JP2003119070A - Phosphor element, radiation detector using the same and medical diagnostic imaging device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a phosphor element having high light emission efficiency and low energy dependence and to provide a radiation detector and a medical diagnostic imaging device which use the element. SOLUTION: The phosphor element uses a phosphor comprising a mother crystal having a garnet structure with Ce as the light emitting element and containing at least Gd, Al, Ga and O. The phosphor element has >=99.8% relative density, >=4 μm average grain size and 550 nm main wavelength for mission of light. The phosphor element shows <=0.6 mm<-1> absorption coefficient μ for light at 550 nm wavelength and has high transmittance. Therefore, the element does not cause decrease in the sensitivity for radiation in a low energy region and is suitable for a radiation detector, particularly for a radiation detector of a medical diagnostic imaging device.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a radiation detector for detecting X-rays, γ-rays, etc., and particularly to a radiation detector suitable for a medical image diagnostic apparatus such as an X-ray CT apparatus.

[0002]

2. Description of the Related Art Conventionally, a radiation detector used in a medical image diagnostic apparatus such as an X-ray CT apparatus has a Xe gas ionization chamber type, a single crystal scintillator such as CdWO ₄ , CsI: Tl, or a rare earth oxide. -Based phosphor (Japanese Patent Publication No. 63-59436)
3-50991, WO99 / 33934) and rare earth sulfide-based phosphors (Japanese Patent Publication No. 60-4856), such as ceramic scintillator and other phosphor elements, combined with Si photodiodes and photomultiplier tubes and other photoelectric conversion elements Solid state detectors have been used. In radiation detectors of recent X-ray CT apparatuses, solid-state detectors have become the mainstream.

The characteristics required for a radiation detector used in such a medical image diagnostic apparatus are that it has a high luminous efficiency with respect to radiation and that it is highly sensitive to radiation in a wide range of energy, that is, energy dependence. It can be mentioned that the sex is small. When the energy dependence is large, the density resolution of the radiation detector becomes low for the radiation of energy having low sensitivity, which causes the deterioration of the image quality. The characteristics of the radiation detector greatly depend on the characteristics of the phosphor element among the elements constituting the radiation detector, and it is desirable to use a phosphor element having high luminous efficiency and low energy dependence.

Up to now, several proposals have been made regarding a phosphor element which has a high luminous efficiency and a small energy dependence and is suitable for a radiation detector, and the applicant of the present invention has also selected Ce as a luminous element and Gd, Al and Ga. We have proposed rare-earth oxide phosphors containing them and have achieved extremely high radiation sensitivity compared to CdWO ₄ . However, in recent X-ray CT apparatuses,
There is a tendency that the scanning time of X-rays that rotate around the subject is shortened to shorten the measurement time, and the detector element tends to be one-dimensional and miniaturized. Energy-dependent products are required.

[0005]

SUMMARY OF THE INVENTION Therefore, an object of the present invention is to provide a phosphor element having high luminous efficiency and low energy dependence. Another object of the present invention is to provide a radiation detector having high sensitivity to radiation having a wide range of energy. A further object of the present invention is to provide a medical image diagnostic apparatus capable of obtaining a high quality image even when the scanning time is shortened.

[0006]

In order to achieve the above object, the inventors of the present invention have found that Gd ₃ (Al, Ga) ₅ O ₁₂ : Ce having high luminous efficiency.
As a result of diligent research on the conditions for obtaining low energy dependence of the phosphor element of the system, as shown in FIG. 1, the energy dependence of the phosphor is related to its light transmittance, and the light transmittance is high. In the case where the light transmittance is low, the decrease in sensitivity is less than that in the case where the light transmittance is 100% over a wide energy range, but when the light transmittance is low, the decrease in sensitivity also increases as the radiation energy decreases. I found that. In particular, by setting the absorption coefficient μ for light of 550 nm, which is the main emission wavelength of this phosphor, to 0.6 mm ^-1 or less, at the minimum thickness that can maintain the radiation absorption rate required as a radiation detector, low energy The inventors of the present invention have found that the reduction in sensitivity to radiation can be reduced and have led to the present invention.

