JP2002082171A - Radiation detector and x-ray diagnostic equipment using the same - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a radiation detector which has high optical conversion efficiency and detection sensitivity of radiation, such as transmitted X-rays and can obtain photographic images of high resolution. SOLUTION: The radiation detector 1 is equipped with plural detection channels, onto which of radiation 13 emitted by a radiation source and transmitted through an object to be inspected is made incident, scintillators 2, which emits scintillation lights corresponding to the intensity of the radiation made incident on the detection channels, and an optical responding means 3 which converts the projection lights from the respective scintillators 2 into electrical signals and is characterized by that a light wavelength converting layer, which converts the wavelength of the scintillation light into the wavelength-matching optical response characteristics of the optical response means 3 is formed between the respective scintillators 2 and optical response means 3.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a radiation detector used in an X-ray diagnostic apparatus or the like. A possible radiation detector.

[0002]

2. Description of the Related Art Various types of radiation, such as X-rays, neutrons, and gamma rays, are irradiated on a subject, and the intensity distribution of the radiation transmitted through the subject is measured by a radiation detector to obtain structural or composition information of the subject. 2. Description of the Related Art A method for obtaining a two-dimensional image (radiography) is widely used as a medical X-ray diagnostic apparatus, a dangerous substance detection apparatus for baggage, and a nondestructive inspection apparatus for various structures.

[0003] For example, neutron radiography is a method of obtaining structural or compositional information of a subject as a two-dimensional image by detecting the intensity distribution of a thermal neutron beam that has passed through the subject and attenuated. Effective for inspection of hydrogen-containing compounds and composite materials composed of metals and light element substances, which are difficult to inspect with X-rays and γ-rays. Used as an effective inspection method in a wide range of fields such as plant equipment, aircraft and automobile parts. Have been.

An X-ray diagnostic apparatus (CT scanner) is
A number of X-ray detectors are arranged around a patient as a subject, and the transmitted X-ray signals received by these detectors are processed by a computer to be reconstructed as a tomographic image, which is then displayed on a display device such as a CRT. It can be displayed or obtained as a photo. Since a tomographic image obtained by the X-ray diagnostic apparatus is obtained as a sliced image of a human body, unlike a normal radiograph or the like, it is possible to diagnose diseases deep in the human body such as internal organs with higher accuracy.

[0005] Here, the main use of the X-ray detector incorporating the solid-state detection element is a computer-processed X-ray tomography apparatus (X-ray CT apparatus: CT scanner) using a computer as described above. As described above, its usefulness has also been confirmed in the field of environment measurement devices for detecting gamma rays and nuclear radiation and in the field of various computer-processed radiography.

Particularly useful as a commercial imaging technique has been computerized X-ray tomography used for medical diagnosis. Initially, X-ray CT detectors of this type include U.S. Pat. No. 4,013,396 to Whetten et al., U.S. Pat. No. 4,198,533 to Shelley et al., And Cotic.
Various detectors described in US Pat. No. 4,161,655, such as c), are commonly used. These detectors operate on the principle that high-pressure xenon gas is used as an X-ray detection medium and the intensity of X-rays is detected by proportional ionization of the xenon gas.

[0007] In the current CT scanner, when the subject is a patient, it is necessary to achieve a high scanning speed at the time of detection in order to minimize the radiation dose harmful to the patient. is there. In addition, the operation speed of the X-ray apparatus needs to be increased when the object to be inspected is the head of the human body, but is indispensable even when scanning the whole body of the human body.

[0008]

However, in the above-mentioned conventional detector, the X-ray detecting medium is X-ray.
Since a gas having a small linear absorption coefficient is used, it is necessary to increase the size of the detector in order to achieve a predetermined detection sensitivity, and the inertial weight inevitably increases, making high-speed scanning difficult. In addition, since it is impossible to reduce the pixel size on the display device screen, it is difficult to obtain a high-resolution captured image.

On the other hand, the latest multi-slice type X-ray CT
Since the apparatus (CT scanner) employs a detector row in which a number of fine detection channels for receiving radiation transmitted through the subject are arranged close to each other, a photomultiplier tube and a photomultiplier tube are provided for each detection channel. It is not possible to equip the associated high-voltage power supply. Therefore, a photodiode is generally used as a light response unit for converting light emitted from a scintillator by transmitted radiation into an electric signal. However, the photodiode itself is
Since almost no amplification effect is exhibited during photoelectric conversion, the luminous efficiency of the scintillator is extremely important in terms of characteristics.

