JP2006208158A - Radiographic image transformation panel, radiographic image detector, and radiographic image photographing system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiographic image transformation panel of high phototransfer rate that is free from image uneveness. <P>SOLUTION: This radiographic image transformation panel 1a is provided with a phosphor particle 11 for absorbing energy of an incident radiation, a charge generation layer for generation a charge by direct or indirect energy transfer from the phosphor particle 11, a charge accumulating element for accumulating the charge obtained in the charge generation layer, an image signal output element for outputting a signal based on the charges accumulated by the charge accumulating element, and a substrate for supporting the constitutive elements hereinbefore, and outputs an image signal of the input radiation, based on the signal output from the image signal output element. In the panel 1a, the charge generation layer contains a conductive polymer compound, and a nanocarbon/π-conjugated compound complex. <P>COPYRIGHT: (C)2006,JPO&NCIPI

Description

  The present invention relates to a radiation image conversion panel, a radiation image detector, and a radiation image capturing system, and more particularly, to a radiation image conversion panel, a radiation image detector, and a radiation image capturing system used in a digital radiation image capturing system.

  Conventionally, radiation images obtained by irradiating a subject with radiation such as X-rays and detecting the intensity distribution of the radiation transmitted through the subject have been widely used.

  As a radiographic imaging system for obtaining a radiographic image, an SF system (screen film system) that obtains a radiographic image by combining a fluorescent intensifying screen and a radiographic film has been used. A digital radiographic imaging system using a CR (Computed Radiography), an FPD (Flat Panel Detector) or the like that detects and converts the signal into an electrical signal and stores it as radiographic image information has been proposed. The digital radiographic imaging system does not use a radiation film unlike the SF system, so there is no complicated process such as development processing, and the screen of an image display device (for example, a cathode ray tube or a liquid crystal display panel) can be quickly displayed. A radiation image can be drawn on top.

  Here, the radiographic image obtained in the digital radiographic imaging system can obtain an image quality equivalent to or higher than that of the SF system, but CR and FPD are expensive, and a lightweight cassette type (moving type) radiographic image conversion panel. It was difficult to realize. Therefore, a radiation image conversion panel has been developed that can obtain a digital radiation image that is inexpensive, lightweight, and high-quality as shown below.

  The radiation image conversion panel described in Patent Document 1 includes a scintillator layer that emits electromagnetic waves according to the intensity of incident radiation, a photoelectric conversion layer that converts electromagnetic waves output from the scintillator layer into electric charges, and the photoelectric conversion layer. In a radiation image conversion panel comprising: an image signal output layer that accumulates the charge obtained in step 1 and outputs a signal based on the accumulated charge; and a substrate that supports the image signal output layer from the scintillator layer. It is made of resin. By forming the substrate with a resin, the substrate can be reduced in weight and resistance to impact can be improved, and the radiation image conversion panel can be reduced in weight and resistance.

  The radiation image conversion panel described in Patent Document 2 is a radiation image conversion panel described in Patent Document 1, in which a phosphor containing oxygen O and rare earth elements gadolinium Gd and europium Eu is used in the scintillator layer. . Since the phosphor has a high radiation absorption rate and a high luminous efficiency, a high-quality radiation image can be obtained.

  In addition, the radiation image conversion panel described in Patent Document 3 is different from the radiation image conversion panel described in Patent Document 1 in that the order of the layers is changed so that the radiation is incident from the substrate side toward the scintillator layer side. It has become. With such a configuration, the electromagnetic waves emitted from the scintillator layer to the photoelectric conversion layer are less scattered, and the sharpness of the radiation image conversion panel can be improved.

Furthermore, the radiation image conversion panel described in Patent Document 4 is a radiation image conversion panel described in Patent Document 1, in which phosphor particles that are constituent elements of a scintillator layer are dispersed in a photoelectric conversion layer. With such a configuration of the photoelectric conversion layer, it is not necessary to separately provide a scintillator layer, and a high-quality radiation image can be obtained.
JP 2003-50280 A JP 2003-57353 A Japanese Patent Laid-Open No. 2003-60181 JP 2003-60178 A

  Here, as a material capable of generating a charge used in the conventional photoelectric conversion layer, a light emitting material used for an organic EL element, such as a conductive polymer material, can be cited. Further, by adding a nanocarbon material having a three-dimensional π electron cloud such as fullerene or carbon nanotube (hereinafter referred to as “CNT”) to the material, the photoelectric conversion rate is improved and the signal value is increased. Can be increased.

  However, when a high concentration of nanocarbon material is added, the nanocarbon materials are aggregated, so that they are not uniformly dispersed in the charge generation layer, resulting in a decrease in photoelectric conversion rate and generation of image unevenness. There was a problem.

  The present invention has been made in view of the above points, and by preventing aggregation of nanocarbon materials and uniformly dispersing them in the charge generation layer, there is no radiation unevenness and a radiological image having a high photoelectric conversion rate. The purpose is to provide a conversion panel.

In order to solve the above problem, the invention according to claim 1 is:
Phosphor particles that absorb the energy of the incident radiation;
A charge generation layer in which energy is transferred directly or indirectly from the phosphor particles to generate charges;
A charge storage element for storing the charge obtained in the charge generation layer;
An image signal output element that outputs an image signal based on the charge accumulated in the charge accumulation element;
A substrate that supports the phosphor particles, the charge generation layer, the charge storage layer, and the image signal output element;
In a radiation image conversion panel comprising:
The charge generation layer contains a conductive polymer compound and a nanocarbon / π-conjugated compound composite.

  According to the first aspect of the present invention, the charge generation layer contains a nanocarbon / π-conjugated compound complex as an additive for improving photoconductivity and carrier mobility. This complex acts as an additive and generates a π-π interaction between the nanocarbon site and the π-conjugated compound site in the molecule of the complex. Therefore, generation | occurrence | production of (pi)-(pi) interaction between composite_body | complex can be suppressed.

The invention according to claim 2 is the radiation image conversion panel according to claim 1,
The nanocarbon / π-conjugated compound composite is formed by a covalent bond between a nanocarbon material and a π-conjugated compound.

  According to the invention described in claim 2, since the nanocarbon / π-conjugated compound complex is formed by a covalent bond, the nanocarbon part and the π-conjugated compound part in the molecule of the complex are between. Strong π-π interaction occurs. Therefore, the occurrence of π-π interaction between the composites can be suppressed.

The invention according to claim 3 is the radiation image conversion panel according to claim 1 or 2,
The π-conjugated compound has a basic skeleton of porphyrin and phthalocyanine.

  According to the invention of claim 3, since the π-conjugated compound has porphyrin or phthalocyanine as a basic skeleton, the charge separation state of the complex is stabilized, and a charge separation complex can be formed by photoexcitation. .

The invention according to claim 4 is the radiation image conversion panel according to any one of claims 1 to 3,
The conductive polymer compound has polyphenylene vinylene, polythiophene, poly (thiophene vinylene), polyacetylene, polypyrrole, polyfluorene, poly (p-phenylene), or polyaniline as a basic skeleton.

  According to invention of Claim 4, since the said electroconductive polymer compound consists of (pi) conjugated polymer compound, it can be set as a charge generation layer with a high photoelectric conversion rate.

The invention according to claim 5 is the radiation image conversion panel according to any one of claims 1 to 4,
The conductive polymer compound has polythiophene as a basic skeleton.