That is, the phosphor element of the present invention is a phosphor element using Ce as a light emitting element and a phosphor composed of a matrix crystal of a garnet structure containing at least Gd, Al, Ga and O, and having a wavelength of The absorption coefficient μ for light of 550 nm is 0.6 mm ^-1 or less. Where μ is the light absorption I (d) = I ₀ e ^-μd
The absorption coefficient is expressed as (d is the thickness of the absorber).

The phosphor element of the present invention is the above phosphor element having a relative density of 99.8% or more and an average crystal grain size of 4 μm or more. Here, the relative density is a relative value when the density (theoretical density) of a single crystal is 100%. With a dense structure having a large crystal grain size and a high relative density, the optical path length of light can be shortened, and thus the light transmittance can be increased.

The phosphor element of the present invention will be described in detail below. The phosphor element of the present invention is at least Gd, Ce, Ga, A
It consists of l and O, and consists of a garnet-structured host crystal. Due to the garnet structure, the atomic ratio of Gd / (Al + Ga + Gd) is 0.3.
It is preferably 3 or more and 0.42 or less. The phosphor composed of such a garnet structure mother crystal is similar to the phosphor described in the patent application specification by the present applicant (JP 2001-4753A), and is described in this specification, A description of the composition and structure of the phosphor is part of this specification.

Further, the phosphor of the present invention has an absorption coefficient μ of 0.6 mm ^-1 or less for light having a wavelength of 550 nm, and since it has such an absorption coefficient, it has a high crystal density of 4 μm or more. It is preferably crystalline. The phosphor made of such a polycrystal has a process of 1) mixing and synthesizing raw material powders such as oxides and carbonates of constituent elements, and 2) powder firing (including sintering) of the powder after synthesis. It can be manufactured through the process described above, and at this time, it is sintered by using a synthetic powder having a large primary particle diameter, or is sintered under the sintering condition that crystal grains grow during sintering. As a result, a sintered body having a large crystal grain size and a high relative density can be obtained. In the present specification, the primary particle size of the powder means the particle size of the smallest unit particle constituting the powder particle, and the crystal grain size of the sintered body is the sintered body which is a polycrystalline body. It means the grain size of the crystal grain of the smallest unit constituting the.

Hereinafter, a specific method for manufacturing the phosphor element of the present invention will be described. As the synthesis process, 1) a method mainly based on a normal oxide mixing method, 2) a method through a liquid phase such as a coprecipitation method or a sol-gel method, 3) a powder synthesized mainly based on the oxide mixing method is mechanically again. There is a method of miniaturizing.

In the case of the oxide mixing method, Gd ₂ O ₃ , Al ₂ O ₃ and Ga ₂ O
₃ , raw material powders such as Ce ₂ O ₃ are wet-mixed, and the mixed powders are fired in the atmosphere at 1400 ° C. to 1700 ° C. for several hours. At this time, the firing temperature is appropriately adjusted depending on the particle size of the raw material powder and the subsequent sintering process. For example, the particle size of the raw material powder is relatively small, for example, submicron size (1 μm or less)
In this case, it is preferable to bake at 1500 ° C or lower. When the raw material powder having such a particle size range is fired at 1500 ° C. or less, the primary particle diameter of the powder produced after firing is about the same as that of the raw material powder, but the submicron-sized powder grows crystals after sintering. As a result, a sintered body having a large crystal grain size can be obtained.

When the particle size of the raw material powder is relatively large, for example, about 0.1 to several μm, 1500 ° C. or higher, preferably 1
Bake at a high temperature of 550 ℃ to 1650 ℃. As a result, a phosphor powder having a relatively large primary particle size can be obtained. Specifically, a phosphor powder having an average particle size of several μm or more is obtained. By sintering the phosphor powder having such a primary particle diameter, a sintered body having a similar crystal grain size can be obtained. In order to obtain a crystal grain size as large as possible as a sintered body, it is preferable to further classify the phosphor powder after firing and use a powder having a predetermined size or more for sintering.

