JP3302463B2 - Radiation detector and method of using the same - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a radiation detector, and more particularly to a radiation detector utilizing a combination of a scintillator and a solid-state light detector.

[0002]

2. Description of the Related Art Conventionally, intensifying screens have been used in various fields of radiography such as medical radiography for radiography for medical diagnosis and industrial radiography for destructive inspection of substances. A so-called radiographic method in combination with a radiographic film is used. According to this method, when radiation such as X-rays transmitted through the subject enters the intensifying screen, the phosphor contained in the intensifying screen absorbs the energy of the radiation and emits fluorescence (instantaneous emission). By this light emission, the radiographic film superimposed on the intensifying screen is exposed, and a radiographic image is formed on the radiographic film. In this way, the radiation image can be obtained directly as an image visualized on the radiation film.

On the other hand, in various fields, a radiographic image recorded on a radiographic film is photoelectrically read to obtain an image signal, the image signal is subjected to appropriate image processing, and the image is reproduced and recorded. Have been done. For example, an X-ray image is recorded using a film having a low gamma value designed to be compatible with the subsequent image processing, and the X-ray image is read from the film on which the X-ray image is recorded and converted into an electric signal. By subjecting the electric signal (image signal) to image processing and reproducing it as a visible image in a copy photograph or the like, a reproduced image having good image quality performance such as contrast, sharpness, and granularity is obtained. (See Japanese Patent Publication No. 61-5193).

Further, the applicant of the present invention reported that radiation (X-ray,
Irradiation with α-rays, β-rays, γ-rays, electron beams, ultraviolet rays, etc.) accumulates a part of this radiation energy, and then irradiates with excitation light such as visible light to produce stimulated emission according to the accumulated energy. Using a stimulable phosphor (stimulable phosphor), radiation image information of a subject such as a human body is temporarily recorded in a sheet-like stimulable phosphor, and the stimulable phosphor sheet is excited with a light beam such as a laser beam. To generate stimulable luminescence light, photoelectrically read the obtained stimulable luminescence light to obtain an image signal,
A radiation image recording / reproducing system for outputting a radiation image of a subject as a visible image on a recording material such as a photographic photosensitive material or a CRT based on the image data has already been proposed (Japanese Patent Application Laid-Open Nos. 55-12429 and 56-56). No. 11395, No. 55-163472,
Nos. 56-104645 and 55-116340).

[0005] This system has the practical advantage of being able to record images over a very wide radiation exposure area compared to conventional radiographic systems using silver halide photography. That is, in the case of the stimulable phosphor, it has been recognized that the amount of emitted light that is stimulated by excitation after accumulation is proportional to the radiation exposure amount over an extremely wide range. Even if fluctuates considerably, the amount of the stimulating light emitted from the stimulable phosphor sheet is read by the photoelectric conversion means with the reading gain set to an appropriate value and converted into an electric signal. Recording materials such as photographic photosensitive materials using
By outputting a radiation image as a visible image on a display device such as a CRT, it is possible to obtain a radiation image that is not affected by a change in radiation exposure.

However, in order to obtain a radiographic image by the above-described radiographic system, when the above-mentioned radiographic image is directly visualized, the radiographic film used for radiography and the sensitivity area of the intensifying screen are matched. There is a need to do.

Further, in the above-described system for reading a radiographic image photoelectrically using the radiographic film and the stimulable phosphor sheet, as described above, the radiographic image is subjected to image processing to obtain a density and contrast according to the purpose. It has to be adjusted to have a radiation image once converted into an electrical signal, it is necessary to perform scanning using an image reading device for that, the operation to obtain the radiation image becomes complicated, It takes a long time to obtain a radiation image.

