JP2006058099A - Radiation scintillator and radiation image detector - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation scintillator that enhances image quality by improving the image resolution. <P>SOLUTION: This radiation scintillator has a scintillator layer 13 of a columnar crystal 31. The film thickness of the scintillator layer 13 is 500μm or larger, and the filling factor of the columnar crystal 31 in the scintillator layer 13 is 70 to 85%. Image resolution is improved, by lowering the filling factor of the columnar crystal 31 in the scintillator layer 13 to 70 to 85%. The lowering of X-ray absorptivity, accompanying the lowering of the filling factor, is compensated by increasing the film thickness of the scintillator layer 13 to 500μm or larger. <P>COPYRIGHT: (C)2006,JPO&NCIPI

Description

  The present invention relates to a radiation scintillator that converts radiation into visible light, and a radiation image detector using the radiation scintillator.

  Conventionally, a radiation image detector is used for medical and industrial nondestructive inspections, and is a planar radiation image detector (hereinafter referred to as FPD) that extracts a radiation image such as an X-ray image or a gamma ray image as an electrical image signal. And an X-ray image tube that extracts a radiation image as a visible image (see, for example, Patent Documents 1 and 2).

  For FPD, for example, a direct method in which electron-hole pairs (e-h pairs) generated in a photoconductive film when X-rays are incident is collected by an electric field and read out as a charge signal, and similar to an X-ray image tube In addition, there are two methods: an indirect method in which X-rays are converted into visible light by a radiation scintillator and read out by a photodiode.

  As shown in FIG. 3, the indirect FPD has a scintillator layer 13 which is a CsI layer composed of a reflective film 12 and a columnar crystal of cesium iodide (CsI) as a phosphor on a substrate 11 through which X-rays are transmitted. A radiation scintillator 14 having at least a radiation scintillator 14 and a radiation scintillator panel 16 as an input means in which a moisture-proof film 15 is sequentially laminated, and a plurality of TFTs (Thin-Film) as switching elements arranged in a matrix on a glass substrate 17 Transistor) and a detector panel 18 as output means in which a photodiode connected to each TFT is laminated. An optical coupling material is injected into the interface between the moisture-proof film 15 and the detector panel 18 as necessary.

  When an indirect FPD is used for medical purposes, the X-ray image 22 transmitted through the human body 21 is transmitted through the substrate 11 and then converted into visible light 24 at the light emitting point 23, and the photodiode of the detector panel 18. It reaches the upper area and is converted into an electrical signal. The electric charge obtained by the photodiode is sequentially read out by the switching operation of the TFT, and after amplification and image processing, it is displayed as an image signal on a monitor (not shown) or stored in a storage medium.

  In such an indirect type FPD, since a solid element is used as a sensor, the problem of noise is the biggest issue. To reduce this noise, the X-ray absorption rate in the scintillator layer 13 is improved. It is important to increase the brightness, and in order to improve the image quality, the image resolution characteristic is important as in the case of the X-ray image tube.

  For this reason, in the radiation scintillator 14, considering the high image quality, the three respective uniformity of image resolution, X-ray absorption rate, and luminance are the most important indexes. As an index of image resolution, the value of CTF in the case of a rectangular wave in which how the intensity of an input signal amount (incident dose) periodically spread in the image space direction is modulated as an output signal is expressed as the most. Commonly used.

  Here, considering three indices excluding uniformity, if only the film thickness is changed in the same process, there is a trade-off relationship with each other. That is, if the film thickness or mass per unit area of the scintillator layer 13 is taken as a parameter, the X-ray absorption rate and luminance increase as the parameter increases, and it becomes easier to obtain an image quality with reduced noise. On the other hand, the image resolution characteristics deteriorate. This is because, as shown in FIG. 3, the visible light 24 from the light emitting point 23 may reach the photodiode via a spread path, so that the thicker the film thickness, the longer the path. This is because the spread of visible light 24 also increases. This is the same as the case where the photocathode is regarded as a photodiode even in the case of an X-ray image tube. Conversely, as the parameter is decreased, the image resolution characteristics are improved, but on the other hand, the X-ray absorption rate and the luminance are decreased.