Therefore, a solid state detector using a scintillator made of a single crystal compound such as CdWO ₄ has been used as a radiation detector of a CT scanner for high speed and high resolution. However, there has been a problem that there are many difficulties in the manufacturability, light emission characteristics, workability, and the like of the scintillator made of the single crystal. That is, most of the scintillators having high luminous efficiency include Eu, Tb,
It is formed by adding an activator such as Pr,
In order to achieve high luminous efficiency, it is necessary to achieve uniformity of the spatial distribution of the activator during the crystal growth process and to strictly control the optimum concentration of the activator, thereby increasing the productivity. There was a problem that it was extremely low. Further, a high-purity base material is required as a scintillator material, and there has been a problem that the manufacturing cost of the high-resolution CT detector becomes enormous.

Furthermore, cadmium tungstate (CdWO ₄ ) has been generally used as a constituent material of scintillators which have been commercially used in X-ray detectors for CT apparatuses up to now. However, CdWO ₄ light emitting efficiency of the X-ray excitation is extremely low defect about 2.5 to 3.0%. Therefore, since the signal output and the noise level reach almost the same level as the electronic noise level, there is a problem that the quality of the output signal is apt to be reduced.

In addition, CdWO ₄ has the property of being easily cleaved like mica at the time of cutting, so that the workability is poor, and it is extremely difficult to obtain a scintillator bar having a predetermined size without damage or defects. In addition, the contained Cd is harmful to the human body, and there is also a problem that the equipment cost for avoiding toxicity rises.

[0013] As an alternative to solve the problems of the scintillator made of a single crystal compounds such as the CdWO _4, for example, it has been proposed also a scintillator made of a polycrystalline material. For example, cadmium oxide activated with europium
A scintillator made of a rare earth oxide phosphor such as yttrium and a scintillator made of a rare earth oxysulfide phosphor such as gadolinium oxysulfide activated with praseodymium have also been put to practical use.

The above-mentioned scintillator has been conventionally used as a constituent material of a detector unit (detection channel) used in various X-ray diagnostic apparatuses (CT scanners). The detector units are arranged in a line to constitute a detector of a single-slice X-ray diagnostic apparatus (CT scanner), while the detector units are arranged two-dimensionally to provide a multi-slice X-ray diagnosis. The detector of the device is configured. In the single-slice CT scanner, only one tomographic image is obtained by one rotation of the X-ray source and the detector, whereas in the multi-slice CT scanner, a plurality of tomographic images are obtained. , More advanced diagnostic operations are possible.

In the single-slice CT scanner, the height of the detector unit is about 30 mm. In the multi-slice CT scanner, the height of each detector unit is increased in order to obtain a clearer tomographic image. Is formed as thin as about 1 mm, so that there is a problem that the light emission characteristics are apt to deteriorate.

That is, in the above-mentioned multi-slice type CT scanner, since the size of each detector is miniaturized, the absolute amount of light emission from the scintillator is also reduced. As a result, the signal current converted by the photodiode is also reduced. There is a disadvantage that it becomes smaller.

On the other hand, in order to clearly capture a moving object such as a heart, which is a living body, and to reduce the amount of X-ray exposure to a patient, the scanning speed is increased by one revolution per minute. It is a premise for practical use to increase the speed from 1 second per rotation to about 0.5 seconds per rotation. In order to respond to such technical requirements, it is necessary for the scintillator used in the detector to have higher luminous efficiency, higher sensitivity, and shorter afterglow characteristics.

However, scintillators made of rare earth oxide phosphors such as gadolinium oxide and yttrium activated with europium, which have been used in CT scanners to date, have a drawback that the afterglow characteristics are long. On the other hand, scintillators made of rare earth oxysulfide phosphors such as gadolinium oxysulfide activated with praseodymium have low sensitivity characteristics, and in any case, the light conversion efficiency and detection sensitivity of radiation such as X-rays are low. There is a problem that it is difficult to obtain a low and high-resolution photographed image.

The present invention has been made to solve the above problems, and in particular, has a high light conversion efficiency and high detection sensitivity for radiation such as transmitted X-rays, and is capable of obtaining a high-resolution captured image. The purpose is to provide a vessel.

[0020]

Means for Solving the Problems To achieve the above object, the present inventors have to increase the intensity of scintillation light incident on a photodiode in order to avoid the difficulties associated with the conventional scintillator made of a polycrystalline compound. Various studies have been made on the configuration of a radiation detector capable of detecting the radiation. As a result, especially when a light wavelength conversion layer made of a transparent resin containing an organic fluorescent dye or an inorganic fluorescent substance is formed between a scintillator and a light responsive means such as a photodiode, the light wavelength conversion layer causes the scintillation light to be generated. Is effectively converted to a wavelength compatible with the optical response characteristics of the optical response means, the detection signal from the optical response means is greatly amplified, and as a result, the light conversion efficiency and detection sensitivity of transmitted X-rays are reduced. It was found that a high-resolution photographed image could be obtained for the first time. The present invention has been completed based on the above findings.