In order to solve the above-mentioned problems in the conventional system, a radiation detector has been proposed (for example, Japanese Patent Application Laid-Open No. 59-211263, Japanese Patent Application Laid-Open No. 2-1640).
No. 67, PCT International Publication No. WO 92/06501, Signa
l, noise, and read out considerations in the develop
ment of amorphous silicon photodiode arrays for rad
iotherapy and diagnostic x-ray imaging, LEAnton
uk et.al, University of Michigan, RAStreet Xero
x 、 PARC 、 SPIE Vol.1443 Medical Imaging V; Image Phy
sics (1991), pp. 108-119).

This radiation detector is composed of a plurality of solid-state photodetectors arranged in a matrix composed of a transparent conductive film and a transparent conductive film with an amorphous semiconductor film interposed therebetween on a substrate made of, for example, quartz glass having a thickness of 3 mm. As described above, a scintillator for converting radiation into visible light is stacked on a solid-state photodetector composed of a plurality of signal lines and scanning lines patterned in a matrix.

The radiation detector is arranged so that the scintillator faces the surface on the radiation incident side, and by irradiating the radiation detector with radiation transmitted through the subject, the radiation directly enters the scintillator and is converted into visible light. Alternatively, when the scintillator is a stimulable phosphor, radiation image information stored and recorded in the scintillator is converted into stimulable luminescence by excitation with excitation light, and this visible light or stimulable luminescence is It is detected by the photoelectric conversion unit of the solid-state photodetector and is photoelectrically converted into an image signal carrying radiation image information. The image signal is read from a transfer unit provided in each solid-state light detecting element of the radiation detector by a predetermined reading unit, and after being subjected to predetermined image processing, is reproduced by a reproducing unit such as a CRT. By using such a radiation detector, a radiation image of a subject can be immediately reproduced without performing a complicated operation. Therefore, a radiation image can be obtained immediately in real time. Can be eliminated.

On the other hand, the same subject is irradiated with X-rays having different energy distributions, and a specific structure (eg, organ, bone, blood vessel, etc.) of the subject has a specific X-ray energy absorption characteristic. The two image signals in which a specific structure is drawn differently are obtained by using the image signal, and after the two image signals are appropriately weighted, subtraction is performed between the two signals to perform subtraction between the two signals. A so-called energy subtraction method for extracting an image is known (for example, JP-A-59-83486).

In this method, one shot energy subtraction for simultaneously storing and recording a radiation image carrying a high energy component and a low energy component of radiation respectively on two stimulable phosphor sheets is performed. Obtaining an image signal carrying a radiation image of the subject;
By performing a subtraction process between the image signals, an image in which a specific structure of the subject is emphasized is obtained.

Various stimulable phosphor sheets for performing such energy subtraction have been proposed. For example, JP-A-3-211500 proposes a stimulable phosphor sheet in which two kinds of stimulable phosphors having different radiation absorption characteristics are laminated and a shutter layer is interposed between the respective stimulable phosphors. ing. According to this stimulable phosphor sheet, a radiation image of a different energy component is stored and recorded in each stimulable phosphor while the shutter layer is open, and from one stimulable phosphor with the shutter layer closed. Excitation light for obtaining an image signal can be prevented from being incident on the other stimulable phosphor. Japanese Unexamined Patent Publication (Kokai) No. 2-298899 discloses a stimulable phosphor formed by separating a plurality of types of stimulable phosphors having different image characteristics (e.g., energy absorption characteristics and sensitivity) from each other in a direction along a sheet surface. Body sheets have been proposed. Using this stimulable phosphor sheet, the radiation images having different energy absorption characteristics are accumulated and recorded in each stimulable phosphor,
The above-described energy subtraction image can be obtained by obtaining an image signal from the radiation image stored and recorded in each of the stimulable phosphors, and performing a weighted subtraction process between the image signals.

[0014]

However, when energy subtraction processing is performed using the above-described stimulable phosphor sheet, a plurality of types of stimulable phosphors are simultaneously irradiated with radiation having different energy absorption characteristics. Even if the energy components are separated by interposing a filter or the like, the separation of the image signal obtained from the high-energy component radiation image and the image signal obtained from the low-energy component radiation image may be insufficient. there were. In addition, when taking an image, radiation can be applied in a state where the stimulable phosphors are superimposed on each other. However, since the reading of the image signal is performed separately from each of the stimulable phosphors, There is a problem that it is difficult to accurately perform the alignment between the image signals when performing the subtraction processing.