In the case of the coprecipitation method, the raw material nitrates, hydrochlorides, sulfates, oxalates, etc. are dissolved in water to form a composite aqueous solution.
Urea, ammonium hydrogen carbonate, etc. are added to this complex aqueous solution to precipitate the Gd-Ce-Al-Ga-O precursor. The precipitate is washed, dried, and then calcined at about 1200 ° C to obtain a synthetic powder. The coprecipitation method yields synthetic powders of submicron or smaller. The method of mechanically refining is a method of refining a synthetic powder obtained by an oxide mixing method by a pulverizing means such as a ball mill, and is used when a synthetic powder of submicron or less is obtained.

The synthetic powder obtained by the above method is pressed into a desired shape, for example, a plate shape, and then sintered. The sintering can be performed by a hot pressing method, a HIP method, an atmospheric pressure sintering method, a combined method of the atmospheric pressure sintering method and the HIP method, or the like. Among these, the hot pressing method and the atmospheric pressure sintering method are particularly preferable. In the hot press method, under an atmosphere of vacuum, 14
Sintering is performed for several hours at a sintering temperature of 00 to 1700 ° C and a pressure of about 500 kgf / cm ² . In the atmospheric pressure sintering method, after performing isostatic pressing at a pressure of about 3000 kgf / cm ² , at the sintering temperature just below the melting point of the phosphor (specifically 1600 to 1700 ° C), in the atmosphere or pure oxygen. Sinter for several hours to several tens of hours. The HIP method can be adopted when using a synthetic powder of submicron or less,
Place synthetic powder in a metal capsule, vacuum seal, and
Sintering is performed at a temperature of about 00 ° C and a pressure of about 2000 atm.

At the time of sintering, it is preferable to use a sintering aid. As the sintering aid, lithium compounds such as LiF and LiCl, potassium compounds such as K ₂ SO ₄ , KNO ₃ , K ₂ CO ₃ and K ₃ PO ₄ , and BaF ₂ can be used. LiF and LiCl are particularly preferable. Such a sintering aid is 0.
It can be used in an amount of about 001 to 10% by weight, and the addition thereof can promote the growth of crystals during sintering and obtain a sintered body having a large crystal grain size.

As described above, in the case of a synthetic powder having a small particle size obtained by firing a raw material powder having a relatively small particle size, the above-mentioned firing is performed without using such a sintering aid. Crystals grow under the binding conditions, and a sintered body having a large crystal grain size is obtained. In addition, by classifying the synthetic powder after firing, even when a material with a large primary particle size is selected for sintering, a sintered body with a large crystal particle size can be obtained without using a sintering additive. To be Therefore, in these cases, a sintering aid may or may not be used, but when the particle size of the synthetic powder after firing has a relatively large and broad distribution, a sintering aid should be added. is necessary.

By sintering under such conditions, a dense sintered body having a relative density of 99.8% or more and a crystal grain size of 4 μm or more can be obtained. This sintered body has a predetermined thickness,
It is preferably cut to a size of 1.8 mm or more to obtain a phosphor element. In a crystal with such a high relative density, the number and size of the vacancies at the triple points of grain boundaries are significantly reduced,
Light scattering is reduced. As a result, the optical path length of the light passing through the sintered body is shortened, and as a result, the light transmittance of the phosphor element can be improved.

The phosphor element of the present invention has a thickness of 1.8 mm.
With the above, high radiation absorbency can be secured, and thus high detectability can be exhibited when used as a radiation detector. Specifically, change the thickness of the scintillator
By setting the length to 1.8 mm or more, it is possible to absorb X-rays by 95% or more when irradiated with X-rays having a tube voltage of 120 kV.

The thus produced sintered body of the present invention has a main emission wavelength of 550 nm when absorbing radiation, and an absorption coefficient μ for light of this wavelength is 0.6 mm ^-1 or less. Converting this absorption coefficient into the light transmittance of a phosphor element having a thickness of 1.8 mm will be 35% or more. When the light transmittance is 35% or more, a decrease in sensitivity to radiation on the low energy side can be reduced, and a phosphor element having small energy dependence can be obtained.