  In the indirect FPD, cesium iodide (CsI) is most frequently used as the phosphor material, as in the case of the X-ray image tube. There are three reasons why CsI is the most useful material. First, CsI has a property of easily forming a columnar crystal structure by a vacuum deposition method, and as a result, a high image resolution can be obtained, and depending on the film formation conditions, the columnar crystals in the scintillator layer 13 can be obtained. A relatively high film with a filling rate (a value obtained by dividing the actual mass by the theoretical density and the apparent volume) of 85% to 92% is obtained, and the vacuum deposition method is uniform over a large area. A film can be formed at low cost. Second, since CsI has a K absorption edge at 33.2 keV and 36.0 keV, it has a film thickness of 85 to 92% with a filling rate of about 400 to 600 μm and does not require much difficulty in manufacturing, and is usually used for medical purposes. It has a high X-ray absorptivity characteristic in the X-ray energy region of 30 to 90 keV used for the above. Third, CsI is the most widely known radiation scintillator 14 along with gadolinium sulfate (GOS) of about 60,000 photons / MeV, and can obtain the highest light emission. For these three reasons of manufacturing and physical properties, CsI can be said to be the most suitable material for obtaining a radiation scintillator 14 having all of image resolution, X-ray absorption capability, luminance and cost.

In the radiation scintillator 14 mainly composed of CsI columnar crystals, the following three points (1) to (3) are given as target specification values necessary for achieving high image quality. (1) The image resolution (CTF) value is 40% or more at 2 Lp / mm. (2) X-ray absorption is 80% or more with an Al filter having an acceleration voltage of 80 kVp and a thickness of 20 mm. (3) Terbium-activated gadolinium sulfate (GOS) intensifying screen having an acceleration voltage of 80 kVp and a thickness of 20 mm, and having a luminance of 200 μm (about 0.1 g / cm ² ) when applied to an FPD. There must be at least 1.7 times. The X-ray absorptance of (2) is determined only by the mass per unit area of CsI and can be obtained from calculation. As shown in FIG. 4, it is understood that about 0.17 g / cm ² is necessary. If the mass per unit area of CsI is 0.17 g / cm ² or more, X-rays with any energy of 30 keV, 50 keV, and 70 keV in an X-ray spectrum produced by an Al filter having an acceleration voltage of 80 kVp and a thickness of 20 mm are 60. % Is absorbed, and high image quality is secured for various applications.
Japanese Examined Patent Publication No. 7-73031 (page 3, FIG. 1) Japanese Patent No. 3126715 (page 3-4, FIG. 2)

  As described above, in the conventional radiation scintillator 14, considering the trade-off relationship when the film thickness is taken as a parameter, the three indices of image resolution, X-ray absorption rate, and luminance are (1) to (3 ) Target specification value is difficult to satisfy at the same time, and sufficient image quality cannot be obtained.

  The present invention has been made in view of the above points, and provides a radiation scintillator capable of improving image resolution and improving image quality without impairing X-ray transmittance, and a radiation image detector using the radiation scintillator. The purpose is to do.

  The present invention has at least a columnar crystal scintillator layer, the scintillator layer has a thickness of 500 μm or more, and the columnar crystal filling rate in the scintillator layer is 70 to 85%.

  And while improving the image resolution by reducing the filling factor of the columnar crystals in the scintillator layer to 70 to 85%, the X-ray absorption rate associated with the reduction of the filling factor by making the film thickness of the scintillator layer 500 μm or more. Make up for the decline. Therefore, when the film thickness of the scintillator layer is 500 μm or more and the filling rate of columnar crystals in the scintillator layer is 70 to 85%, the image resolution is improved and the image quality is improved without impairing the X-ray transmittance. It becomes possible.

  According to the present invention, the resolution of the columnar crystals in the scintillator layer is reduced to 70 to 85%, thereby improving the image resolution, and by increasing the film thickness of the scintillator layer to 500 μm or more, the X accompanying the decrease in the filling rate Decrease in the linear absorption rate can be compensated. Therefore, the film thickness of the scintillator layer is 500 μm or more, and the filling rate of columnar crystals in the scintillator layer is 70 to 85% without impairing the X-ray transmittance. The image resolution can be improved and the image quality can be improved.

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

  The same components as those described in the background art will be described using the same reference numerals.

  As shown in FIG. 3, an indirect planar radiation image detector (hereinafter referred to as FPD) as a radiation image detector has a reflective film 12, at least a columnar shape, on a substrate 11 that transmits radiation such as X-rays and gamma rays. A radiation scintillator 14 having a crystal scintillator layer 13, a radiation scintillator panel 16 as an input means in which a moisture-proof film 15 is sequentially laminated, and a plurality of TFTs (not shown) as switching elements arranged in a matrix on a glass substrate 17 And a detector panel 18 as an output means in which a photodiode connected to each TFT is laminated. If necessary, an optical coupling material can be injected into the interface between the detector panel 18 and the moisture-proof film 15.

  As the material of the substrate 11, glass, graphite, light metals (beryllium, titanium, aluminum) or alloys thereof, ceramics (alumina, zirconia, beryllia), polymer materials (polycarbonate, polyimide), and the like are used.