  According to the fifth aspect of the present invention, since the conductive polymer compound has polythiophene as a basic skeleton, a charge generation layer having a higher photoelectric conversion rate can be obtained.

The invention according to claim 6 is the radiation image conversion panel according to any one of claims 1 to 5,
A scintillator layer comprising the phosphor particles;
A photoelectric conversion layer comprising the charge generation layer and converting the energy into electrical energy;
In order from the radiation irradiation direction.

  According to the invention described in claim 6, the radiation image conversion panel includes a scintillator layer including the phosphor particles, a photoelectric conversion layer including the charge generation layer, the charge storage element, and the image signal output element, Are sequentially provided from the radiation irradiation direction, so that a radiation image conversion panel having a high photoelectric conversion rate can be obtained.

The invention according to claim 7 is the radiation image conversion panel according to any one of claims 1 to 5,
The photoelectric conversion layer and the scintillator layer are provided in order from a radiation irradiation direction.

  According to invention of Claim 7, since the said radiographic image conversion panel is equipped with the said photoelectric conversion layer and the said scintillator layer in order from a radiation irradiation direction, it is set as a radiographic image conversion panel with higher sharpness. Can do.

The invention according to claim 8 is the radiation image conversion panel according to any one of claims 1 to 5,
The photoelectric conversion layer includes the phosphor particles.

  According to the invention described in claim 8, since the photoelectric conversion layer includes the phosphor particles, it is not necessary to provide the scintillator layer separately from the photoelectric conversion layer, and a so-called direct radiation image conversion panel is provided. Can do.

The invention according to claim 9 is the radiation image conversion panel according to any one of claims 1 to 7,
The substrate that supports the photoelectric conversion layer is made of a resin.

  According to the ninth aspect of the present invention, it is possible to reduce the weight and improve the durability of the radiation image conversion panel by using a resin as a substrate that supports the components.

The invention according to claim 10 is the radiation image conversion panel according to any one of claims 1 to 9,
The phosphor particles are made of cesium iodide (CsI: Tl) or gadolinium oxysulfide (Gd ₂ O ₂ S: Tb).

According to the invention described in claim 10, since the phosphor particles are made of cesium iodide (CsI: Tl) or gadolinium oxysulfide (Gd ₂ O ₂ S: Tb), the X-ray absorption efficiency and the light emission efficiency are high. It can be set as a radiation image conversion panel.

Invention of Claim 11 in the radiographic image conversion panel as described in any one of Claims 1-10,
The phosphor particles have a crystallite size of 10 nm to 100 nm.

According to the eleventh aspect of the present invention, since the crystallite size of the phosphor particles is 10 nm to 100 nm, a radiation image conversion panel having high luminous efficiency and good granularity can be obtained.
Here, the crystallite size of the phosphor particles is an index indicating the size of the crystallites constituting the phosphor particles, and is measured by the Wilson method.

Invention of Claim 12 in the radiographic image conversion panel as described in any one of Claims 1-11,
The phosphor particles have a particle size of 0.2 μm to 5 μm.

  According to invention of Claim 12, since the particle size of the said fluorescent substance particle is 0.2 micrometer-5 micrometers, it can be set as a radiation image conversion panel with high luminous efficiency and high sharpness.

Invention of Claim 13 in the radiographic image conversion panel as described in any one of Claims 1-12,
The charge storage element that stores the charge obtained in the charge generation layer and the image signal output element that outputs an image signal based on the charge are made of an organic semiconductor.

  According to the thirteenth aspect of the present invention, since the charge storage element and the image signal output element are made of an organic semiconductor, the radiation image conversion panel can be manufactured at a low manufacturing cost.

The invention according to claim 14 is the radiation image conversion panel according to any one of claims 1 to 13,
The charge storage element and the image signal output element are composed of divided silicon stacked structure elements.

  According to the fourteenth aspect of the present invention, since the charge storage element and the image signal output element are composed of divided silicon stacked structure elements, a radiation image conversion panel with low manufacturing cost can be obtained.

The invention according to claim 15 is a portable radiographic image detector,
It has the radiographic image conversion panel as described in any one of Claims 1-14, It is characterized by the above-mentioned.

  According to invention of Claim 15, since it is a portable structure radiographic image detector provided with the radiographic image conversion panel as described in any one of Claims 1-14, imaging | photography of a radiographic image is performed easily. be able to.

The invention according to claim 16 is the radiological image detector according to claim 15,
The apparatus includes a power supply unit that supplies power necessary to drive the radiation image detector.

  According to the sixteenth aspect of the present invention, the radiographic image detector includes power supply means for supplying electric power necessary for driving the radiographic image detector, so that radiographic images can be easily captured.

The invention according to claim 17 is the radiological image detector according to claim 15 or 16,
Storage means for storing the image signal is provided.

  According to the seventeenth aspect of the present invention, the radiological image detector includes the storage unit that stores the image signal output from the radiographic image conversion panel, so that radiographic images can be easily captured.

The invention according to claim 18 is the radiological image detector according to claim 17,
The storage means is detachable.

  According to the invention described in claim 18, since the radiation image detector includes the removable storage means, it is possible to easily take a radiation image.

The invention according to claim 19 is a radiographic imaging system,
A radiographic imaging device having the radiographic image detector according to any one of claims 15 to 18,
A console connected to the radiographic imaging device and operating the radiographic image detector;
It is characterized by providing.

  According to the nineteenth aspect, a radiographic image can be taken by the operator operating the console.

The invention according to claim 20 provides
A radiographic imaging apparatus comprising the radiographic image detector according to claim 17 or 18,
An input operation unit that inputs an instruction from an operator and a control unit that controls image signal transfer processing by the radiographic image detector, and is connected to the radiographic imaging apparatus to operate the radiographic image detector Console,
In the radiographic imaging system characterized by comprising:
The control unit causes the radiographic image detector to transfer the image signal based on an input instruction from the input operation unit.

  According to the invention described in claim 20, since the image signal can be stored in the storage means of the radiation image detector and transferred at an arbitrary timing, for example, during the transfer of the image signal, the patient Thus, it is possible to capture radiographic images efficiently and continuously without causing the subject to wait.

The invention according to claim 21
A radiographic imaging apparatus comprising the radiographic image detector according to claim 17 or 18,
An input operation unit that inputs an instruction from an operator and a control unit that controls image signal transfer processing by the radiographic image detector, and is connected to the radiographic imaging apparatus to operate the radiographic image detector Console,
In the radiographic imaging system characterized by comprising:
The control unit causes the radiographic image detector to delete the image signal based on an input instruction from the input operation unit.

  According to the twenty-first aspect, the image signal stored in the storage unit can be deleted according to the judgment of the operator.

  According to the first aspect of the invention, since the occurrence of π-π interaction between the composites can be suppressed, aggregation due to the π-π interaction between the composites can be prevented. Therefore, the composite can be uniformly dispersed in the charge generation layer, and the conductive polymer material and the additive are uniformly dispersed in the photoelectric conversion layer. An image conversion panel can be obtained, and a high-quality radiation image can be obtained.

  According to the invention described in claim 2, since the occurrence of π-π interaction between the composites can be suppressed, aggregation due to the π-π interaction between the composites can be prevented. Therefore, the composite can be uniformly dispersed in the charge generation layer, and a radiation image conversion panel having no image unevenness and a high photoelectric conversion rate can be obtained, and a high-quality radiation image can be obtained.

  According to the invention described in claim 3, since the charge separation state of the composite is stabilized and charge separation is promoted, the photoelectric conversion rate of the radiation image conversion panel can be improved.

