in the case where the radiation detectors are arranged in a line in the X-ray incident direction, the measured current value from the detector installed at the rear is reduced to the same level as the noise current value, and the backscattered X-ray is reduced. A phenomenon such as incidence on a front detector does not occur, and X-ray energy information can be acquired with high accuracy. In particular, when calculating energy information of X-rays using a response function obtained by a deterministic method, the accuracy of the calculation result can be improved.

  Further, by using a flat panel detector for the radiation detector array 5, a large number of radiation detectors 50 and a current readout circuit can be integrally configured. Also, by disposing two types of absorbers 4A and 4B in parallel on the radiation incident surface side of the radiation detector array 5, three types of radiation absorption rates (four types if including no absorber) are obtained. It is possible to easily provide a radiation absorbing filter having the same. In the radiation absorbing filter having such a configuration, the width and the arrangement interval of the absorbers 4A and 4B can be easily changed, and the width and the arrangement interval can be set to be suitable in consideration of attenuation by the absorbers 4A and 4B. .

  Further, by arranging the absorbers 4A and 4B in parallel at different positions, the spatial resolution in the horizontal direction of the CT image can maintain the high spatial resolution originally possessed by the radiation detector array 5.

  On the other hand, the vertical spatial resolution of the CT image is coarser than the spatial resolution originally held by the radiation detector array 5. However, in the CT for the human body, the human body is scanned while moving in the body axis direction. By this operation, the spatial resolution in the body axis direction can be made smaller than the length of the combination of the absorbers 4A and 4B.

  Furthermore, in order to cope with a helical scan that rotates in a spiral around the human body, it is possible to provide the absorber with a certain angle and to install the absorber in front of the radiation detector 50.

  In the above-described embodiment, an example in which a flat panel detector is used for the radiation detector array 5 has been described. However, the present invention is not limited to a detector that reads out the amount of electric charge at every set time like a flat panel detector. It can also be realized using a detector that stores the applied energy. For example, an imaging plate, a thermoluminescence plate, a glass dosimeter and the like can be used. However, these detectors need to irradiate laser or raise the temperature after X-ray irradiation to read out the output in proportion to the amount of stored energy. It is preferable to use the method for simply measuring the substance distribution inside the subject by transmission imaging.

  As shown in FIGS. 5 and 6, the width of the absorber 4B may be twice the width W of the absorber 4A. The absorber 4B is arranged at an interval of 2W in the vertical direction. The absorbers 4A are arranged so that those that overlap the absorber 4B and those that do not overlap are alternately arranged. Even with such a configuration, the region R1 having neither the absorber 4A nor the absorber 4B, the region R2 having only the absorber 4A, the region R3 having only the absorber 4B, and the region R4 having the absorber 4A and the absorber 4B overlapped. As described above, the regions R1 to R4 having different radiation absorptivity can be provided one-dimensionally and periodically.

  As shown in FIGS. 7 and 8, four types of absorbers 4A to 4D having different radiation absorptances may be arranged in the vertical direction. The absorbers 4A to 4D have the same size, and are arranged in order in the vertical direction. Even with such a configuration, four regions having different radiation absorptivity can be provided one-dimensionally and periodically.

  7 and 8, any one of the absorbers 4A to 4D may be omitted. That is, any one of the absorbers 4A to 4D may be air. Even with such a configuration, four regions having different radiation absorptivity can be provided one-dimensionally and periodically.

<Indirect conversion method>
The radiation detector array 5 may use an indirect conversion type (scintillator type) flat panel detector instead of the direct conversion type. In this case, the X-ray detection device 3 includes a substrate on which a thin film transistor switch and a photoelectric conversion element are formed, and a radiation image conversion screen formed on the substrate. The radiation image conversion screen includes a phosphor-containing layer (scintillator layer), and two or more radiation absorption filters are arranged one-dimensionally and periodically on the phosphor-containing layer. As the radiation absorbing filter, those shown in FIGS. 2 and 3 can be used. Further, it is also possible to assign the role of the radiation absorption filter to the scintillator, and in this case, the shape of the scintillator is as shown in FIGS.