When energy subtraction is performed, an image signal carrying a high energy component of the radiation and an image signal carrying a low energy component are required. Since images carrying low energy components cannot be obtained at the same time, it is necessary to perform so-called two-shot energy subtraction in which two shots are taken. However, when the two-shot energy subtraction is performed, a positional shift may occur between the corresponding two images, and as a result, there is a problem that a good subtraction image cannot be obtained.

Further, it is necessary not only to perform the energy subtraction but also to perform the addition processing of the image signals in order to perform the alignment between the image signals.

The present invention has been made in view of the above circumstances, and has high accuracy in the positioning of a plurality of image signals for performing the above-described energy subtraction processing or addition processing. It is an object of the present invention to provide a radiation detector that can perform satisfactorily.

[0018]

A first radiation detector according to the present invention comprises: an instantaneous light-emitting phosphor which emits instantaneous light which carries radiation image information when irradiated with radiation; A scintillator including a stimulable phosphor that emits stimulating luminescence light that carries the radiation image information upon irradiation of excitation light, where radiation image information is stored and recorded,
A single solid state photodetector that converts the instantaneous emission light and the photostimulated emission light emitted from the scintillator into respective image signals and outputs the image signals is stacked.

The second radiation detector according to the present invention is the second radiation detector according to the present invention, wherein the instantaneous light-emitting phosphor and the stimulable phosphor have different radiation absorption characteristics. It is characterized by the following.

Further, a third radiation detector according to the present invention is the first or second radiation detector according to the present invention, wherein the scintillator includes the instantaneous light-emitting phosphor and the stimulable phosphor, respectively. Characterized by the above scintillator.

A fourth radiation detector according to the present invention is the radiation detector according to the above-described present invention, wherein the instantaneous emission light and the photostimulated emission light are transmitted between the scintillator and the solid state light detector. And a filter that does not transmit the excitation light.

According to a fifth radiation detector of the present invention, in the above-described radiation detector of the present invention, the instantaneous emission light and the photostimulated emission light are transmitted between the two scintillators, and the excitation light is emitted. It is characterized in that a filter that does not transmit light is interposed.

Further, a sixth radiation detector according to the present invention is the radiation detector according to the above-described present invention, wherein the scintillator is interposed on the opposite side to the solid-state light detector,
An excitation light source that emits the excitation light is further provided.

Further, the image signal reading method according to the present invention comprises:
The above-described first, third, fourth, fifth or sixth radiation detector according to the present invention is irradiated with radiation, and the solid is generated by the instantaneous emission light emitted from the instantaneous emission phosphor by the irradiation of the radiation. Reading the first image signal output from the photodetector, and then irradiating the scintillator with the excitation light to read the second image signal output from the solid-state photodetector by the stimulated emission light; 1st read
And superimposing processing by adding the second image signal and the second image signal.

Further, another image signal reading method according to the present invention is directed to irradiating the first, third, fourth, fifth or sixth radiation detector according to the present invention with radiation, and irradiating the radiation with the radiation. The first image signal output from the solid-state photodetector is read by the instantaneous light emitted from the instantaneous light-emitting phosphor, and then the scintillator is irradiated with the excitation light to emit the solid by the stimulated emission light. A second image signal output from the photodetector is read, and the read first and second image signals are weighted and subtracted to perform an energy subtraction process.

[0026]

According to the first radiation detector of the present invention, a scintillator including an instantaneous light-emitting phosphor that emits instantaneously emitted light and a stimulable phosphor that emits stimulated light is provided. A single solid-state photodetector that detects emitted light, converts it into an image signal, and outputs the image signal is stacked. For this reason, the image signal output from the solid-state photodetector by the instantaneous emission light and the image signal output from the solid-state photodetector by the stimulable emission light emitted from the stimulable phosphor by the irradiation of the excitation light. , And their positions correspond to each other.
For this reason, by adding the respective image signals, it is possible to perform a superposition process of the image signals with high positioning accuracy with a simple configuration.