  According to invention of Claim 4, since the said electroconductive polymer compound consists of (pi) conjugated polymer compound, it can be set as a charge generation layer with a high photoelectric conversion rate.

  According to the invention described in claim 5, since the conductive polymer compound is composed of a π-conjugated polymer compound, a charge generation layer having a higher photoelectric conversion rate can be obtained.

  According to invention of Claim 6, it can be set as a radiographic image conversion panel with a high photoelectric conversion rate, and a high quality radiographic image can be obtained.

  According to invention of Claim 7, it can be set as a radiographic image conversion panel with high sharpness and a high photoelectric conversion rate, and a high quality radiographic image can be obtained.

  According to invention of Claim 8, it can be set as a direct-type radiographic image conversion panel with a high photoelectric conversion rate, and a high quality radiographic image can be obtained.

  According to the ninth aspect of the present invention, it is possible to reduce the weight and improve the durability of the radiation image conversion panel, and it is possible to obtain a light-weight radiation image detector having sufficient durability.

  According to invention of Claims 10-12, it can be set as a radiographic image conversion panel with high X-ray absorption efficiency and luminous efficiency, and a high quality radiographic image can be obtained.

  According to invention of Claim 13, 14, it can be set as the radiation image conversion panel with low manufacturing cost.

  According to invention of Claim 15-18, it can be set as a radiographic image detector of a portable structure, and a radiographic image can be obtained easily.

  According to the nineteenth aspect, a radiographic image can be taken by the operator operating the console.

  According to the invention described in claim 20, since the image signal can be stored in the storage means of the radiation image detector and transferred at an arbitrary timing, for example, during the transfer of the image signal, the patient Thus, it is possible to capture radiographic images efficiently and continuously without causing the subject to wait.

  According to the twenty-first aspect, the image signal stored in the storage unit can be deleted according to the judgment of the operator.

  Embodiments of the radiation image conversion panel of the present invention will be described below with reference to the drawings. However, the scope of the invention is not limited to the illustrated examples, and it is needless to say that various modifications can be made without departing from the scope of the invention.

[First Embodiment]
As shown in FIG. 1, the radiation image conversion panel 1a of this embodiment includes a scintillator layer 10 that emits an electromagnetic wave corresponding to the intensity of incident radiation, and a photoelectric that generates an electric charge by absorbing the electromagnetic wave from the scintillator layer 10. The conversion layer 20a, the image signal output layer 30 for accumulating the charge obtained in the photoelectric conversion layer 20a and outputting a signal based on the accumulated charge, the scintillator layer 10, the photoelectric conversion layer 20a, and the image signal output layer 30 And a substrate 40 that supports the substrate.
Each configuration will be described below.

  The scintillator layer 10 has phosphor particles 11 made of a phosphor as a main component, absorbs the energy of incident radiation, and absorbs electromagnetic waves having a wavelength of 300 nm to 800 nm, that is, from ultraviolet light centering on visible light. Emits electromagnetic waves (light) over infrared light.

The phosphor used in the scintillator layer 10 is a tungstate phosphor such as CaWO ₄ , CaWO ₄ : Pb, MgWO, Y ₂ O ₂ S: Tb, Gd ₂ O ₂ S: Tb, La ₂ O ₂ S: Terbium-activated rare earth oxysulfide phosphors such as Tb, (Y, Gd) ₂ O ₂ S: Tb, (Y, Gd) ₂ O ₂ S: Tb, Tm, YPO ₄ : Tb, GdPO ₄ : Tb, LaPO ₄ : Terbium-activated rare earth phosphate phosphor such as Tb, LaOBr: Tb, LaOBr: Tb, Tm, LaOCl: Tb, LaOCl: Tb, Tm, GdOBr: Tb, GdOBr: Tb, Tm, GdOCl: Tb, GdOCl: Terbium activated rare earth oxyhalide phosphors such as Tb and Tm, thulium activated rare earth oxyhalide phosphors such as LaOBr: Tm, LaOCl: Tm Gadolinium-activated rare earth oxyhalide phosphors such as LaOBr: Gd and LuOCl: Gd, cerium-activated rare earths such as GdOBr: Ce, GdOCl: Ce, (Gd, Y) OBr: Ce, and (Gd, Y) OCl: Ce Oxyhalide phosphors, BaSO ₄ : Pb, BaSO ₄ : Eu ²⁺ , (Ba, Sr) SO ₄ : Eu ²⁺ and other barium sulfate phosphors, Ba ₃ (PO ₄ ) ₂ : Eu ²⁺ , (Ba ₂ PO ₄ ) ₂ : Eu ²⁺ , Sr ₃ (PO ₄ ) ₂ : Eu ²⁺ , (Sr ₂ PO ₄ ) ₂ : Eu ^2+, etc., divalent europium activated alkaline earth metal phosphate-based fluorescence Body, BaFCl: Eu ²⁺ , BaFBr: Eu ²⁺ , BaFCl: Eu ²⁺ , Tb, BaFCl: Eu ²⁺ , Tb, BaF ₂ .BaCl ₂ .KCl: Eu ²⁺ , (Ba, Mg) F _2. BaCl ₂・ KCl : Divalent europium-activated alkaline earth metal fluoride halide phosphors such as Eu ²⁺ , iodide phosphors such as CsI: Na, CsI: Tl, NaI, KI: Tl, ZnS: Ag, (Zn , Cd) S: Ag, (Zn, Cd) S: Cu, sulfide-based phosphors such as (Zn, Cd) S: Cu, Ag, HfP ₂ O ₇ , HfP ₂ O ₇ : Cu, Hf ₃ (PO ₄ ) Hafnium phosphate phosphors such as ₄ , YTaO ₄ , YTaO ₄ : Tm, YTaO ₄ : Nb, (Y, Sr) TaO ₄ : Nb, LuTaO ₄ , LuTaO ₄ : Tm, LuTaO ₄ : Nb, (Lu, Sr) tantalate phosphors such as TaO ₄ : Nb, GdTaO ₄ : Tm, Mg ₄ Ta ₂ O ₉ : Nb, Gd ₂ O ₃ .Ta ₂ O ₅ .B ₂ O ₃ : Tb, Gd ₂ O ₂ S: Eu ³⁺ , (La, Gd, Lu) ₂ Si ₂ O ₇ : Eu, ZnSiO ₄ : Mn, Sr ₂ P ₂ O ₇ : Eu, or the like can be used.

In particular, cesium iodide (CsI: Tl) and gadolinium oxysulfide (Gd ₂ O ₂ S: Tb), which have high X-ray absorption and emission efficiency, are preferable. By using these, high-quality images with low noise can be obtained. Can do.

  The crystallite size (by the Wilson method), which is an index indicating the size of the crystallite constituting the phosphor particle 11, is 10 nm to 100 nm. This is because if the crystallite size is smaller than 10 nm, the light emission efficiency is lowered, and if the crystallite size is larger than 100 nm, the phosphor collection rate at the time of production deteriorates. The particle size of the phosphor particles 11 is 0.2 μm to 5 μm. This is because if the particle size is smaller than 0.2 μm, the light emission efficiency is lowered, and if it is larger than 5 μm, the scattering of light emission in the scintillator layer 10 is increased and the sharpness is deteriorated. The particle size of the phosphor particles 11 is preferably 0.5 μm to 2 μm.