[substrate]
The substrate used for the X-ray detection device is not particularly limited as long as it is a substrate having radiation transparency, and a known material, for example, various kinds of glass, polymer material, metal, or the like can be used.

  For example, plate glass made of glass such as quartz, borosilicate glass, and chemically strengthened glass; ceramic substrate made of ceramic such as sapphire, silicon nitride, and silicon carbide; semiconductor made of silicon, germanium, gallium arsenide, gallium phosphide, gallium nitrogen, and the like Semiconductor substrate consisting of; cellulose acetate film, polyester film, polyethylene terephthalate film, polyamide film, polyimide film, triacetate film, polycarbonate film, carbon fiber reinforced resin sheet, etc .; polymer film (plastic film); aluminum sheet, iron sheet, copper A metal sheet such as a sheet; a metal sheet having a metal oxide coating layer, an amorphous carbon substrate, or the like can be used.

  Since the weight of the radiation image conversion screen is being reduced in view of the convenience of carrying the radiation image conversion screen, the thickness of the substrate is preferably 2.0 mm or less, more preferably 1.0 mm or less. .

  Further, when the substrate has a high reflectance, the sensitivity of the flat panel detector of the radiation image conversion screen can be increased by reflecting the visible light emitted by the phosphor toward the photoelectric conversion element, which is preferable. In order to increase the reflectance, it is possible to use a substrate in which a pigment or ceramic powder is dispersed in the above-mentioned glass or polymer film, and a ceramic having a high reflectance is provided between the substrate and the phosphor layer. And a buffer layer such as a metal film.

[Scintillator layer]
The scintillator layer in the present invention corresponds to the phosphor-containing layer. Note that the phosphor in the present invention refers to an electromagnetic wave having a wavelength of 300 nm to 800 nm, that is, an electromagnetic wave having a wavelength ranging from 300 nm to 800 nm, that is, an electromagnetic wave ranging from ultraviolet light to infrared light (light). ).

  The type of the phosphor contained in the scintillator layer is not particularly limited as long as the effects of the present invention are not impaired, but those having high conversion efficiency from X-rays to visible light and hardly absorbing the emitted visible light are preferable. .

For example, as described in JP-A-2000-162394 and JP 2003-82347, the gadolinium oxysulfide phosphor _{(Gd 2 O 2} S), terbium (Tb), Jirupuroshiumu (Dy), cerium (Ce) And the like containing an activator such as

  Further, as described in JP-A-2006-38870, CsI, a mixture of CsI and sodium iodide (NaI) at an arbitrary molar ratio, or a mixture of CsI and CsI as described in JP-A-2001-59899. CsI containing an activator such as indium (In), thallium (Tl), lithium (Li), potassium (K), rubidium (Rb), and sodium (Na) may be used.

One of the above phosphors may be used alone, or two or more different phosphors may be used in combination.
Further, the phosphor contained in the scintillator layer may be a columnar crystal, a single crystal, or a sintered body.

(Phosphor filling rate)
The filling rate of the phosphor in the scintillator layer is usually 40% or more, preferably 50% or more, more preferably 60% or more, and usually 100% or less. It is preferable if it is within this range since the incident X-rays can be efficiently converted into visible light.

The method for measuring the filling rate is as follows.
The weight (W ₀ ) of the substrate before forming the scintillator layer and the weight (W ₁ ) after forming the scintillator layer are measured. The weight per unit volume is calculated from the thickness and area of the formed layer containing the phosphor and the measured weight of the film (W ₁ −W ₀ ).

  On the other hand, the weight per unit volume of the phosphor contained in the layer is calculated from the weight ratio between the phosphor and the medium. From the calculated content per unit volume of the phosphor and the specific gravity of the phosphor, it is possible to calculate the filling rate (volume%) of the phosphor.

[Method of forming scintillator layer]
Examples of the method for forming the scintillator include a method of forming a layer by a vacuum evaporation method, a wet film formation method, and a sintering method. Hereinafter, a method for forming a wet film-forming method using the phosphor-containing composition will be described in detail, but the present invention is not limited thereto. Hereinafter, the case of the wet film forming method will be described.