Further, the instantaneous light-emitting phosphor and the stimulable phosphor have different radiation absorption characteristics, and the image signal output from the solid-state photodetector is weighted and subtracted from each of the instantaneous light and the stimulable light. By performing the processing, it is possible to perform the energy subtraction processing with high alignment accuracy with a simple configuration.

Further, by using the scintillator as two scintillators each containing an instantaneous luminescent phosphor and a stimulable phosphor, the instantaneous luminescent light and stimulable luminescent light emitted from each phosphor can be well separated. Thus, the information carried by the image signal output from the solid-state photodetector can be satisfactorily separated by the instantaneous emission light and the stimulated emission light emitted from each scintillator.

Further, by providing a filter between the scintillator and the solid-state photodetector, which transmits the instantaneous emission light and the stimulated emission light but does not transmit the excitation light, the scintillator is irradiated with the excitation light to emit the stimulated emission light. The excitation light does not enter the solid-state photodetector at the time of obtaining, and an image signal with less noise due to the excitation light can be obtained.

Further, with respect to the scintillator separated into two pieces, by providing the above-mentioned filter between each scintillator, the width of the wavelength of light emitted from the phosphor contained in the scintillator in contact with the solid-state photodetector is determined. The degree of freedom can be increased, and the width of the phosphor that can be used can be increased.

Further, by providing the two-dimensional excitation light source for emitting the excitation light on the opposite side of the scintillator from the solid-state light detector, the excitation light and the radiation detector can be integrally formed. An image signal reading system using the detector can be configured compactly.

[0032]

Embodiments of the present invention will be described below with reference to the drawings.

FIG. 1 is a diagram showing a first embodiment of the radiation detector according to the present invention. As shown in FIG. 1, a radiation detector 1 according to an embodiment of the present invention includes a planar scintillator 3 and a solid-state photodetector 2 stacked with a filter 7 interposed therebetween. Here, the scintillator 3 contains two types of phosphors, a stimulable phosphor and a normal phosphor. In this embodiment, the stimulable phosphor has an emission peak at 400 nm. BaF (BrI): Eu ²⁺ ,
CaW with emission peak at 420 nm as a normal phosphor
It contains O ₄ . In the present embodiment, the stimulable phosphor is excellent in the absorption characteristic of the low energy component of the radiation, and the ordinary phosphor is excellent in the absorption characteristic of the high energy component of the radiation. The filter 7 cuts the excitation light 8A so that the excitation light 8A emitted from the excitation light source 8 having LEDs arranged two-dimensionally does not enter the solid-state photodetector 2. n
m, the light having a wavelength of 600 nm is cut, and the light emitted from the stimulable phosphor and the normal phosphor is emitted.
, 420-nm optical filter is used to pass light having a wavelength of 420 nm.

Here, the details of the solid-state photodetector 2 will be described. FIG. 2 is a diagram illustrating details of the solid-state photodetector. As shown in FIG. 2, the solid-state photodetector 2 includes a plurality of solid-state photodetectors.
The solid-state photodetector 18 has signal lines 12A and 12B made of a conductive film patterned on a substrate 11 made of a resin sheet, and a photoelectric conversion element made of an amorphous silicon 13 and a transparent electrode 14. It comprises a photodiode 15 as a unit, a thin film transistor 17 as a transfer unit comprising an amorphous silicon 16 and a transfer electrode 16A (gate) provided in the amorphous silicon 13. By arranging a plurality of the solid-state photodetectors 18 having such a configuration in a two-dimensional manner, a solid-state photodetector 2 is formed. By stacking the solid-state photodetector 2 on a scintillator 3, the radiation detector 1 It is configured. Here, the thickness of the resin sheet 11 is about several hundred microns, and the X-ray absorptivity is low. The thickness of the amorphous silicon 13 is about 1 micron.