  Furthermore, the shape of the phosphor particles 11 is preferably spherical. This is because the dispersibility of the phosphor particles 11 becomes high and the filling rate of the phosphor particles 11 can be increased due to the spherical shape, so that the graininess can be improved.

  The phosphor particles 11 may be dispersed in the following binder 12. For example, polyurethane, vinyl chloride copolymer, vinyl chloride-acrylonitrile copolymer, butadiene-acrylonitrile copolymer, polyamide resin, polyvinyl butyral, cellulose derivative, styrene-butadiene copolymer, various synthetic rubber resins, phenol resin, Examples thereof include an epoxy resin, a urea resin, a melanin resin, a phenoxy resin, a silicon resin, an acrylic resin, and a urea formamide resin. Among these, polyurethane, polyester, vinyl chloride copolymer, polyvinyl butyral, and nitrocellulose are preferably used. By using such a preferred binder 12, it is possible to increase the dispersibility of the phosphor particles 11 and to increase the filling rate of the phosphor.

  The photoelectric conversion layer 20a is provided on the surface side opposite to the radiation irradiation surface side of the scintillator layer 10, and is simplified in FIG. 1, but in order from the scintillator layer 10 side, a diaphragm, a transparent electrode film, a charge generation layer, and A conductive layer is provided. Hereinafter, description will be made sequentially.

  The diaphragm is for separating the scintillator layer 10 from other layers, and for example, Oxi-nitride is used.

The transparent electrode film is an electrode that transmits an electromagnetic wave to be photoelectrically converted, and may be formed on a substrate or another layer. The transparent electrode film is formed using a conductive transparent material such as indium tin oxide (ITO), SnO ₂ , or ZnO.

  The charge generation layer contains a compound that can be converted photoelectrically, that is, a compound that separates charges by transferring energy from radiation by electromagnetic waves (light), and serves as an electron donor that can generate charges. And an additive as an electron acceptor.

The conductive polymer compound used in the charge generation layer has the basic skeleton of polyphenylene vinylene, polythiophene, poly (thiophene vinylene), polyacetylene, polypyrrole, polyfluorene, poly (p-phenylene) or polyaniline shown in Chemical formula 1. Are preferable (in the formula 1, x is preferably an integer of 1 or more). Preferable specific examples of polyphenylene vinylene and derivatives thereof are shown in Chemical Formula 2 (In Chemical Formula 2, n, m, k, and j are integers of 0 or more, and x is an integer of 2 or more). Preferable specific examples of polythiophene and derivatives thereof are shown in Chemical formula 3 (In Chemical formula 3, n and m are integers of 0 or more, k is an integer of 1 or more, and x is an integer of 2 or more). A preferred specific example of poly (thiophene vinylene) and derivatives thereof is shown in Chemical Formula 4 (In Chemical Formula 4, n, m, k, and j are integers of 0 or more, and x is an integer of 2 or more). A preferred specific example of polyacetylene and derivatives thereof is shown in Chemical formula 5 (in Chemical formula 5, n and m are integers of 0 or more, and x is an integer of 2 or more). A preferred specific example of polypyrrole and derivatives thereof is shown in Chemical formula 6 (In Chemical formula 6, n is an integer of 0 or more, k is an integer of 1 or more, and x is an integer of 1 or more). Preferable specific examples of polyfluorene and derivatives thereof are shown in Chemical formula 7 (In chemical formula 7, n and m are integers of 0 or more, and x is an integer of 1 or more). Preferable specific examples of poly (p-phenylene) and derivatives thereof are shown in Chemical formula 8 (In chemical formula 8, n and m are integers of 0 or more, and x and y are integers of 1 or more). Preferable specific examples of polyaniline and derivatives thereof are shown in Chemical formula 9 (In chemical formula 9, n is an integer of 0 or more and x is an integer of 2 or more). The conductive polymer is not limited to the specific examples shown in Chemical Formulas 2 to 9, and polythiophene and derivatives thereof are most preferable from the viewpoint of combination with the phosphor particles.

  In the charge generation layer, an additive having a three-dimensional π electron cloud is added in order to improve photoconductivity and carrier mobility between a plurality of π-conjugated polymer compounds. The additive is a nanocarbon / π-conjugated compound composite in which a nanocarbon material and a π-conjugated compound are combined as described below (see FIGS. 2 and 3).

The nanocarbon material is a material in which carbon atoms are connected in a spherical shape or a cylindrical shape, and the diameter is nano-sized. For example, fullerene C-60, fullerene C-70, fullerene C-76, fullerene C-78, fullerene C-84, fullerene C-240, fullerene C-540, mixed fullerene, fullerene nanotube, multi-wall nanotube, single-wall nanotube, One having a nano horn (conical shape) or the like as a basic skeleton is exemplified. A specific example of a fullerene derivative is shown in Chemical formula 10. A specific example of a CNT derivative is shown in Chemical formula 11 (in Chemical formula 11, n is an integer of 1 or more). The nanocarbon material is not limited to the specific examples shown in Chemical Formula 10 and Chemical Formula 11.

  The π-conjugated compound is an aromatic compound having 7 or more π electrons. Therefore, for example, benzene having only 6 π electrons does not correspond to a π-conjugated compound, and naphthalene corresponds to a π-conjugated compound because it has 10 π electrons. Since the π-conjugated compound generates a π-π interaction with the nanocarbon material, the π-π interaction occurs within the composite, and the π electron cloud of the nanocarbon material spreads more widely, It becomes an extended additive.

The π-conjugated compound is preferably a compound having a basic skeleton of porphyrin or phthalocyanine. The π-conjugated compound having such a structure has a large number of π electrons and emits electrons when irradiated with light. Therefore, upon receiving electromagnetic waves (light) from the phosphor, the nanocarbon / π-conjugated compound complex easily becomes a charge-separated complex. A preferred specific example of the porphyrin derivative is shown in Chemical formula 12 (In Chemical formula 12, n is an integer of 0 or more and m is an integer of 1 or more). A preferred specific example of the phthalocyanine derivative is shown in Chemical formula 13 (In Chemical formula 13, n is an integer of 0 or more). Note that the π-conjugated compound is not limited to the specific examples shown in Chemical formulas 12 and 13.

  The nanocarbon / π-conjugated compound composite is a composite in which at least one π-conjugated compound is combined with one nanocarbon material. Examples of the composite method include ion bonding, electrostatic interaction, metal and ligand. Coordination bond using π, π-π interaction, covalent bond and the like. Here, it is sufficient that the interaction described above occurs between the two, and the composite is not limited to a combination of the nanocarbon material and the π-conjugated compound, but the viewpoint of expanding the π electron cloud. Therefore, it is preferably a covalent bond, particularly preferably a π-π interaction.

  In addition, a spacer may be interposed between the nanocarbon material and the π-conjugated compound. Examples of the spacer include an alkyl group, a phenyl group, an amide group, an ester group, an acetylene group, an oligothiophene, and a structure in which a substituent is bonded to these, and those containing an alkyl group are particularly preferable.

  Since the nanocarbon / π-conjugated compound complex generates a π-π interaction in the complex, the π-π interaction between the complexes is suppressed. Therefore, since the nanocarbon / π-conjugated compound complex is prevented from aggregating with each other, it can be uniformly dispersed as an additive in the charge generation layer, and the photoelectric conversion rate of the photoelectric conversion layer 20a can be improved and obtained. Image unevenness of the radiation image to be produced is prevented.