That is, the radiation detector according to the present invention includes a plurality of detection channels for receiving radiation projected from a radiation source and transmitted through a subject, and scintillation light corresponding to the intensity of the radiation incident on each detection channel. A radiation detector comprising: a scintillator for emitting light; and a light responsive means for converting light emitted from each scintillator to an electric signal, wherein the light for converting the wavelength of the scintillation light into a wavelength suitable for the light response characteristic of the light responsive means. A wavelength conversion layer is formed between each of the scintillators and the light responding means.

The light wavelength conversion layer is preferably composed of a resin layer in which an organic fluorescent dye is uniformly dissolved in a transparent synthetic resin. Further, the light wavelength conversion layer may be composed of a resin layer in which an inorganic phosphor is uniformly dispersed in a transparent synthetic resin.

The organic fluorescent dye is at least one organic compound selected from a coumarin compound, a rhodamine compound, a sulfolodamine compound, a dicyanomethyl compound (DCM), a styryl compound and a pyromethane compound. It is preferred that Further, the inorganic phosphor includes manganese-activated magnesium fluorogermanate (2.5 MgO.MgF ₂ : Mn ⁴⁺ ),
Cerium-activated yttrium aluminate (Y ₃ Al ₅ O
₁₂ : Ce) and at least one phosphor particle selected from europium-activated yttrium oxysulfide (Y ₂ O ₃ S: Eu).

Each of the scintillators has a general formula Ba
It is preferable to use a polycrystalline ceramic formed by adding europium to a barium fluorohalide represented by FX (where X is at least one halogen element of Cl, Br and I). Further, the porosity of each scintillator is preferably 0.1% or less.

An X-ray diagnostic apparatus according to the present invention is characterized in that the above-mentioned radiation detector is provided as an X-ray detector.

According to the radiation detector of the present invention,
Since the wavelength of the scintillation light emitted from the scintillator is effectively converted by the optical wavelength conversion layer into a wavelength that matches the optical response characteristics of the optical response means, the detection signal from the optical response means increases significantly, resulting in Thus, the light conversion efficiency and detection sensitivity of radiation such as transmitted X-rays are increased, and a high-resolution captured image is obtained.

[0027]

Next, an embodiment of the present invention will be described.
This will be described with reference to the accompanying drawings. FIG. 1 shows an X-ray diagnostic apparatus (X-ray diagnostic apparatus) having an X-ray detector as a radiation detector according to the present invention.
FIG. 1 is a cross-sectional view illustrating a schematic configuration of an example of a line CT scanner.

An X-ray diagnostic apparatus 10 shown in FIG. 1 includes a radiation source 11 that emits radiation such as X-rays, a collimator 12 that converts the emitted X-rays into a thin fan-shaped X-ray beam 13, and a subject 14. An X-ray detector 1 as a radiation detector for detecting a transmitted X-ray beam 13; a computer 15 for performing computer processing on a detection signal of the X-ray detector 1 to reconstruct a subject image 17; And a display (CRT) 16 for displaying the information.

The radiation source 11 is usually composed of a rotating anode type X-ray tube. As shown in FIG. 1, in a multi-tomography type CT apparatus, a large number of detection channels (detection cells) 1a are two-dimensionally arranged. The X-ray detector 1 is configured as a detector row that is arranged in a random manner.

As shown in FIGS. 2 and 1, each detection channel (detection cell) 1a constituting the X-ray detector 1 is a radiation (X-ray) 13 projected from a radiation source 11 and transmitted through a subject 14. And a plurality of detection channels 1a for allowing the radiation
In a radiation detector 1 including a scintillator 2 that emits scintillation light corresponding to the intensity of the light 3 and an optical response unit 3 that converts the light emitted from each scintillator 2 into an electric signal, the wavelength of the scintillation light is changed by an optical response. The light wavelength conversion layer 4 for converting the light into a wavelength suitable for the light response characteristic of the means 3 is formed between each of the scintillators 2 and the light response means 3.

The X-rays projected from the X-ray source 11 pass through a collimator (slit window) 12 to become a quadrangular pyramid-shaped X-ray beam 13 having a predetermined aperture angle. The X-ray beam 13 passes through a subject 14 such as a patient arranged and fixed in the passage, and reaches the X-ray detector 1 configured by arranging a large number of detection channels 1a two-dimensionally. The X-ray beam 13 transmitted through the subject 14 attenuates according to the density of each part of the subject 14. Each transmitted X-ray beam 13 is X
Each detection channel 1 of the X-ray detector 1 facing the radiation source 11
a.