The X-ray 5 emitted from the X-ray source 4 is
And is transmitted through the subject 6. The X-ray 5 transmitted through the subject 6 is applied to the radiation detector 1. Radiation detector 1
The X-rays 5 radiated onto the scintillator 3 constituting the radiation detector 1 are first radiated. As a result, the normal phosphor contained in the scintillator 3 emits light instantaneously according to the radiation image information of the subject 6, and this instantaneous light is emitted by the solid-state photodetector 2.
The light is received by the photodiode 15 as a photoelectric conversion unit of each of the solid-state light detecting elements 18 that constitute the solid-state photodetector 18. At the same time, radiation image information of the subject 6 is accumulated and recorded on the stimulable phosphor contained in the scintillator 3. Then, the instantaneous light is photoelectrically converted, and signal charges are generated in the photodiode 15 according to the light emission intensity. Thereafter, the signal charges are read out, and an image signal S1 as an electric signal carrying radiation image information of a low energy component of the subject 6 is output. The output image signal S1 is logarithmically converted by the logarithmic conversion unit 20, and then digitized by the A / D conversion unit, and the digital image signal logS1 obtained in this manner.
Are temporarily stored in the memory 22.

Next, the excitation light 8A is emitted from the excitation light source 8 and radiated to the scintillator 3. As a result, the stimulable phosphor contained in the scintillator 3 is excited, and stimulable luminescent light carrying radiation image information of the subject 6 stored and recorded in the stimulable phosphor is emitted. The light is received by a photodiode 15 as a photoelectric conversion unit of each solid-state light detecting element 18 constituting the solid-state light detector 2. At this time, since the excitation light 8A is cut by the filter 7, the excitation light 8A
Is not detected by the solid-state light detector 2. Then, the stimulated emission light is photoelectrically converted, and signal charges are generated in the photodiode 15 according to the emission intensity. Thereafter, the signal charges are read out, and an image signal S2 is output as an electric signal carrying a high-energy component radiation image of the subject 6. The output image signal S2 is logarithmically converted by the logarithmic conversion means 20, and then digitized by the A / D conversion means 21, and the digital image signal log thus obtained is obtained.
S1 is stored in the memory 22.

When the image signals S1 and S2 are both stored in the memory 22 as described above, both signals are read out from the memory and input to the calculating means 23.

The arithmetic means 23 is a circuit for inputting two image signals.
logS1, logS2 on which was suitably weighted and subtracted for each corresponding solid-state photodetector element, a digital difference signal _{Ssub = a · logS 1 -b ·} logS 2 + c [a · b is the weighting factor, c is Bias component]. The difference signal Ssub is subjected to image processing such as gradation processing and frequency processing in the image processing means 24, and then sent to the reproduction means 25 for reproduction of a radiation image. The reproducing means 25 may be a display means such as a CRT, or a recording device for performing optical scanning recording on a photosensitive film, or may temporarily store an image signal in an image file such as an optical disk or a magnetic disk. It may be replaced by a device.

The reason why the digital image signal is treated as a logarithmic value is that the band of the image data is compressed and unnecessary image information can be completely removed. It is also possible to do the same with image signals.

If the coefficients a and b are appropriately determined when performing the above-described subtraction operation, the obtained difference signal Ssu
In b, the signal components of the part other than the specific structure in the subject 6 are deleted. Therefore, if an image is reproduced based on the difference signal Ssub, a radiation image in which only the specific structure is extracted can be obtained.
When performing the subtraction calculation, it is necessary to perform the calculation between the corresponding pixels as described above. Here, according to the present invention, the position of each solid-state light detecting element 18 constituting the solid-state light detector 2 of the radiation detector 1 is invariable, and the instantaneous emission light emitted from two kinds of phosphors included in the scintillator 3 is used. Since the positional relationship between the radiation image information carried by the stimulated emission light is the same, no special operation such as using a marker or fixing the reading position is required, and the image signal S1 is simply output from the radiation detector 1. , S2, and the image signals S1 and S2 are read out and the corresponding solid-state light detecting elements 18 are subtracted in correspondence with each other, whereby the alignment has already been performed.