  As the material of the reflective film 12, silver alloy, silver, aluminum, gold, copper, barium, platinum, chromium, nickel, titanium oxide, magnesium oxide, or the like is used.

  The scintillator layer 13 has a CsI layer composed of columnar crystals of cesium iodide (CsI) grown in a direction crossing the substrate 11. FIG. 1 is a micrograph (SEM photograph) taken from the surface of a radiation scintillator, wherein 31 is a columnar crystal, 32 is a gap, and the filling rate of the columnar crystal 31 in the scintillator layer 13 (actual mass, theoretical density and apparent density) The value divided by the volume is 70-85%. Further, for the purpose of improving the luminance, a continuous layered phosphor layer of cesium iodide (CsI) may be formed so as to fill a gap between the tip portions of the columnar crystals 31.

  The material of the moisture-proof film 15 includes polybaraxylylene, polyimide, polyurea, polyamide, polyurethane, polyazomethine, nitrocellulose, cellulose acetate, polycarbonate, polyester, polyethylene, polyvinylidene chloride, nylon, vinylidene chloride-vinyl chloride copolymer , Vinylidene chloride-acrylonitrile copolymer, polyolefins, epoxy resins, polyacrylates, or thin films such as silicon oxide, silicon oxynitride, indium oxide, silicon nitride, silicon carbide, alumina against visible light A multilayer film with a transparent inorganic film is used.

  When the indirect FPD is used for medical purposes, the X-ray image 22 transmitted through the human body 21 is transmitted through the substrate 11 and then converted into visible light 24 at the light emitting point 23 of the scintillator layer 13, and the detector panel. It reaches the area on 18 photodiodes and is converted into an electrical signal. The electric charge obtained by the photodiode is sequentially read out by the switching operation of the TFT, and after amplification and image processing, it is displayed as an image signal on a monitor (not shown) or stored in a storage medium.

  The radiation scintillator panel 16 is manufactured by the following procedure when applied to an FPD. The substrate 11 on which the reflective film 12 is formed is set in a vacuum chamber of a vacuum vapor deposition apparatus, and an evaporation source containing CsI is disposed in a direction facing the substrate 11. Then, the temperature of the substrate 11 in the vacuum chamber is adjusted to a condition that a desired columnar crystal and a filling rate can be obtained by the heating pressure adjusting means in the vacuum chamber. Next, the evaporation source is heated to boil CsI, vapor of CsI is sent onto the substrate 11, and a columnar crystal 31 of CsI grown in a crossing direction with respect to the substrate 11 is formed. When the film thickness of the CsI columnar crystal 31 or the mass per unit area reaches a predetermined value, heating of the evaporation source is stopped and film formation is completed. Next, when the temperature of the substrate 11 reaches room temperature, the substrate 11 is taken out from the vacuum chamber of the vacuum evaporation apparatus.

  Next, the substrate 11 on which the scintillator layer 13 is formed is subjected to moisture-proof treatment, and then the target specification values (1) to (3) described above are inspected, and the process proceeds to an assembly process with the detector panel 18.

That is, the target specification values of (1) to (3) are (1) 40% or more when the image resolution (CTF) value is 2 Lp / mm, and (2) an Al filter with an acceleration voltage of 80 kVp and a thickness of 20 mm. The X-ray absorption rate is 80% or more, and (3) the brightness when applied to an FPD with an Al filter having an acceleration voltage of 80 kVp and a thickness of 20 mm has a conventional film thickness of 200 μm (about 0.1 g / cm ² ). The terbium activated gadolinium sulfate (GOS) intensifying screen is 1.7 times or more.

  Next, the evaluation test of each sample 1-11 was implemented about the relationship between the film thickness of the scintillator layer 13, and the filling rate of the columnar crystal 31 in the scintillator layer 13, and the result is shown in FIG. In this evaluation test, a glass plate having a thickness of 0.5 mm was used for the substrate 11, and after a silver alloy reflective film 12 and a magnesium oxide protective film for protecting the reflective film 12 were laminated on the substrate 11, After thallium-doped cesium iodide (CsI / TI) was vacuum-deposited as scintillator layer 13 and polybarxylylene was deposited to a thickness of 5 μm as a moisture-proof film 15 by a CVD method, an evaluation test was performed to see if the target specification value was satisfied. .

  In Samples 1 to 4, aiming at a filling rate of 85 to 92%, the resulting filling rates were 87.7%, 88.4%, 88.1%, and 88.7%, all from a filling rate of 85% It became high. FIG. 5 shows a micrograph (SEM photograph) having a filling rate of about 88%.