Then, as shown in FIG. 2, while the transmitted X-ray beam 13 entering each detection channel 1a passes through the scintillator 2, scintillation light corresponding to the intensity of the transmitted X-ray beam 13 is emitted. The wavelength of the scintillation light is converted by the light wavelength conversion layer 4. That is, the wavelength is converted to a wavelength that matches the optical response characteristics of the optical response unit 3 such as a photodiode disposed on the secondary side of the optical wavelength conversion layer 4. Therefore, the electric signal output from the light responding means 3 can be greatly increased, and as a result, the X-ray detection sensitivity can be increased.

The intensity of the electric signal output from the light responding means 3 is a measure indicating the degree of attenuation of the X-ray beam 13 in each part of the subject through which the X-ray beam has passed.

In actual X-ray photography, the X-ray source 11
The X-ray detector 1 is rotated while performing X-ray imaging around a subject (patient) 14 fixed inside the opening, and at a plurality of angular positions with respect to the subject 14, each of the detection channels 1 a The electric signal is read as image information. The electric signal read by the X-ray photography is digitized and processed by the computer 15 by computer. That is, the computer 15 calculates an image of the cross section of the subject 14 through which the fan-shaped X-ray beam 13 has passed by using one of a plurality of available algorithms, and reconstructs the image as the subject image 17. I do. The reconstructed image is displayed on the display 16. The subject image 17 is, for example, a tomographic image of the subject 14. In a multi-slice type X-ray CT apparatus, the subject 14
Are photographed at the same time. According to such a multi-slice type X-ray CT apparatus, it is also possible to stereoscopically describe an imaging result.

X-ray diagnostic apparatus (CT scanner) shown in FIG.
In 10, the X-ray source 11 is fixed so as to face the X-ray detector 1, and is configured to rotate around the subject 14 at a high speed in a state where the relative positions of both are fixed. . In this case, in order to achieve a higher spatial resolution, at the position of the detector cell at the center of a large number of two-dimensionally arranged detection channels 1a, the cell interval should be as small as possible. Is required.

From the above viewpoint, the detection channel constructed according to the present embodiment is a very desirable form for constituting the above-mentioned detection channel row. As shown in FIG. 2, the structure of each detection channel (single cell) 1 a of the X-ray detector 1 used in the present embodiment is such that the cell width of each detection channel 1 a is in the area facing the fan-shaped X-ray beam 13. A pair of collimator plates 6 for defining the thickness of the entrance window 5 for receiving X-rays along the
7 is determined.

Here, in order to achieve a good spatial resolution, it is important to divide the fan-shaped X-ray beam into as many pixel channels as possible, and for that purpose, the X-ray detector 1 is constructed. It is necessary to make the cell size of each detection channel 1a as small as possible. In this embodiment, the cell length of the entire detection channels arranged two-dimensionally, that is, the cell size in the direction perpendicular to the plane of the fan-shaped X-ray beam is set to 40 to 60 mm, and the length of one row is set to 1 ~ 2mm
And the cell width is specified to be about 1 mm.

In the detection channel 1a, the transmitted X-rays 13 passing between the collimator plates 6 and 7 and entering the scintillator 2 are converted into low-energy radiation (scintillation light) in the visible or near-visible spectrum range. You.

In the conventional X-ray detector, the scintillation light emitted from the scintillator is directly supplied to a light responsive means such as a photodiode, and converted into an electric signal proportional to the intensity of the scintillation light in the light responsive means. I was However, since the wavelength of the scintillation light is not a wavelength suitable for the spectral response characteristics (energy conversion efficiency) of the photodiode, there is a fatal problem that the light conversion efficiency is low.

FIG. 3 is a graph showing a spectral response characteristic (energy conversion efficiency) of a photodiode as an optical response means generally used in an X-ray diagnostic apparatus (CT scanner). In order to convert scintillation light into an electric signal with high conversion efficiency, the scintillation light needs to be 550 to 850.
It is preferably radiation having a wavelength in the nm spectral range. However, phosphor materials that convert X-rays into radiation such as visible light are widely known, but their emission spectrum is not necessarily limited to a spectral region of 550 to 850 nm, which is an emission wavelength suitable for the spectral response characteristics of the photodiode. Highly efficient scintillator phosphors have not been developed yet.

Therefore, in the present invention, as shown in FIG.
A light wavelength conversion layer 4 is provided between the scintillator 2 and a photodiode assembly 3 as a light responsive means. The light wavelength conversion layer 4 converts the light emission wavelength of the scintillator 2 into radiation having a light emission wavelength suitable for the light response characteristics of the photodiode, and exhibits an effect of further increasing the X-ray detection sensitivity.

The light wavelength conversion layer may be composed of a resin layer in which an organic fluorescent dye is uniformly dissolved in a transparent synthetic resin, but may be composed of a resin layer in which an inorganic phosphor is uniformly dispersed in a transparent synthetic resin. May be.