Next, the second embodiment of the radiation detector according to the present invention will be described.
An example will be described.

FIG. 3 is a view showing a second embodiment of the radiation detector according to the present invention. As shown in FIG.
The radiation detector 1 according to the embodiment includes the above-described BaF (Br).
I): Eu ²⁺ or the like and a normal phosphor such as CaWO ₄ are contained in separate scintillators 3A and 3B, and a scintillator 3A containing a stimulable phosphor and a normal phosphor are used. A scintillator 3B is laminated, and the solid state detector 2 is laminated via the scintillators 3A and 3B and the filter 7. Note that arrows shown in FIG. 3 indicate directions in which radiation and excitation light are incident.

When the X-rays 5 passing through the subject are irradiated on the radiation detector 1 in the same manner as in the first embodiment of the present invention,
The scintillator 3A stores and records the radiation image information of the subject 6, and the scintillator 3B emits light instantaneously according to the radiation image information of the subject 6. This instantaneous emission light is detected by each solid-state light detection element 18 constituting the solid-state light detector 2, and an image signal S1 corresponding to the emission intensity of the instantaneous emission light is output from the radiation detector 1.

Next, excitation light 8A is emitted from the excitation light source 8, and the excitation light 8A is irradiated on the scintillator 3A. As a result, the scintillator 3A is excited, and stimulated emission light carrying radiation image information of the subject 6 stored and recorded in the scintillator 3A is emitted. An image signal S2 detected by the detection element 8 and corresponding to the emission intensity of the stimulated emission light is output from the radiation detector 1.

The two image signals S1 and S2 output from the radiation detector 1 are subjected to energy subtraction processing in the same manner as in the first embodiment of the present invention described above, and the image signals subjected to the energy subtraction processing are reproduced by the reproducing means. At 25, it is reproduced as an energy subtraction image.

In the above-described second embodiment of the present invention, a scintillator 3B including a normal phosphor that emits light instantaneously is disposed closer to the solid-state photodetector 2 among the two scintillators 3A and 3B. However, the scintillator 3A including a phosphor that stimulates emission may be arranged closer to the solid-state photodetector 2.

Next, a third embodiment of the present invention will be described.

FIG. 4 is a diagram showing a third embodiment of the present invention. As shown in FIG. 4, the radiation detector 1 according to the third embodiment of the present invention, as shown in the second embodiment of the present invention, separates a stimulable phosphor and a normal phosphor into separate scintillators 3A and 3B. The scintillators 3A and 3B are provided with a filter 7 that transmits instantaneous emission light and photostimulated emission light but does not transmit excitation light.

A radiation detector 1 according to a third embodiment of the present invention
The readout of the image signal from the third embodiment is the same as that of the above-described second embodiment of the present invention, and therefore detailed description is omitted here. However, in the third embodiment of the present invention, the instantaneous light emitted from the scintillator 3B is used. Since the emitted light is directly incident on the solid-state photodetector 2 without passing through the filter 7, the degree of freedom of the wavelength of the phosphor that emits light instantaneously increases, and the wavelength of the stimulated emission light emitted from the stimulable phosphor is limited. Instead, a phosphor that emits light instantaneously can be freely selected.

Next, a fourth embodiment of the present invention will be described.

FIG. 5 is a diagram showing a fourth embodiment of the present invention. As shown in FIG. 5, the radiation detector 1 according to the fourth embodiment of the present invention includes a scintillator 3 containing a stimulable phosphor and a normal phosphor and a solid state, similarly to the above-described first embodiment of the present invention. The photodetector 2 is laminated with a filter 7 interposed therebetween. The solid-state photodetector 2 is arranged on the side where X-rays are incident, and a two-dimensional shape that emits excitation light on the side of the scintillator 3. Are arranged so as to be integrated with the radiation detector 1.