[0021]

EXAMPLES Examples of the phosphor element of the present invention will be described below.

Example 1 As a starting material powder, Gd with a particle size of about 0.1-5 μm₂O₃, A
l₂O₃, Ga₂O₃, Ce₂(C₂O _Four)₃・ 9H₂Use O powder and mix these
After combining, the primary temperature is set by firing at 1550 ℃ -1650 ℃.
Gd with a particle size of 2 to 6 μm₃(Al, Ga)_FiveO₁₂: Ce phosphor powder
Obtained. From this powder, only powder with a primary particle size of 4 μm or more
Separately, press-molded, 1500 ° C in vacuum, 500kgf / cm²
And hot-sintered with the same crystal grains as the phosphor powder
A sintered body having a diameter was obtained. Machine the obtained sintered body
A ceramic scintillator having a thickness of 1.8 mm was created.

The relative density of this sintered body was 99.9%.
The scintillator had an absorption coefficient μ for light having a wavelength of 550 nm of 0.57 mm ⁻¹ and a light transmittance of 36% for light having a wavelength of 550 nm. In this case, the sensitivity to X-rays of 60 keV is 76 or more when the sensitivity at 100% light transmittance is 100, and it is shown that the detection sensitivity is sufficient even in the low energy region.

Comparative Example 1 The same raw material powder as in Example 1 was used and fired at a temperature of 1500 ° C. to give Gd ₃ (Al, Ga) having a primary particle size of about 1 to ₃ μm.
₅ O ₁₂ : Ce phosphor powder was obtained. After this phosphor powder was pressure-molded, the same sintering conditions as in Example 1 (in vacuum, 1500 ° C., 50 ° C.) were used.
It was sintered with a hot press of 0 kgf / cm ² ) to obtain a sintered body having a crystal grain size (about 1 to 3 μm) similar to the primary grain size. This was machined to obtain a ceramic scintillator having a thickness of 1.8 mm.

The relative density of this sintered body was 99.9%.
The absorption coefficient μ of the scintillator having a wavelength of 550 nm was 0.63 mm ^-1 , and the light transmittance of the light having a wavelength of 550 nm was 32%. In this case, the sensitivity to X-rays of 60 keV was reduced to 68 when the sensitivity was 100 when the light transmittance was 100%, indicating that sufficient detection sensitivity could not be obtained in the low energy region.

Example 2 As a starting material powder, Gd having a particle size of about 0.01 to 0.1 μm
_{2 O 3, Al 2 O 3} , Ga 2 O 3, Ce 2 (C 2 O 4) 3 · 9H 2 with O powder were mixed them by firing at 1400 ° C. to 1500 ° C., primary particle Gd ₃ (Al, Ga) ₅ O ₁₂ : Ce with a diameter of 0.05 to 0.1 μm
A phosphor powder was obtained. This powder is pressure-molded and vacuumed for 15
Hot press sintering was performed at 00 ° C and 500 kgf / cm ² . Under these sintering conditions, the crystals grew larger than the primary grain size of the phosphor powder, and a sintered body having a grain size of about 10 μm was obtained. This sintered body was machined to produce a ceramic scintillator having a thickness of 1.8 mm.

The relative density of this sintered body was 99.9% or more. The scintillator had an absorption coefficient μ of 0.30 mm ⁻¹ for light having a wavelength of 550 nm and a light transmittance of 58% for light having a wavelength of 550 nm. In this case, the sensitivity to X-rays of 60 keV was 87 when the sensitivity at 100% light transmittance was 100, and the detection sensitivity in the low energy region was further improved as compared with the scintillator of Example 1.