Examples of the resin material serving as a carrier for the organic fluorescent dye and the inorganic fluorescent material include polymethacrylate,
Polyvinyl chloride, vinyl chloride-vinyl acetate copolymer, alkyd resin, aromatic sulfonamide resin, urea resin,
Melamine resin, benzoguanamine resin and the like can be suitably used.

The organic fluorescent dye is at least one organic compound selected from a coumarin compound, a rhodamine compound, a sulfolodamine compound, a dicyanomethyl compound (DCM), a styryl compound and a pyromethane compound. Preferably, there is. DC above
Specific examples of M include 4- (dicyanomethylene) -2-
Methyl-6- (P-dimethylaminostyryl) -4H-
Piran and others.

The above-mentioned inorganic phosphor is preferably a manganese-activated magnesium fluorogermanate (2.5MgO.Mg).
F ₂ : Mn ⁴⁺ ), at least one phosphor particle selected from cerium-activated yttrium aluminate (Y ₃ Al ₅ O ₁₂ : Ce) and europium-activated yttrium oxysulfide (Y ₂ O ₃ S: Eu) It is preferred that

The light wavelength conversion layer can be composed of not only a single layer but also a plurality of layers containing fluorescent materials having mutually different characteristics. For example, scintillation light having a wavelength of about 400 mm is converted to a longer wavelength of 450 to 550 n.
m into blue-green light having a wavelength of about m, and further convert this converted light into longer wavelength scintillation light,
By finally converting the light into red light of rhodamine type or sulfoldamine type, it is possible to further increase the photoelectric conversion efficiency of the photodiode.

In the X-ray detector 1 shown in FIG.
In order to efficiently guide the scintillation light emitted by the scintillator 2 to the active surface of the photodiode 3, the following surface treatment may be performed. For example, the surface of the scintillator 2 other than the side surface in contact with the photodiode 3 may be treated so that the scintillation light is reflected toward the inside of the scintillator 2 to prevent scattering of the scintillation light.

The material constituting the scintillator 2 used in the X-ray detector 1 according to the present invention is not particularly limited as long as it can convert X-rays into radiation such as visible light (scintillation light) with high efficiency. It is not something to be done. However, from the viewpoint of the stability of characteristics, each of the scintillators is prepared by adding europium to a barium fluorohalide represented by a general formula BaFX (where X is at least one halogen element of Cl, Br and I). More preferably, it is made of a polycrystalline ceramic of an alkaline earth metal fluoride halide phosphor formed by the above method.

The scintillator made of the above polycrystalline ceramic is prepared by the following manufacturing method. That is, a phosphor powder of a predetermined composition is cold isostatically pressed (C
IP) method or a molding method, and the obtained molded body is sintered in a vacuum, an inert gas atmosphere or a weak reducing atmosphere, and the obtained sintered body is cut, polished, etc. By performing machining, a rod-shaped scintillator having a predetermined dimension can be easily prepared.

In particular, in order to form a high-density and dense polycrystalline ceramic, it is necessary to use a phosphor raw material powder whose particle diameter is adjusted to an optimum range. Specifically, the particle diameter of the raw material powder is preferably 3 μm or less, and more preferably 0.5 to
A range of 1.5 μm is more preferred.

When the phosphor raw material powder is molded by the CIP process, 147 MPa (1.5 ton / cm) is used.
² ) The above hydrostatic pressure may be added. The packing density of the molded product obtained by this molding operation is 60 to the theoretical density.
Reaches 80%.

Further, the sintering operation is carried out by using a high-temperature compression technique such as a hot press method or a hot isostatic press (HIP) method. The sintering temperature is preferably set in the range of 600 to 1200 ° C. If the sintering temperature is lower than 600 ° C., it is difficult to remove pores formed in the ceramic sintered body. On the other hand, if the sintering temperature exceeds 1200 ° C., the crystal size of the polycrystalline ceramics becomes coarse, mechanical strength and the like are reduced, and workability is deteriorated and finished dimensional accuracy of processing is also unfavorably reduced. Also,
The sintering pressure needs to be 40 MPa or more in order to form a particularly dense scintillator.

In the X-ray detector according to the present invention, the porosity of each scintillator is preferably 0.1% or less in order to improve the transparency of scintillation light. If the porosity exceeds 0.1%, the scintillation light is easily scattered, and the scintillation light is not easily transmitted to the light responsive means.

As a preferred method for producing the scintillator used in the present invention, the following method can be adopted. That is, the CIP molded body formed as described above is accommodated in a mild steel capsule lined with platinum foil, and HIP treatment is performed at a temperature of 900 ° C. for 2 hours under a pressure of 45 MPa in an argon gas atmosphere. Or the above-mentioned CIP molded body is placed in a graphite mold,
Under vacuum of × 10 ⁻² Torr, temperature 650 ° C., 50M
A method of performing a hot press treatment for 3 hours by applying a pressure of Pa is preferably used.