Here, when X-rays enter the scintillator 3, the low energy components of the X-rays are
The high-energy component of the X-rays stays near the surface on the side where the X-rays are incident, and on the back side away from the surface. For this reason, when a stimulable phosphor having excellent X-ray high-energy component absorption characteristics is used, the high-energy component of X-rays stays on the side of the scintillator 3 where the X-ray was not irradiated. Therefore, when the excitation light is irradiated from the side where the high-energy component of X-rays is stagnated, the stimulated emission that carries the radiation image information containing more high-energy components of X-rays stored and recorded in the scintillator 3 is more preferable. Light is emitted.

Therefore, in the radiation detector 1 according to the fourth embodiment of the present invention, the energy of the image signal S1 obtained by the instantaneous emission and the image signal S2 obtained by the stimulated emission can be better separated. As a result, better energy subtraction processing can be performed.

Here, in the above-described fourth embodiment of the present invention, since the solid-state photodetector 2 is disposed on the side where X-rays are incident, each solid-state photodetector constituting the solid-state photodetector 2 is provided. The substrate on which the element 18 is provided must have a property of transmitting X-rays.

In the first to fourth embodiments of the present invention described above, the energy subtraction processing is performed on the two image signals S1 and S2 output from the solid-state photodetector. Signals S1 and S2
May be subjected to a weighted addition process (superposition process).

That is, two image signals logS1, l converted into digital signals by the arithmetic means 23 shown in FIG.
For ogS2, Sad = k1 logS1 + k2 logS2 (2) where k1, k2: weighting coefficient may be calculated. By performing the superposition process on the two image signals S1 and S2 in this manner, the image signals can be emphasized, and a reproduced image with higher sharpness can be obtained.

In the above-described embodiment, a filter that transmits instantaneous light and stimulated emission light but does not transmit excitation light is interposed between the scintillator and the solid state light detector or between the two scintillators. However, as long as the solid-state photodetector has a radiation detection characteristic of detecting instantaneous light emission and stimulated emission light but not excitation light, it is not particularly necessary to interpose a filter.

Further, in the above-described embodiment, the stimulable phosphor that emits stimulable light and the ordinary phosphor that emits light instantaneously have different radiation absorption characteristics. In particular, it is not necessary to make the radiation absorption characteristics different for each phosphor.

In the above embodiment, CaWO ₄ is used as a normal phosphor that emits light instantaneously, and BaF (BrI): Eu ²⁺ is used as a stimulable phosphor that emits light. There is no particular limitation, and any phosphor may be used as the ordinary phosphor as long as it emits light instantaneously. For example, SrFBr: Eu ²⁺ , Gd
₂ O ₂ S: Tb ³⁺ or the like may be used. Also,
As the stimulable phosphor, any phosphor may be used as long as it emits stimulable light.

In the above-described embodiment, a B-410 optical filter is used as a filter, and a wavelength of 410 nm is used.
Although only the light in the vicinity is transmitted, any filter may be used as long as it transmits instantaneous light and stimulated light and does not transmit excitation light.

Further, in the above-described embodiment, the amorphous silicon layer is used as the semiconductor layer. However, the present invention is not limited to this, and any semiconductor layer may be used.

[0062]

As described above in detail, in the radiation detector and the image signal reading method according to the present invention, a phosphor that emits light instantaneously and a phosphor that emits stimulated light are separately provided and emitted from each phosphor. The instantaneous emission light and the stimulated emission light are each detected by one solid-state photodetector to obtain separate image signals. Therefore, the alignment of the image signals can be performed easily and accurately, and each image signal can be detected. Separation can be performed favorably. Therefore, in the superimposition processing using each image signal, the image quality can be further improved, and in the energy subtraction processing, an energy subtraction image with good energy separation can be obtained.