  The nanocarbon / π-conjugated compound composite has a π-electron cloud that is larger than the π-electron cloud of the nanocarbon itself due to the π-π interaction in the composite, and also performs charge separation in the composite. Is generated. Therefore, the charge generated in the charge generation layer is stabilized and recombination during charge transport is prevented, so that the photoelectric conversion rate of the photoelectric conversion layer 20a can be improved.

  The film thickness of the charge generation layer is preferably 50 nm or more (particularly 100 nm or more) from the viewpoint of increasing the amount of light absorption, and is preferably 1 μm or less (particularly 300 nm or less) from the viewpoint that the electric resistance does not become too large.

  The conductive layer is made of, for example, chromium. In addition, a general metal electrode or the transparent electrode can be selected, but in order to obtain good characteristics, a metal, an alloy, an electrically conductive compound and a mixture thereof having a small work function (4.5 eV or less) are used. What is used as an electrode material is preferable.

  The image signal output layer 30 is provided on the side opposite to the radiation irradiation surface side of the photoelectric conversion layer 20a, and accumulates charges obtained by the photoelectric conversion layer 20a and outputs signals based on the accumulated charges. ing. The image signal output layer 30 uses a capacitor 31 that is a charge storage element that accumulates the charges generated in the photoelectric conversion layer 20a for each pixel, and a transistor 32 that is an image signal output element that outputs the accumulated charges as a signal. Configured.

  As the transistor 32, for example, a TFT (Thin Film Transistor) is used. This TFT may be an inorganic semiconductor type used in a liquid crystal display or the like, or an organic semiconductor, and is preferably a TFT formed on a plastic film. As the TFT formed on the plastic film, an amorphous silicon type is known, but other than that, FSA (Fluidic Self Assembly) technology developed by US Alien Technology, that is, made of single crystal silicon. A TFT may be formed on a flexible plastic film by arranging minute CMOS (Nanoblocks) on an embossed plastic film. Further, it may be a TFT using an organic semiconductor as described in documents such as Science 283, 822 (1999), Appl. Phys. Lett, 771488 (1998), Nature, 403, 521 (2000).

  Thus, as the transistor 32 used in the present invention, a TFT manufactured by the FSA technique and a TFT using an organic semiconductor are preferable, and a TFT using an organic semiconductor is particularly preferable. If a TFT is formed using this organic semiconductor, equipment such as a vacuum deposition apparatus is not required as in the case where a TFT is formed using silicon, and the TFT can be formed by utilizing printing technology or inkjet technology. Cost is low. Furthermore, since the processing temperature can be lowered, it can also be formed on a plastic substrate that is vulnerable to heat.

  The transistor 32 accumulates electric charges generated in the photoelectric conversion layer 20 a and is electrically connected to a collection electrode (not shown) that serves as one electrode of the capacitor 31. The capacitor 31 accumulates the charge generated in the photoelectric conversion layer 20 a and reads the accumulated charge by driving the transistor 32. That is, by driving the transistor 32, a signal for each pixel of the radiation image can be output.

  The substrate 40 is a support for the radiation image conversion panel 1a. A resin, a glass substrate, or the like is used for the substrate 40, but it is preferable to use a resin from the viewpoint of improving resistance or reducing the weight.

Next, output of an image signal by the radiation image conversion panel 1a will be described.
The radiation incident on the radiation image conversion panel 1a of the present embodiment is incident from the scintillator layer 10 side toward the substrate 40 side.

  First, radiation is incident on the scintillator layer 10 and the phosphor particles 11 are contained in the scintillator layer 10. Therefore, the phosphor particles 11 absorb the energy of the radiation in the scintillator layer 10, and an electromagnetic wave corresponding to the intensity is generated. Emits light. Among the emitted electromagnetic waves, the electromagnetic waves incident on the photoelectric conversion layer 20a penetrate the diaphragm and the transparent electrode of the photoelectric conversion layer 20a and reach the charge generation layer. Since the charge generation layer contains the conductive polymer and the nanocarbon / π-conjugated compound complex, the electromagnetic wave is absorbed in the charge generation layer, and a hole-electron pair ( Charge separation state) is formed. The generated electric charges are transported to different electrodes (transparent electrode film and conductive layer) by the internal electric field, and the photocurrent is detected. At this time, since the nanocarbon / π-conjugated compound complex is uniformly added to the charge generation layer, the generated charges can exist stably and recombination during charge transport is suppressed.

  In this way, the generated charges are carried to the conductive layer and accumulated in the capacitor 31 of the image signal output layer 30. The accumulated charge drives the transistor 32 connected to the transistor 31 and outputs an image signal.

As described above, according to the present embodiment, in the charge generation layer, the nanocarbon / π conjugated compound composite is added to improve the photoelectric conversion rate. A uniform charge generation layer can be obtained. Therefore, the radiation image conversion panel 1a without image unevenness and high photoelectric conversion rate can be obtained.
Further, by combining the π-conjugated compound with the nanocarbon material, the π electron cloud is further expanded and the charge generated in the charge generation layer is stabilized, so that the radiation image conversion panel 1a having a higher photoelectric conversion rate is obtained. be able to.

  Next, the radiation image detector 50 provided with the radiation image conversion panel 1a of this embodiment will be briefly described.

  An example of the structure of the radiation image detector 50 is shown in FIG. The radiation image detector 50 is output from the radiation image conversion panel 1a using a radiation image conversion panel 1a, a control unit 51 that controls the operation of the radiation image detector 50, a rewritable read-only memory (for example, a flash memory), and the like. A memory unit 52 that is a storage unit for storing the image signal, a power supply unit 53 that is a power supply unit that supplies power necessary to obtain the image signal by driving the radiation image conversion panel 1a, and the like. The body 54 is provided in the body 54. The housing 54 has a communication connector 55 for performing communication from the radiation image detector 50 to the outside as needed, and an operation for switching the operation of the radiation image detector 50. A unit 56, a display unit 57 indicating completion of radiographic imaging preparation, a predetermined amount of image signal written in the memory unit 52, and the like are provided.

  Here, as shown in FIG. 4, the radiographic image detector 50 is provided with a power supply unit 53 and a memory unit 52 that stores an image signal of the radiographic image, and the radiographic image detector 50 is detachable via a connector 55. If it carries out, it can be set as the portable structure which can carry the radiographic image detector 50. FIG. Further, if the memory unit 52 is configured to be detachable using a non-volatile memory, the memory unit 52 can be attached to the image processing unit 63 described later without connecting the radiation image detector 50 and the image processing unit 51. Since the image signal can be supplied to the image processing unit 63, the radiographic image can be easily captured and processed, and the operability can be improved. When the radiological image detector 50 is used as a stationary type, the supply of electric power and the reading of an image signal via the connector 55 allow the radiographic image to be read without providing the memory unit 52 and the power source unit 53. Of course, an image signal can be obtained.

  Next, an outline of a radiographic imaging system using such a radiographic image detector 50 will be described with reference to FIG.

  FIG. 5 is a diagram showing a schematic configuration of an embodiment of a radiographic image capturing system 60 to which the radiographic image detector according to the present invention is applied.

  As shown in FIG. 5, the radiographic image capturing system 60 according to the present embodiment includes a server 61 that manages information related to radiographic image capturing, an image capturing operation device 62 that performs operations related to radiographic image capturing, and a wireless LAN (Local Area Network), for example. ) And the like, and a console 64 for operating the radiation image detector 50 are connected through a network 65. A radiographic imaging device 67 is connected to the imaging operation device 62 via a cable or the like to irradiate a subject (a patient 66 in the present embodiment) with radiation to take a radiographic image. Here, the network 65 may be a dedicated communication line for the system. However, the network 65 is an existing line such as Ethernet (registered trademark) because the degree of freedom of the system configuration is low. Is preferred.