In these preferred production methods, B
E as an activator for a phosphor body having a composition of aFX
A rare earth element such as u is added in a range of 0.1 to 5 mol%. The activator made of the rare earth element is used to enhance the efficiency of converting transmitted X-rays incident on the scintillator into scintillation light having a wavelength suitable for the optical response characteristics of an optical response means such as a photodiode. It is added as a contained component.

According to the above manufacturing method, a scintillator made of dense polycrystalline ceramics having good workability can be obtained without requiring excessive manufacturing costs. The polycrystalline ceramic for a scintillator obtained by the above-described manufacturing method has a high density substantially equal to the theoretical density such that the porosity is 0.1% or less, and has high visible light conversion efficiency by X-ray irradiation. It is.

However, the above-mentioned scintillator made of polycrystalline ceramics exhibits a blue or ultraviolet emission having an emission peak at a wavelength of about 380 to 400 nm, and has a drawback that it does not conform to the above-described spectral sensitivity characteristics of the photodiode. That is, since the scintillator emits light in the blue to ultraviolet region, there is a disadvantage that the photoelectric conversion efficiency is low even when the scintillator directly irradiates a silicon photodiode that converts scintillation light into an electric signal.

Therefore, in the present invention, as shown in FIG. 2, the light wavelength conversion layer 4 is formed between the scintillator 2 and the photodiode 3. The light wavelength conversion layer 4 adjusts the wavelength of blue or ultraviolet light emitted from the scintillator 2 to 550-7.
By converting the light into a wavelength in the red or infrared region of 00 nm to match the spectral response characteristics of the photodiode 3, the photoelectric conversion efficiency of the photodiode 3 is greatly increased.

Among the various alkaline earth metal fluorinated halides studied as constituent materials of the scintillator by the present inventors, BaFCl is the most characteristically preferable for the following reasons. That is, the BaFCl has a considerably large X-ray stopping power over the entire energy range of radiation generally used in the field of radiology including the diagnostic X-ray apparatus. X
Irradiation of radiation is reduced, and the amount of exposure to harmful radiation can be significantly reduced.

BaFCl is made of a polycrystalline ceramic which is close to the theoretical density and has little porosity and high luminous efficiency by using a high-temperature compression technique generally used as a method for producing a ceramic sintered body. It can be formed as Incidentally, the luminous efficiency of the BaFCl: Eu scintillator to which europium (Eu) is added as an activator reaches about 13%.

Other phosphors which have been found to have crystallographic properties similar to BaFCl and scintillation properties useful for use in CT scanner scintillators include BaFBr and BaFI. These phosphors are advantageous in that they have an X-ray blocking ability equal to or higher than that of BaFCl. However, BaFI has the disadvantage that it is deliquescent and is slightly unstable chemically. Therefore, a polycrystalline ceramic such as BaFCl, which has good chemical stability and low radiation embrittlement, is optimal as a scintillator component.

In the above-mentioned scintillator made of polycrystalline ceramic, the wavelength of the scintillation light,
The conversion characteristics such as the decay time, the amount of residual light, and the degree of the hysteresis phenomenon largely depend on the type of activator made of a rare earth element contained in the scintillator. When europium is added at an optimal concentration of 0.1 to 5 mol% as an activator, the main emission wavelength of the scintillator is 360 to 4
The range is 00 nm.

In order to adapt the wavelength of the scintillation light emitted from the scintillator to the spectral response characteristics of the photodiode, the present invention uses a light wavelength conversion layer containing an organic fluorescent dye or an inorganic fluorescent material as shown in FIG. 4
It is provided between the scintillator 2 and the photodiode 3.

FIG. 4 is a graph showing an emission spectrum of a scintillator in the case where a light wavelength conversion layer containing sulfolodamine or the like as an organic fluorescent dye is formed. As shown in FIG. 4, the peak wavelength of the emission spectrum of the scintillator having the light wavelength conversion layer is 560 to 620 nm.
This value is very close to the peak response wavelength of a silicon photodiode as an optical response means used for a CT scanner as shown in FIG. Therefore, by forming the light wavelength conversion layer as described above, a scintillator that can convert light having a wavelength suitable for the light responsive means can be obtained, and at the same time, the photoelectric conversion efficiency in the light responsive means greatly increases, Is improved, and a subject image having high resolution can be obtained.

All the conversion characteristics of the scintillator formed of the polycrystalline ceramic as described above can be similarly suitably used as a scintillator of a radiation detector for CT (computer tomography) and DR (digital radiography). . The primary decay time of the scintillation light emitted from the scintillator is 1
Although it is 0 microsecond or less, it can be further reduced by increasing the concentration of europium as an activator contained in the scintillator, or by adding cerium (Ce) in combination.