  In Samples 5 to 11, aiming at a filling rate of 70 to 85%, the resulting filling rate was 82.2%, 80.9%, 81.1%, 73.2%, 72.5%, 73.4%. 73.3%, both of which were in the range of 70 to 85% filling rate.

  From the results of the evaluation test shown in FIG. 2, the samples 1 to 4 having a filling rate higher than 85% are either (1) or (2), which is either the image resolution CTF or the X-ray absorption rate regardless of the film thickness. The specification value was not met.

  Among Samples 5 to 11 having a filling rate of 70 to 85%, Samples 5, 6, 8, and 9 having a film thickness of less than 500 μm did not satisfy the target specification value of X-ray absorption rate (2).

  Of Samples 5 to 11 having a filling rate of 70 to 85%, Samples 7, 10, and 11 having a film thickness of 500 μm or more satisfied all the target specification values (1) to (3).

  Thus, at the expense of a decrease in the X-ray absorption rate, the image resolution can be improved by reducing the filling rate of the columnar crystals 31 in the scintillator layer 13 to 70 to 85%. By increasing the film thickness of the scintillator layer 13 to 500 μm or more, it was possible to compensate for the decrease in the X-ray absorption rate accompanying the decrease in the filling rate. Therefore, by setting the film thickness of the scintillator layer 13 to 500 μm or more and the filling rate of the columnar crystals 31 in the scintillator layer 13 to 70 to 85%, the image resolution is improved without impairing the X-ray transmittance, and high The image quality was improved.

  When the filling rate is smaller than 70%, even if the thickness of the scintillator layer 13 is increased to 500 μm or more, the decrease in the X-ray absorption rate cannot be sufficiently compensated, and the filling rate is larger than 85%. In this case, although the X-ray absorption rate increases as in the conventional case, the image resolution decreases, and when the film thickness is smaller than 500 μm, the decrease in the X-ray absorption rate cannot be sufficiently compensated.

  By applying such a radiation scintillator to the FPD, the uniformity of each of the image resolution, the X-ray absorption rate, and the luminance is improved, and the image quality can be improved.

  Further, when such a radiation scintillator is applied to an X-ray image tube, since the phosphor additive used for the scintillator layer 13 is sodium iodide, the emission wavelength is different, and the target specification value of luminance is not particularly set. The target specification value can be satisfied by setting the thickness of the scintillator layer 13 to 500 μm or more and setting the filling rate of the columnar crystals 31 in the scintillator layer 13 to 70 to 85%. In the case of an X-ray image tube, an input surface that converts an X-ray image into an electronic image and an output surface that converts an electronic image converted on the input surface into a visible image are arranged in a vacuum envelope. Yes, the input surface is the input means 16, and the output surface is the output means 18.

  Therefore, the radiation scintillator having the CsI columnar crystal 31 has a film thickness of the scintillator layer 13 of 500 μm or more and the columnar crystal in the scintillator layer 13 regardless of whether it is applied to any radiation image detector of an FPD or an X-ray image tube. By setting the filling rate of 31 to 70 to 85%, the target specification value can be satisfied.

  When manufacturing a radiation scintillator to be applied to an FPD, a detector panel 18 having a TFT array and a photodiode formed on a glass substrate 17 is set in a vacuum deposition apparatus, and a scintillator layer 13 is formed on the detector panel 18. It may be formed.

  Further, as the protective film for protecting the sales company film 12, in addition to magnesium oxide, silicon oxide, titanium oxide, calcium fluoride, magnesium fluoride, or the material of the moisture-proof film 15 may be used.

It is the microscope picture (SEM photograph) which image | photographed the radiation scintillator which shows one embodiment of this invention from the surface. It is explanatory drawing which shows the result of the evaluation test of each sample about the relationship between the film thickness of a radiation scintillator, and the filling rate of the columnar crystal in a scintillator layer. It is sectional drawing of a radiographic image detector. It is a graph which shows the X-ray absorption rate with respect to the mass per unit area of CsI. It is the microscope picture (SEM photograph) which image | photographed the conventional radiation scintillator from the surface.

Explanation of symbols

13 Scintillator layer
14 Radiation scintillator
16 Radiation scintillator panel as input means
18 Detector panel as output means
31 columnar crystals

Claims (2)

Having at least a columnar scintillator layer;
A radiation scintillator characterized in that the film thickness of the scintillator layer is 500 μm or more and the filling rate of columnar crystals in the scintillator layer is 70 to 85%. Input means having a radiation scintillator according to claim 1 for converting radiation into visible light;
An output unit that outputs image information based on visible light converted by the radiation scintillator.