Example 3 The same raw material powder as in Example 1 was used and fired at a temperature of 1500 ° C. to obtain Gd ₃ (Al, Ga) having a primary particle size of about 1 to ₃ μm.
₅ O ₁₂ : Ce phosphor powder was obtained. 0.1% by weight of LiF as a sintering aid was added to this phosphor powder, and the mixture was pressure-molded and then sintered under the same sintering conditions as in Example 1 (in vacuum, 1500 ° C., 500 kgf / cm ² hot press). did. Under these sintering conditions, crystals grew larger than the primary grain size of the phosphor powder, and a sintered body having a grain size of about 5 μm was obtained. This was machined to obtain a ceramic scintillator having a thickness of 1.8 mm.

The relative density of this sintered body was 99.9%.
The scintillator had an absorption coefficient μ for light having a wavelength of 550 nm of 0.53 mm ^-1 , and a light transmittance of 39% for light having a wavelength of 550 nm. The sensitivity to X-rays of 60 keV was 78 when the sensitivity at 100% light transmittance was 100, and it was shown that sufficient detection sensitivity can be obtained even in the low energy region.

Example 4 The same raw material powder as in Example 1 was used and fired at a temperature of 1500 ° C. to obtain Gd ₃ (Al, Ga) having a primary particle size of about 1 to ₃ μm.
₅ O ₁₂ : Ce phosphor powder was obtained. After 0.1% by weight of LiF as a sintering aid was added to the phosphor powder and pressure molding was performed, the phosphor powder was sintered under atmospheric pressure at 1600 to 1700 ° C., which is a temperature just below the melting point of the phosphor. Even under these sintering conditions, crystals grew larger than the primary particle size of the phosphor powder, and a sintered body having a crystal particle size of about 5 μm was obtained. Machine this to the thickness
A 1.8 mm ceramic scintillator was obtained.

The relative density of this sintered body was 99.9%.
The scintillator had an absorption coefficient μ for light having a wavelength of 550 nm of 0.51 mm ⁻¹ and a light transmittance of 39% for light having a wavelength of 550 nm. In this case, the sensitivity to X-rays of 60 keV was 78 when the sensitivity at 100% light transmittance was 100, and it was shown that sufficient detection sensitivity can be obtained even in the low energy region.

As can be seen from the above examples, the phosphor element of the present invention has an absorption coefficient of 0.6 mm for light having a wavelength of 550 nm.
^{When it was -1} or less, the luminous efficiency was high, and the sensitivity was low even with low-energy radiation. Such a phosphor element is suitable as a phosphor element for a radiation detector as described below, and a radiation detector with high detectability can be obtained.

[0033]

BEST MODE FOR CARRYING OUT THE INVENTION A radiation detector according to the present invention and a medical image diagnostic apparatus using the same will be described below.

FIG. 2 shows a cross section of the radiation detector of the present invention. This radiation detector includes a plurality of phosphor elements 1, a plurality of photoelectric conversion elements 2 bonded to the phosphor elements 1 by an optical adhesive that does not absorb the light emitted from the phosphor elements 1, and these photoelectric conversion elements. An element array substrate 3 in which 2 are arranged, and a light reflector 4 that covers the phosphor element 1 and reflects light from the phosphor element 1 to prevent crosstalk of the channel tube are provided. In the figure, the phosphor element 1
Although only four are shown, the number of phosphor elements can be arbitrarily changed depending on the application. The array of phosphor elements and photoelectric conversion elements may be linear or planar. For example, in the case of a radiation detector of an X-ray CT apparatus, which will be described later, several 100 phosphor elements and photoelectric conversion elements are arranged in an arc shape.

The phosphor element 1 is a ceramic scintillator having a garnet structure containing Ce as a light emitting element and Gd, Al, Ga and O as main elements, and has a main emission wavelength of 550 nm and absorbs light of this wavelength. The coefficient μ is 0.6 mm ^-1 or less. The thickness is preferably 1.8 mm or more in order to absorb 95% or more of radiation.

The photoelectric conversion element 2 converts the light emitted by the phosphor element 1 into an electric signal, and is preferably made of a material having a high response speed and excellent conversion efficiency at a wavelength of 550 nm. As such a material, for example, a Si-PIN photodiode can be used. The light-reflecting material 4 is a material that transmits radiation but reflects light emitted by the phosphor element.
TiO ₂ resin is used, but is not limited thereto.