The X-ray detector according to the present embodiment shown in FIG. 2 is used as an X-ray detector for a multi-slice type CT apparatus as shown in FIG. 6 in addition to a single slice type CT apparatus as shown in FIG. The same can be applied. In the CT apparatus shown in FIG. 5, the X-ray beam 1
3 transmits through the subject 14. The transmitted X-rays are detected by the X-ray detectors 1 arranged in a single row and converted into electric signals, while the CT apparatus shown in FIG. The light passes through the subject 14. The transmitted X-rays are detected by the X-ray detectors 1 arranged in double rows and converted into electric signals.

In the single-slice CT apparatus shown in FIG. 5, only one tomographic image can be obtained by an imaging operation in which the X-ray source 11 and the X-ray detector 1 make one rotation. On the other hand, in the multi-slice type CT apparatus shown in FIG. 6, a plurality of tomographic images can be obtained, so that a more advanced and precise diagnosis can be made. The tomographic image is displayed on a display (CRT) 17 as shown in FIG. 1 and stored in the form of a film or the like in order to contribute to an examination by a diagnostician at a later date.

Next, a more specific embodiment of the radiation detector according to the present invention will be described.

Example First, BaFCl: Eu (Eu concentration = 0.2 mol%) phosphor powder having an average particle diameter of 2 μm was molded by a rubber press. This molded body is housed in a Ta capsule,
While applying a hydrostatic pressure of 45 MPa by argon gas to the capsule, the capsule was heated to a temperature of 900 ° C. and subjected to hot isostatic pressing (HIP) for 2 hours to produce a large number of polycrystalline ceramics. The porosity of the obtained polycrystalline ceramic was 0.05%. This polycrystalline ceramic was cut and polished to produce scintillators for each example.

On the other hand, as shown in Table 1, 2% by weight of FM17C (rhodamine-based organic dye: manufactured by Shinroka Co., Ltd.) or an organic fluorescent dye was added to a solution of polyvinyl butyral as a transparent synthetic resin material dissolved in ethanol. Resin slurries were prepared by blending sulfo-rhodamine at a ratio of 0.02% by weight.

Next, the operation of applying the resin slurry and the operation of drying and curing the resin slurry on the side opposite to the X-ray incidence side of the scintillator for each embodiment prepared as described above were repeated. The scintillators for each of Examples 1 to 8 in which the light wavelength conversion layer having the thickness shown in FIG.

FIG. 4 is a graph showing the emission spectra of the scintillators 2 for Examples 1 to 5 in which the light wavelength conversion layer 4 containing sulfolodamine and the like as an organic fluorescent dye was formed. As shown in FIG. 4, the peak wavelength of the emission spectrum of the scintillator 2 having the light wavelength conversion layer 4 is 560 to 560.
620 nm, which is very close to the peak response wavelength of a silicon photodiode as an optical response means used in a CT scanner application as shown in FIG.

Next, as shown in FIG. 2, each of the scintillators 2 having the above-mentioned light wavelength conversion layer 4 integrally formed at the end of a detection cell having a cell width of 1 mm divided by a pair of collimator plates 6 and 7.
The X-ray detector 1 as a radiation detector according to the first to eighth embodiments can be mounted by integrally incorporating a photodiode assembly 3 as a light response unit on the secondary side of the light wavelength conversion layer 4. Each was prepared.

COMPARATIVE EXAMPLE On the other hand, except that no resin slurry containing an organic fluorescent dye or the like was applied and no light wavelength conversion layer was formed, the same scintillator and photodiode assembly as those of the examples were used. An X-ray detector according to a comparative example having the same dimensions as the examples was prepared.

In order to compare and evaluate the photoelectric conversion effect of the X-ray detectors according to each of the examples and the comparative examples prepared as described above, X-rays of a predetermined intensity are incident on each X-ray detector.
The electric signal output from the photodiode assembly was measured. However, the electric signal intensity is shown as a relative value with the electric signal intensity output from the detector of the comparative example in which the light wavelength conversion layer is not formed as a reference value (100%). The measurement results are shown in Table 1 below.

[0076]

[Outside 1]

[0077]

[Table 1]

As is clear from the results shown in Table 1 above,
According to the X-ray detector 1 according to each embodiment in which the light wavelength conversion layer 4 containing a predetermined phosphor is formed between the scintillator 2 and the photodiode 3 as a light responsive means, the light is emitted from the scintillator 2. Since the wavelength of the scintillation light is effectively converted by the light wavelength conversion layer 4 into a wavelength suitable for the light response characteristics of the photodiode 3, the photoelectric conversion efficiency of the photodiode 3 is significantly higher than that of the detector of the comparative example. It was found to be improved. as a result,
It has been found that the detection sensitivity of the X-ray detector 1 is improved, and a high-resolution captured image can be obtained.