  The method for forming a scintillator layer in the present invention includes a coating step and a drying step. After the drying step, a post-treatment step (a step of performing washing and drying again) may optionally be included.

(Coating process)
In the coating step of the present invention, a phosphor-containing composition is used. The phosphor-containing composition of the present invention may be in the form of a powder or a slurry.

  When the film is formed by a wet film forming method, a composition containing the phosphor and, if necessary, other media, for example, a binder resin, a dispersant, a plasticizer, a photopolymerization initiator / a thermal polymerization initiator, or the like is used. Further, if necessary, an organic solvent or the like may be contained in order to adjust the viscosity of the composition.

  The binder resin that may be contained in the phosphor-containing composition is not particularly limited as long as the effects of the present invention are not impaired.For example, besides nitrified cotton, cellulose acetate, ethyl cellulose, polyvinyl butyral, and linear Polysaccharides such as polyester, polyvinyl acetate, vinylidene chloride-vinyl chloride copolymer, vinyl chloride-vinyl acetate copolymer, polyalkyl- (meth) acrylate, polycarbonate, polyurethane, cellulose acetate butyrate, polyvinyl alcohol, gelatin, dextrin, and gum arabic And the like.

  The dispersant is not particularly limited as long as the effects of the present invention are not impaired, and examples thereof include phthalic acid and stearic acid. Further, examples of the plasticizer include triphenyl phosphate and diethyl phthalate. One of these materials may be used alone, or two or more different materials may be used in combination.

  The organic solvent that may be contained in the phosphor-containing composition is not particularly limited as long as it can dissolve or disperse the phosphor and the medium, for example, ethanol, methyl ethyl ether, butyl acetate, and ethyl acetate. , Ethyl ether, xylene and the like. As the organic solvent, one type may be used alone, or two or more different types may be used in combination.

  A coating film is formed using the composition (or coating solution) adjusted as described above. The method for forming the coating film is not particularly limited as long as the effects of the present invention are not impaired, and known techniques can be applied. Examples thereof include a die coating method, a screen printing method, an inkjet method, and a spin coater method.

  In the case of a photocurable composition (or a coating liquid) after forming a coating film, a method of curing the coating film by using a UV irradiation device or the like through a heating and drying step, if necessary, may be mentioned. In the case of a thermosetting composition (or a coating solution), a method of curing using a hot plate, a hot air drier, or the like may be used.

  The scintillator layer may be formed of a plurality of films by overlapping two or more different films. The two or more different layers include, for example, those having different average particle sizes and particle size distributions of the phosphors, and different media contained.

[Protective layer]
After the formation of the scintillator layer, a protective layer may be further formed on the scintillator layer. The material for forming the protective layer is not particularly limited as long as the effects of the present invention are not impaired. For example, (A) urethane (meth) acrylate, (B) monofunctional (meth) acrylate, (C) polyfunctional (meth) A) A cured product of a radiation-curable composition containing an acrylate, and a film such as PET having an adhesive layer. The above-mentioned radiation-curable composition may optionally contain materials other than those described above.

  Hereinafter, the present invention will be described more specifically with reference to experimental examples that replace the examples of the present invention. The present invention is not limited to these.

  An experiment of energy-resolved CT measurement was performed using an X-ray detector having a radiation absorption filter in which two types of absorbers were arranged in parallel. As shown in FIG. 9, in the X-ray detector, a tin absorber having a width of 1 mm and a thickness of 100 μm and a zirconium absorber having a width of 2 mm and a thickness of 100 μm were arranged in parallel on the surface of a flat panel detector. The zirconium absorbers were arranged at 2 mm intervals, and the tin absorbers were arranged at 1 mm intervals. The measured value of each area is used as the detector number of the radiation detector, 1ch (detector α corresponding to the area without either absorber), 2ch (detector β corresponding to the area only with the zirconium absorber), 3ch (tin absorption) It is considered that the measurement was performed using a detector γ corresponding to the body-only region and 4ch (a detector δ corresponding to the region where the zirconium absorber and the tin absorber overlapped). The flat panel detector was 50 mm × 50 mm and the pixel size was 48 μm × 48 μm.