The radiation detector having such a structure is
As the phosphor element 1, a ceramic scintillator having a garnet structure containing Gd, Al, Ga, and O as main elements is used, so that the luminous efficiency is significantly higher than that of conventional CdWO ₄ , and the absorption coefficient μ at a wavelength of 550 nm is 0.6. mm ^-1 or less, light transmittance
Since it is 35% or more, high sensitivity can be maintained even for low-energy radiation. INDUSTRIAL APPLICABILITY The radiation detector of the present invention can be applied not only as a radiation detector of a medical image diagnostic apparatus described below but also as an industrial radiation detector, and can realize high radiation detection ability.

Next, a medical image diagnostic apparatus using the above radiation detector will be described. FIG. 3 is a diagram showing an outline of an X-ray CT apparatus which is an embodiment of the medical image diagnostic apparatus of the present invention. This X-ray CT apparatus includes a scan gantry unit 10 for imaging a subject and an image reconstructing unit 20 for reconstructing a CT image based on the measurement data obtained by the scan gantry unit 10, The gantry unit 10 has a rotating disk 11 having an opening 14 into which a subject is loaded, an X-ray tube 12 mounted on the rotating disk 11, and an X-ray tube 12 attached to the rotating disk 11. A collimator 13 for controlling the radiation direction and an X-ray detector 15 mounted on the rotating disk 11 so as to face the X-ray tube 12.
A detector circuit 16 for converting the X-rays detected by the X-ray detector 15 into a predetermined signal; and a scan control circuit 17 for controlling the rotation of the rotating disk 11 and the width of the X-ray flux. .

The image reconstructing section 20 includes an input device 21 for inputting the name of the subject, inspection date and time, inspection conditions, etc., and detector circuit 16.
An image calculation circuit 22 for performing a calculation process on the measurement data S1 sent from to reconstruct a CT image, the CT image created by the image calculation circuit 22, the subject name and examination date and time input from the input device 21. , An image information adding section 23 for adding information such as inspection conditions, and a CT image signal S2 to which image information is added
And a display circuit 24 that adjusts the display gain of and output to the display monitor 30.

Scan gantry unit 10 and image reconstruction unit
The structure of 20 is the same as that of the conventional X-ray CT apparatus, but here the radiation detector of the present invention described above is used as the X-ray detector 15. That is, the X-ray detector 15 includes a large number of detection elements (for example 96
(0 pieces) arranged in an arc shape, and the phosphor element of the present invention is used as a phosphor element that constitutes a scintillator.

In this X-ray CT apparatus, X-rays are irradiated from the X-ray tube 12 with the subject lying on a bed (not shown) installed in the opening 14 of the scan gantry unit 10. The X-ray has directivity by the collimator 13 and is irradiated to the subject as a fan-beam-shaped X-ray. Then, the X-rays transmitted through the subject are detected by the X-ray detector 15. On this occasion,
By continuously irradiating X-rays and rotating the rotary disk 11 around the subject, transmitted X-rays are detected while changing the X-ray irradiation direction. The measurement data S1 for obtaining the CT image of the subject is obtained by one rotation of the rotating disk in the case of full scan and half rotation in the case of half scan.

The rotation time of this rotating disk tends to be shortened in recent years, specifically, it is 0.5 to 1 second / revolution. X
The ray detector 16 turns on and off several hundred times during this one rotation, and detects the X-rays that have passed through the subject. X-ray detector
Although the X-ray dose detected by 16 will decrease as the rotation time is shortened, the X-ray CT apparatus of the present invention uses the X-ray detector 16 having high radiation sensitivity and low energy dependence. Therefore, even if the rotation time is shortened, high output can be obtained,
As a result, a high quality CT image can be obtained. The obtained image is displayed on the display monitor 30.