  The server 61 is configured by a computer, and the server 61 includes a control unit that controls each unit that configures the server 61, an input operation unit that inputs various types of information and instructions from an operator, and an external unit that stores various types of information. A storage device or the like is provided (none of which is shown).

  The control unit of the server 61 is configured to store the patient information, imaging information, and the like input from the input operation unit in an external storage device in a state where they are associated with each other. Patient information is information about the patient 66 such as the patient's 66 name, age, sex, date of birth, patient ID number for identifying the patient 66, and the like. The imaging information is information necessary to perform imaging such as an imaging part (a part to be imaged on the subject's body), an imaging direction and a method.

  Instead of the server 61, an information reading device such as a card reader for reading patient information or imaging information written in advance on the ID card may be provided.

  The radiographic imaging device 67 has a radiation source 67a, and generates radiation by applying a tube voltage to the radiation source 67a. A diaphragm device 67b for adjusting the radiation field is provided at the radiation irradiation port of the radiation source 67a so as to be freely opened and closed. A bed 67c on which the patient 66 is placed via the radiation image detector 50 is provided below the radiation source 67a and in the radiation irradiation range.

  The imaging operation device 62 is configured by a computer including a display unit for displaying information and an input operation unit (none of which is shown) for inputting instructions from a radiographer who is an operator, and corresponds to imaging conditions. The radiation source 67a, the diaphragm device 67b, and the like of the radiation image photographing device 67 are controlled so that the photographing is performed based on the tube voltage value, the radiation dose, and the irradiation field diaphragm value.

  As shown in FIG. 6, the console 64 includes a control unit 64a, a RAM 64b, a ROM 64c, a display unit 64d, an input operation unit 64e, a communication unit 64f, and the like, and each unit is connected by a bus 64g.

  The display unit 64d includes, for example, a CRT (Cathode Ray Tube), an LCD (Liquid Crystal Display), and the like, and displays various screens according to instructions of display signals output from the control unit 64a. ing.

  The input operation unit 64e is composed of, for example, a keyboard, a mouse, and the like, and outputs a key press signal pressed by the keyboard or an operation signal from the mouse as an input signal to the control unit 64a. .

  The communication unit 64f communicates various types of information with the radiation image detector 50 via the base station 63 by a wireless communication method such as a wireless LAN.

  The control unit 64a is composed of, for example, a CPU or the like, reads a predetermined program stored in the ROM 64c, develops it in the work area of the RAM 64b, and executes various processes according to the program.

  The control unit 64 a transmits various received information to the radiation image detector 50 wirelessly via the base station 63.

  The control unit 64a is configured to cause the radiation image detector 50 to transfer an image signal based on an input instruction from the input operation unit 64e.

  Further, the control unit 64a is configured to delete the image signal stored in the radiation image detector 50 based on an input instruction from the input operation unit 64e.

  Next, the operation of the radiation image capturing system 60 will be described.

  When radiographic images are captured, patient information and imaging information are transmitted from the server 61 to the imaging operation device 62. The imaging operation device 62 appropriately displays the received information on the display unit, and the radiologist captures a radiographic image while confirming the information.

  The imaging operation device 62 controls the radiation source 67a and the aperture device 67b of the radiographic imaging device 67 based on the tube voltage value, the dose, and the irradiation field aperture value included in the received imaging information. The patient 66 is irradiated with radiation under these conditions.

  At this time, the radiation image detector 50 is inserted on the bed 67c and under the patient 66, detects the amount of radiation transmitted through the patient 66, converts the detected radiation into an electrical signal, and outputs an image signal. And the acquired image signal is stored in the memory unit 52.

  Then, the radiographic image signal stored in the memory unit 52 of the radiographic image detector 50 is transmitted to the console 64 wirelessly via the base station 63.

  The console 64 transmits the image signal received through the communication unit 64f to the server 61 in association with the patient information and the like. Thereafter, each image signal is subjected to image processing by the server 61 and then output as appropriate, and is used as a radiation image for diagnosis by a doctor.

  As described above, according to the radiographic image capturing system 60 according to the present embodiment, the radiographic image detector 50 including the radiographic image conversion panel 1a having no image unevenness and a high photoelectric conversion rate is used. Can be obtained.

[Second Embodiment]
Next, a second embodiment of the present invention will be described.
As shown in FIG. 7, the radiation image conversion panel 1b of the second embodiment includes a scintillator layer 10, a photoelectric conversion layer 20a, an image signal output layer 30, and a substrate 40, as in the first embodiment. In the first embodiment, the radiation image conversion panel 1a includes the scintillator layer 10, the photoelectric conversion layer 20a, the image signal output layer 30, and the substrate 40 in this order from the radiation irradiation direction. The substrate 40, the image signal output layer 30, the photoelectric conversion layer 20a, and the scintillator layer 10 are provided in this order from the radiation irradiation direction, and the radiation is incident on the scintillator layer 10 side from the substrate 40.

  The output of the image signal by the radiation image conversion panel 1b having such a configuration will be described. First, radiation enters the substrate 40, passes through the image signal output layer 30 and the photoelectric conversion layer 20a in this order, and reaches the scintillator layer 10. To do. Then, the electromagnetic wave emitted according to the intensity of radiation in the scintillator layer 10 is less scattered. Thereafter, the electromagnetic wave is photoelectrically converted in the photoelectric conversion layer 20 a and an image signal is output from the image signal output layer 30. Therefore, the same effect as that of the first embodiment can be obtained while improving the sharpness of the radiation image conversion panel 1b.

[Third Embodiment]
Next, a third embodiment of the present invention will be described.
As shown in FIG. 8, the photoelectric conversion layer 20 c in the radiation image conversion panel 1 c according to the third embodiment is a so-called direct type photoelectric module in which the photoelectric conversion layer 20 a according to the first embodiment includes the phosphor particles 11. It is the conversion layer 20c. The radiation image conversion panel 1c according to the present embodiment includes a photoelectric conversion layer 20c, an image signal output layer 30, and a substrate 40 in order from the radiation irradiation direction.

  The radiation image conversion panel 1c having such a configuration is a so-called direct radiation image conversion panel. Here, the direct-type radiation image conversion panel refers to a photoelectric conversion layer containing phosphor particles and a conductive polymer compound, so that radiation is directly absorbed and converted into electric energy in the photoelectric conversion layer. It has become. On the other hand, the radiation image conversion panels according to the first embodiment and the second embodiment are so-called indirect radiation image conversion panels in which radiation energy is indirectly absorbed in the photoelectric conversion layer.

  Even in direct-type radiation image conversion panels, phosphor particles absorb radiation, energy is transferred from the phosphor particles to the conductive polymer compound, and the conductive polymer compound is in a charge-separated state to generate electrical energy. The principle of doing is the same. However, the two are different in the manner in which energy is transferred from the phosphor particles that have absorbed the radiation to the conductive polymer compound. In the direct type, when the molecular orbital of the phosphor particles and the molecular orbital of the conductive polymer compound are close to each other, in addition to the energy transfer by electromagnetic waves emitted from the phosphor particles as in the indirect type, the phosphor particles Energy may be directly transferred to the conductive polymer compound by electrons or the like earlier than the light emitting electromagnetic wave.