  FIG. 10 is a diagram showing an experimental system. The X-ray tube was equipped with an aluminum filter having a thickness of 2 mm, the tube voltage was set to 120 kV, and the tube current was set to 3.0 mA. The acrylic phantom was cylindrical and provided with an iodine region having a diameter of 30 mm and a diameter of 5 mm at the center. The concentration of iodine was 30 μm per 5 mm of X-ray transmission distance.

  First, the acrylic thickness was changed from 12 to 42 mm at 10 mm intervals, the iodine thickness was changed from 0 μm, 15 μm, 30 μm, 45 μm, and 60 μm, and the signal amount of the flat panel detector with respect to the X-rays passed for all combinations was measured. And the response function was determined.

  Next, the amount of light emitted by X-rays transmitted through the cylindrical acrylic phantom in the arrangement shown in FIG. 10 was measured.

  The output profile along the diameter of the cylindrical acrylic phantom is shown in FIG. The energy ranges in Table 2 below were set for this output profile, and the number of X-rays in each energy range was determined by unfolding.

  In performing the calculation, the rotation of 10 degrees and the measurement of the X-ray number distribution are repeated 17 times, and the rotation measurement is performed by performing the measurement 18 times in total. The acrylic phantom and the iodine region, which are the subject, are both cylindrical. Therefore, assuming that substantially the same data can be obtained even when rotated by 10 degrees, the following analysis was performed by regarding one measurement result as 18 pieces of measurement data.

  FIG. 12A shows a CT image reconstructed by the filter back projection method using the number of X-rays in the energy range E2. Energy range E2 is the energy range above the K-absorption edge of iodine and was sensitive to iodine attenuation of X-rays.

[Comparative example]
In the above-mentioned experimental example, as shown in FIG. 13, the CT was carried out in the same manner as in the experimental example, except that a tin absorber and a zirconium absorber having a width of 1 mm and a thickness of 100 μm were arranged in a lattice as shown in FIG. Images were reconstructed. The CT image is shown in FIG.

  In this comparative example, since the width of the absorber was 1 mm and the size of one pixel was 2 mm square, the image was rough. On the other hand, in the image according to the experimental example, since the spatial resolution of the pixel of the flat panel detector was obtained, it was confirmed that the image was smooth.

  FIG. 14 shows a CT value profile along the diameter of the phantom in the CT images of FIGS. 12A and 12B. In addition, a CT value profile obtained by calculation is also shown. The CT value is not a Hounsfield unit that is common in the medical field, but is represented by its linear attenuation coefficient.

  In the CT value profile of FIG. 14, when the radiation absorption filter according to the comparative example was used, irregularities in the CT value were observed. On the other hand, when the radiation absorption filter of the experimental example according to the present invention was used, the change in the CT value was smooth, and the transition from the acrylic portion to the iodine portion was clear.

DESCRIPTION OF SYMBOLS 1 X-ray inspection apparatus 2 X-ray tube 3 X-ray detectors 4A to 4D Absorber 5 Radiation detector array 6 Current measuring device 7 Computing device 8 Imaging device 50 Radiation detector

Claims (6)

In a radiation detection device having a radiation detector array in which a plurality of radiation detectors that convert radiation energy into electric charges are arranged vertically and horizontally,
The radiation incidence side, two or more radiation absorbing filter is disposed one-dimensionally periodic, characterized Rukoto which have a and overlapping areas, the radiation detecting apparatus.   The radiation detecting apparatus according to claim 1, wherein the two or more radiation absorbing filters in a stripe shape are arranged in parallel. A radiation image conversion screen having a layer containing a phosphor,
On the layer containing the phosphor, two or more types of radiation absorbing filters are arranged one-dimensionally and periodically, have overlapping regions , and include three or more regions having different radiation absorptivity. Radiation image conversion screen.   The radiation image conversion screen according to claim 3, wherein the radiation absorption filters are arranged in a stripe shape.   The radiation image conversion screen according to claim 3, wherein the two or more radiation absorbing filters in a stripe shape are arranged in parallel.   A radiation detection apparatus comprising: the radiation image conversion screen according to claim 3; and a photoelectric converter.

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