In the above embodiments, the X-ray CT apparatus has been described as an example of the medical image diagnostic apparatus, but the medical image diagnostic apparatus of the present invention can also be applied to an apparatus using γ rays as a radiation source.

[0044]

According to the present invention, it is possible to provide a phosphor element having high luminous efficiency and high sensitivity to radiation having a wide range of energy. Further, by using such a phosphor element, it is possible to provide a radiation detector having high radiation detection ability and low energy dependence. Furthermore, by providing this radiation detector, it is possible to provide a medical image diagnostic apparatus which has excellent detectability of radiation passing through the subject and is capable of obtaining a high-quality CT image.

[Brief description of drawings]

FIG. 1 is a diagram illustrating the energy dependence of light transmittance and radiation sensitivity of a phosphor element.

FIG. 2 is a diagram showing an embodiment of a radiation detector of the present invention.

FIG. 3 is a diagram showing an embodiment of a medical image diagnostic apparatus of the present invention.

[Explanation of symbols]

DESCRIPTION OF SYMBOLS 1 ... Phosphor element, 2 ... Photoelectric conversion element, 11 ... Rotating disk,
12 ... X-ray tube, 15 ... X-ray detector, 20 ... Image reconstruction unit

─────────────────────────────────────────────────── ─── Continuation of front page (51) Int.Cl. ⁷ Identification code FI theme code (reference) G01T 1/20 G01T 1/20 G 1/202 1/202 H04N 5/321 H01L 27/14 C04B 35/00 H H04N 5/321 H01L 27/14 K (72) Ichiro Miura Ichiro Miura 1-14-1 Uchikanda, Chiyoda-ku, Tokyo F-term (reference) 2H088 EE01 EE02 FF02 FF04 GG10 GG19 GG20 JJ04 JJ05 JJ37 4G030 AA11 AA14 AA34 AA36 BA00 CA04 GA11 GA23 4H001 CA02 CA04 CA08 CF02 XA08 XA13 XA31 XA64 YA58 4M118 AB01 CA01 CB11 5C024 AX11 AX16 CX41 CY47 GX02 GX09

Claims (8)

[Claims] 1. Ce as a light emitting element, at least Gd, Al,
What is claimed is: 1. A phosphor element using a phosphor composed of a garnet-structured host crystal containing Ga and O, wherein the absorption coefficient μ for light having a wavelength of 550 nm is 0.6 mm ^-1 or less. 2. The phosphor element according to claim 1, wherein the relative density is 99.8% or more and the average crystal grain size is 4 μm.
A phosphor element having the above features. 3. The phosphor element according to claim 1, having a thickness of 1.8 mm or more. 4. The phosphor is manufactured by a manufacturing method including a step of synthesizing raw material powders to obtain a phosphor powder having an average particle size of 4 μm or more, and a step of sintering the phosphor powder. The phosphor element according to any one of claims 1 to 3, characterized in that. 5. The phosphor is manufactured by a manufacturing method including a step of synthesizing raw material powders to obtain a phosphor powder having an average particle size of 1 μm or less, and a step of sintering the phosphor powder. The phosphor element according to any one of claims 1 to 3, characterized in that. 6. The phosphor is manufactured by a manufacturing method including a step of synthesizing a raw material powder to obtain a phosphor powder, and a step of sintering the phosphor powder at a temperature just below a melting point under normal pressure. The phosphor element according to any one of claims 1 to 3, characterized in that. 7. A radiation detector comprising a phosphor element that emits light by radiation and a photoelectric conversion element that detects light emission by the phosphor, wherein the phosphor element is used.
7. A radiation detector comprising the phosphor element according to any one of items 1 to 6. 8. A radiation source, a radiation detector arranged to face the radiation source, a rotating disk which holds the radiation source and the radiation detector, and is driven to rotate around a subject, The radiation detector according to claim 7, which is used as the radiation detector in a medical image diagnostic apparatus including image reconstruction means for reconstructing a tomographic image of the subject based on the intensity of the radiation detected by the radiation. A medical image diagnostic apparatus characterized by being used.