  The output of the image signal by the radiation image conversion panel 1c in the present embodiment will be described. Since the scintillator layer is not separately provided in the radiation image conversion panel 1c of the present embodiment, the radiation is incident on the photoelectric conversion layer 20c. Then, the phosphor particles 11 included in the photoelectric conversion layer 20c absorb radiation. Part of the absorbed energy is immediately transmitted to the conductive polymer compound using electrons as a medium, and the remaining energy is transmitted to the conductive polymer compound as electromagnetic waves. Therefore, charges are generated from the charge generation layer included in the photoelectric conversion layer 20c, and an image signal is output from the image signal output layer 30 based on this charge. Therefore, the direct radiation image conversion panel 1c can achieve the same effects as those of the indirect radiation image conversion panel in the first embodiment.

  Hereinafter, the present invention will be described more specifically with reference to examples and comparative examples. However, the present invention is not limited to these examples.

[Production of radiation image conversion panel]
First, each radiation image conversion panel was produced as shown below.

(1) Creation of Example 1 For the scintillator layer, cesium iodide (CsI: Tl) was used as a phosphor to form a scintillator layer having a thickness of 400 μm.
The photoelectric conversion layer was provided with a transparent electrode film, a charge generation layer, and a conductive layer. For the transparent electrode film, ITO was used for the transparent electrode, and the thickness was 100 nm. On the transparent electrode film, a mixed film in which a CNT derivative (Chemical Formula 15) and a naphthimide derivative (Chemical Formula 16) are added as additives to a conductive polymer polycarbazole (Chemical Formula 14) is formed, and a charge generation layer having a thickness of 150 nm is formed. Formed. A conductive layer having a thickness of 250 nm made of aluminum was formed on the charge generation layer.
Here, a method for manufacturing the charge generation layer of the photoelectric conversion layer will be described. First, Chemical 15 and Chemical 16 were mixed and dissolved in dichlorobenzene at an equivalent ratio of 1:20, and then 20 equivalents of Chemical 14 were added to Chemical 15 to obtain a mixed solution. This mixed solution was spin-coated on the transparent electrode film to form a mixed film, which was used as a charge generation layer.
Amorphous silicon was used for the image signal output layer, and a thin film transistor (TFT) was formed.
The substrate was made of polyethylene terephthalate (PET).
These scintillator layer to image signal output layer were provided on the substrate to obtain the radiation image conversion panel of Example 1.

(2) Production of Example 2 As shown in Table 1, Compound 1: Compound 6 = 20: 1 with an additive ratio of CNT / pyromellitimide complex (Chemical Formula 17) and an equivalent ratio of the composition of the mixed film. A charge generation layer was produced in the same manner as in Example 1 except that, and Example 2 according to the present invention was obtained.

(3) Production of Example 3 As shown in Table 1, the additive was used as a CNT / porphyrin complex (Chemical Formula 18), and the composition of the mixed film was set to an equivalent ratio of Compound 1: Compound 7 = 20: 1. Except for the above, a charge generation layer was produced in the same manner as in Example 1, and Example 3 according to the present invention was made.

(4) Production of Example 4 As shown in Table 1, the conductive polymer compound was polyphenylene vinylene (Chemical Formula 19), and the composition of the mixed film was equivalent to the compound 2: compound 7 = 20: 1. Except for the above, a charge generation layer was produced in the same manner as in Example 1, and this was designated as Example 4 according to the present invention.

(5) Production of Example 5 As shown in Table 1, the conductive polymer compound was polythiophene (Chemical Formula 20), and the composition of the mixed film was equivalent to the ratio of Compound 3: Compound 7 = 20: 1 in an equivalent ratio. Prepared a charge generation layer in the same manner as in Example 1, and designated Example 5 according to the present invention.

(6) Creation of Example 6 The photoelectric conversion layer was provided with a transparent electrode film, a charge generation layer, and a conductive layer. ITO was used as a transparent electrode for the transparent electrode film. An insulating film made of polyurethane is coated on the transparent electrode film, and a CNT derivative (Chemical Formula 15) or a naphthimide derivative (Chemical Formula 16) as an additive to the conductive polymer compound polycarbazole (Chemical Formula 14) is coated on the insulating film. And a phosphor (Gd ₂ O ₂ S: Tb) added film was formed to form a 150 μm thick charge generation layer. A conductive layer made of aluminum and having a thickness of 200 nm was formed on the charge generation layer.
Here, a method for manufacturing the charge generation layer of the photoelectric conversion layer will be described. First, compound 15 and compound 16 are mixed and dissolved in dichlorobenzene in an equivalent ratio of 1:20, and then 20 equivalents of compound 14 and 1 equivalent of phosphor (Gd ₂ O ₂ S: Tb) was added to obtain a mixed solution. This mixed solution was spin-coated on the insulating film to form a mixed film, which was used as a charge generation layer.
Amorphous silicon was used for the image signal output layer, and a thin film transistor (TFT) was formed.
The substrate was made of polyethylene terephthalate (PET).
The photoelectric conversion layer and the image signal output layer were provided on the substrate to obtain a radiation image conversion panel of Example 6 without a scintillator layer.

(7) Production of Examples 7 to 10 As shown in Table 2, the charge generation layer was produced by the production method of Example 6 in the same manner as in Examples 2 to 5 with the composition of the mixed film, It was set as Example 7-10 Example which concerns on this invention.

(11) Preparation of Comparative Example 1 Next, as shown in Table 1, except that the additive was only a CNT derivative, and the composition of the mixed film was changed to an equivalent ratio of Chemical 14: Chemical 15 = 20: 1. A charge generation layer was produced in the same manner as in Example 1 and used as Comparative Example 1 according to the present invention.

(12) Preparation of Comparative Example 2 As shown in Table 2, Example 6 was used except that the additive was only a CNT derivative, and the composition of the mixed film was changed to an equivalent ratio of Chemical formula 14: Chemical formula 15 = 20: 1. Similarly, a charge generation layer was produced and used as Comparative Example 2 according to the present invention.

[Evaluation]
Evaluation of each of the radiation image conversion panels of Comparative Examples 1 and 2 and Examples 1 to 10 obtained as described above was performed by the following two types of evaluation methods.

(1) Evaluation method A
The radiation image conversion panels of Comparative Example 1 and Examples 1 to 5 are provided with a scintillator layer, a photoelectric conversion layer, an image signal output layer, and a substrate in order from the radiation irradiation direction, and X from the scintillator layer toward the substrate side. Line (50 kV) irradiation was performed, the signal value from the thin film transistor in the image signal output layer was measured, and the signal value in Comparative Example 1 was expressed as a relative signal value of 1. Table 1 shows the results.
Similarly, the radiation image conversion panels of Comparative Example 2 and Examples 6 to 10 are provided with a photoelectric conversion layer, an image signal output layer, and a substrate in order from the radiation irradiation direction, from the photoelectric conversion layer toward the substrate side. X-ray (50 kV) irradiation was performed, the signal value from the thin film transistor in the image signal output layer was measured, and the signal value of Comparative Example 2 was expressed as a relative signal value of 1. Table 2 shows the results.

(2) Evaluation method B
About the radiographic image conversion panel of the comparative example 1 and Examples 1-5, it shall be equipped with a board | substrate, an image signal output layer, a photoelectric converting layer, and a scintillator layer in an order from a radiation irradiation direction, and it X-rays toward the scintillator layer side from a board | substrate. (50 kV) irradiation was performed, the signal value from the thin film transistor in the image signal output layer was measured, and the signal value of Comparative Example 1 was represented by a relative signal value of 1. Table 1 shows the results.