[Brief description of the drawings]

FIG. 1 is a diagram showing a radiation detector according to a first embodiment of the present invention.

FIG. 2 is a cross-sectional view illustrating a solid-state light detecting element included in the solid-state light detector of the radiation detector.

FIG. 3 is a diagram illustrating a radiation detector according to a second embodiment of the present invention.

FIG. 4 is a diagram showing a radiation detector according to a third embodiment of the present invention.

FIG. 5 is a view showing a radiation detector according to a fourth embodiment of the present invention;

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 Radiation detector 2 Solid-state photodetector 3 Scintillator 4 X-ray source 5 X-ray 6 Subject 7 Filter 8, 8 'Excitation light source 8A, Excitation light 20 Logarithmic conversion means 21 A / D conversion means 22 Memory 23 Arithmetic means 24 Image processing Means 25 Reproduction means S1, S2, S 'Image signal

──────────────────────────────────────────────────続 き Continued on the front page (51) Int.Cl. ⁷ Identification code FI H04N 5/32 H04N 5/325 5/325 H01L 31/00 A (56) References JP-A-5-264799 (JP, A) JP-A-64-12300 (JP, A) JP-A-5-66298 (JP, A) JP-A-5-87935 (JP, A) JP-A-63-241500 (JP, A) JP-A-5-257000 (JP, A) JP-A-58-215575 (JP, A) JP-A-62-69182 (JP, A) JP-A-57-161674 (JP, A) JP-T-5-502764 (JP, A) ( 58) Field surveyed (Int.Cl. ⁷ , DB name) G01T 1/20 G01T 1/00 G03B 42/02 H01L 31/09 H04N 5/32 H04N 5/325

Claims (8)

(57) [Claims] An instantaneous light emitting phosphor that emits instantaneous light that carries radiation image information when irradiated with radiation,
A scintillator comprising: a stimulable phosphor that emits stimulating light that carries the radiation image information by irradiating the radiation image information with the radiation image information being accumulated and irradiated with the excitation light; and the instantaneous emission emitted from the scintillator. A radiation detector, comprising: a single solid-state photodetector that converts light and the stimulated emission light into an image signal and outputs the image signal separately. 2. The radiation detector according to claim 1, wherein the instantaneous light-emitting phosphor and the stimulable phosphor have different radiation absorption characteristics. 3. The radiation detector according to claim 1, wherein the scintillator comprises two scintillators each containing the instantaneous light emitting phosphor and the stimulable phosphor. 4. Transmitting the instantaneous emission light and the stimulated emission light between the scintillator and the solid state light detector,
4. The radiation detector according to claim 1, wherein a filter that does not transmit the excitation light is interposed. 5. The radiation detection according to claim 3, wherein a filter that transmits the instantaneous emission light and the stimulating emission light but does not transmit the excitation light is interposed between the two scintillators. vessel. 6. The excitation light source according to claim 1, further comprising an excitation light source that emits the excitation light, on an opposite side of the scintillator from the solid-state photodetector. Radiation detector. 7. The solid state light detection by irradiating the radiation detector according to claim 1, 3, 4, 5, or 6 with the instantaneous light emitted from the instantaneous light emitting phosphor by the irradiation of the radiation. Reading out the first image signal output from the detector, and thereafter irradiating the scintillator with the excitation light to read out the second image signal output from the solid-state photodetector by the stimulated emission light; An image signal reading method, wherein the first and second image signals obtained are added to perform a superimposition process. 8. The solid state light detection by irradiating the radiation detector according to claim 2, 3, 4, 5, or 6 with the instantaneous emission light emitted from the instantaneous emission phosphor by the irradiation of the radiation. Reading out the first image signal output from the detector, and thereafter irradiating the scintillator with the excitation light to read out the second image signal output from the solid-state photodetector by the stimulated emission light; An image signal reading method, wherein weighted subtraction processing is performed on the obtained first and second image signals to perform energy subtraction processing.