In the above embodiment, an X-ray detector used in a medical X-ray diagnostic apparatus has been described as an example. However, the radiation detector according to the present invention is not limited to the X-ray medical diagnostic apparatus. The present invention can be similarly applied to not only the inspection apparatus but also an X-ray non-destructive inspection apparatus for industrial use. The radiation detector according to the present invention greatly contributes to improvement of detection accuracy by the X-ray nondestructive inspection device.

[0080]

As described above, in the radiation detector according to the present invention, the wavelength of the scintillation light emitted from the scintillator is effectively adjusted to the wavelength suitable for the optical response characteristics of the optical response means by the optical wavelength conversion layer. Since the signal is converted, the detection signal from the light responding means is greatly increased, and as a result, the light conversion efficiency and detection sensitivity of radiation such as transmitted X-rays are increased,
A high-resolution captured image can be obtained.

[Brief description of the drawings]

FIG. 1 is a sectional view showing a schematic configuration of an embodiment of an X-ray diagnostic apparatus (CT apparatus) using a radiation detector according to the present invention.

FIG. 2 is a cross-sectional view showing a configuration of one detection channel constituting a radiation detector used in the X-ray diagnostic apparatus shown in FIG.

FIG. 3 is a graph showing a spectral response characteristic of a silicon photodiode.

FIG. 4 is a graph showing an emission spectrum of the light wavelength conversion layer.

FIG. 5 is a schematic view showing a diagnostic operation by a single-slice type CT apparatus.

FIG. 6 is a schematic diagram showing a diagnostic operation by a multi-slice type CT apparatus.

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 Radiation detector (X-ray detector) 1a Detection channel (detection cell) 2 Scintillator 3 Photoresponse means (photodiode) 4 Optical wavelength conversion layer 5 Incident window 6,7 Collimator plate 10 X-ray diagnostic apparatus (CT scanner) 11 Radiation source (X-ray tube) 12 Collimator (slit window) 13 X-ray beam 14 Subject (patient) 15 Computer 16 Display (CRT) 17 Subject image

──────────────────────────────────────────────────の Continued on the front page (51) Int.Cl. ⁷ Identification FI FI Theme Court ゛ (Reference) G01T 1/24 G01T 1/24 1/29 1/29 C (72) Inventor Eiji Koyanatsuzu Isogo, Yokohama-shi, Kanagawa 8 Shinsugita-cho, Toku-ku, Tokyo Toshiba Yokohama Office (72) Inventor Yukihiro Fukuda 8 Shinsugita-cho, Isogo-ku, Yokohama, Kanagawa, Japan 1 Kokai Toshiba Town F-term in Toshiba R & D Center (reference) 2G088 EE02 FF02 GG20 GG21 JJ02 JJ14 4C093 AA22 BA07 CA02 CA31 EB12 EB18 EB20

Claims (8)

[Claims] 1. A plurality of detection channels for receiving radiation projected from a radiation source and transmitted through a subject, a scintillator for emitting scintillation light corresponding to the intensity of the radiation incident on each detection channel, and A light response means for converting the emitted light into an electric signal, and a light wavelength conversion layer for converting the wavelength of the scintillation light into a wavelength suitable for the light response characteristic of the light response means, and each of the scintillators and A radiation detector formed between the radiation detector and the optical response means. 2. The radiation detector according to claim 1, wherein the light wavelength conversion layer comprises a resin layer in which an organic fluorescent dye is uniformly dissolved in a transparent synthetic resin. 3. The radiation detector according to claim 1, wherein the light wavelength conversion layer comprises a resin layer in which an inorganic phosphor is uniformly dispersed in a transparent synthetic resin. 4. The organic fluorescent dye is at least one organic compound selected from a coumarin compound, a rhodamine compound, a sulfolodamine compound, a dicyanomethyl compound (DCM), a styryl compound and a pyromethane compound. 3. The method according to claim 2, wherein
A radiation detector as described. 5. The method according to claim 1, wherein the inorganic phosphor is manganese-activated magnesium fluorogermanate (2.5MgO.MgF).
₂ : Mn ⁴⁺ ), at least one phosphor particle selected from cerium-activated yttrium aluminate (Y ₃ Al ₅ O ₁₂ : Ce) and europium-activated yttrium oxysulfide (Y ₂ O ₃ S: Eu). The radiation detector according to claim 3, wherein the radiation detector is provided. 6. Each of the scintillators has a general formula BaFX
Wherein said barium fluorohalide is a polycrystalline ceramic formed by adding europium to barium fluorohalide represented by the formula (where X is at least one halogen element of Cl, Br and I). 2. The radiation detector according to 1. 7. A porosity of each of the scintillators is 0.1%.
The radiation detector according to claim 1, wherein: 8. An X-ray diagnostic apparatus comprising the radiation detector according to claim 1 as an X-ray detector.