  As shown in Table 1, when Comparative Example 1 and Example 1 are compared, in the charge generation layer, the chemical polymer 14 is a conductive polymer compound, the nanocarbon material is 15 and the π-conjugated compound is 16 It can be seen that the relative signal value is increased to 3.2 by adding the mixture. That is, it can be seen that the photoelectric conversion rate is improved by adding the nanocarbon material and the π-conjugated compound to the charge generation layer.

  When Example 1 and Example 2 are compared, in the charge generation layer, by adding the chemical compound 17 of the nanocarbon / π-conjugated compound complex to the chemical polymer 14 that is a conductive polymer compound, the relative signal value is increased. It can be seen that it increases to 10.2. That is, it can be seen that the photoelectric efficiency is improved by adding a composite of both of them before forming a mixture of the nanocarbon material and the π-conjugated compound.

  When Example 2 and Example 3 are compared, the relative signal value increases to 15.1 by using the compound π of a π-conjugated compound having a porphyrin as a basic skeleton for the complex added in the charge generation layer. I understand. That is, it can be seen that it is preferable to use a compound having a porphyrin as a basic skeleton as the π-conjugated compound used in the nanocarbon / π-conjugated compound composite.

  Comparing Example 3 and Example 4, it can be seen that the relative signal value increases to 35.5 by using polyphenylene vinylene represented by Chemical Formula 19 as the conductive polymer compound in the charge generation layer. Similarly, when Example 4 is compared with Example 5, it can be seen that the relative signal value is increased to 72.1 by using polythiophene represented by Chemical Formula 20 as the conductive polymer compound. That is, as the conductive polymer compound, polyphenylene vinylene (Chemical Formula 19) is preferable, and polythiophene (Chemical Formula 20) is particularly preferable.

  Further, as can be seen from the result of the relative signal value by the evaluation method B shown in Table 1, the same effect can be obtained also in the radiation image conversion panel having a configuration in which X-rays are incident from the substrate side toward the scintillator layer side. By adopting such a configuration, the electromagnetic wave (light) from the phosphor of the scintillator layer supplied to the photoelectric conversion layer is less scattered, so that the photoelectric conversion rate is improved in each radiation image conversion panel. It appears more prominently.

  Furthermore, although the radiation image conversion panel shown in Table 2 includes phosphor layers in the photoelectric conversion layer and is composed of three layers, it shows the same tendency as in Table 1, and each additive is added to the charge generation layer. It can be seen that the same effects as in Table 1 can be obtained without adding a scintillator layer.

It is a schematic block diagram which shows the radiographic image conversion panel 1a in 1st Embodiment. It is a figure which shows an example of a fullerene / (pi) share type compound complex. It is a figure which shows an example of a CNT / (pi) share type compound composite. 3 is a diagram illustrating an example of the structure of a radiation image detector 50. FIG. It is a figure which shows an example of the radiographic imaging system 60 using the radiographic image detector. 3 is a block diagram showing a main configuration of a console 64. FIG. It is a schematic block diagram which shows the radiographic image conversion panel 1b in 2nd Embodiment. It is a schematic block diagram which shows the radiographic image conversion panel 1c in 3rd Embodiment.

Explanation of symbols

1a, 1b, 1c Radiation image conversion panel 10 Scintillator layer 11 Phosphor particle 20a, 20c Photoelectric conversion layer 30 Image signal output layer 31 Capacitor 32 Transistor 40 Substrate 50 Radiation image detector

Claims (21)

Phosphor particles that absorb the energy of the incident radiation;
A charge generation layer in which energy is transferred directly or indirectly from the phosphor particles to generate charges;
In a radiation image conversion panel comprising:
The charge generation layer contains a conductive polymer compound and a nanocarbon / π-conjugated compound composite, wherein the radiation image conversion panel.   The radiation image conversion panel according to claim 1, wherein the nanocarbon / π-conjugated compound complex is formed by a covalent bond between a nanocarbon material and a π-conjugated compound.   The radiation image conversion panel according to claim 1, wherein the π-conjugated compound has porphyrin and phthalocyanine as a basic skeleton.   The conductive polymer compound has a basic skeleton of polyphenylene vinylene, polythiophene, poly (thiophene vinylene), polyacetylene, polypyrrole, polyfluorene, poly (p-phenylene), or polyaniline. The radiation image conversion panel as described in any one of Claims.   The radiation image conversion panel according to claim 1, wherein the conductive polymer compound has polythiophene as a basic skeleton. A scintillator layer comprising the phosphor particles;
A photoelectric conversion layer comprising the charge generation layer and converting the energy into electrical energy;
The radiation image conversion panel according to claim 1, wherein the radiation image conversion panel is provided in order from a radiation irradiation direction.   The radiation image conversion panel according to any one of claims 1 to 5, wherein the photoelectric conversion layer and the scintillator layer are provided in order from a radiation irradiation direction.   The radiation image conversion panel according to claim 1, wherein the photoelectric conversion layer includes the phosphor particles.   The radiation image conversion panel according to claim 1, wherein the substrate that supports the photoelectric conversion layer is made of a resin. The radiographic image conversion according to any one of claims 1 to 9, wherein the phosphor particles are made of cesium iodide (CsI: Tl) or gadolinium oxysulfide (Gd ₂ O ₂ S: Tb). panel.   The radiation image conversion panel according to any one of claims 1 to 10, wherein a crystallite size of the phosphor particles is 10 nm to 100 nm.   The radiation image conversion panel according to claim 1, wherein a particle size of the phosphor particles is 0.2 μm to 5 μm.   13. The charge storage element for storing the charge obtained in the charge generation layer and the image signal output element for outputting an image signal based on the charge are made of an organic semiconductor. The radiation image conversion panel as described in any one of Claims.   The radiation image conversion panel according to any one of claims 1 to 13, wherein the charge storage element and the image signal output element are composed of divided silicon stacked structure elements.   A radiation image detector having a portable structure comprising the radiation image conversion panel according to claim 1.   The radiographic image detector according to claim 15, further comprising power supply means for supplying electric power necessary for driving the radiographic image detector.   The radiographic image detector according to claim 15, further comprising a storage unit that stores an image signal output from the radiographic image conversion panel.   The radiographic image detector according to claim 17, wherein the storage unit is detachable. A radiographic imaging device having the radiographic image detector according to any one of claims 15 to 18,
A console connected to the radiographic imaging device and operating the radiographic image detector;
A radiographic imaging system comprising: A radiographic imaging apparatus comprising the radiographic image detector according to claim 17 or 18,
An input operation unit that inputs an instruction from an operator and a control unit that controls image signal transfer processing by the radiographic image detector, and is connected to the radiographic imaging apparatus to operate the radiographic image detector Console,
In the radiographic imaging system characterized by comprising:
The control unit causes the radiographic image detector to transfer the image signal based on an input instruction from the input operation unit. A radiographic imaging apparatus comprising the radiographic image detector according to claim 17 or 18,
An input operation unit that inputs an instruction from an operator and a control unit that controls image signal transfer processing by the radiographic image detector, and is connected to the radiographic imaging apparatus to operate the radiographic image detector Console,
In the radiographic imaging system characterized by comprising:
The control unit causes the radiographic image detector to delete the image signal based on an input instruction from the input operation unit.

Images