JP2005091538A - Radiographic system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To obtain an image with inconspicuous noise, of high sharpness, high image quality and high detectability by reading an image obtained by phase contrast radiography by using a both-side light condensation reading system. <P>SOLUTION: The radiographic system is provided with a phase contrast breast image photographing device for photographing a breast with radiation from a radiation source and recording the breast as the radiograph on a stimulable phosphor plate or a phase contrast image photographing device for photographing an object with radiation from a radiation source and recording the object as the radiograph on the stimulable phosphor plate, and a radiograph both-side reading device for irradiating the stimulable phosphor plate with exciting light and reading the stimulated phosphorescent light on both sides of the stimulable phosphor plate. <P>COPYRIGHT: (C)2005,JPO&NCIPI

Description

  In this invention, phase contrast radiographic imaging is performed, and the photostimulable phosphor plate on which the radiographic image is recorded is read from both sides of the photostimulable phosphor plate. The present invention relates to a radiographic imaging system that provides a radiographic image with improved lesion detection.

  Conventionally, a radiographic imaging apparatus capable of obtaining a digital radiographic image for diagnosis of a disease or the like is known, and a brightness corresponding to the accumulated energy is obtained by storing a part of radiation energy and irradiating excitation light. CR (Computed Radiography), which is a radiographic image capturing apparatus using a stimulable phosphor that exhibits exhaustion, is widely used. A radiographic image obtained by this apparatus is obtained by detecting a two-dimensional distribution of a radiation dose emitted from a radiation source and transmitted through a subject with a radiographic image detector. That is, the image is based on the radiation absorption contrast of the subject. The reading system that reads the photostimulated luminescence from the photostimulable phosphor that stores a part of the radiation energy as image information usually irradiates one side of the photostimulable phosphor plate on which the radiation image is recorded, The photostimulated light is condensed by a light guide and converted into image information. However, as a means for reading the photostimulated luminescence light, the excitation light is scanned on both sides or one side of the photostimulable phosphor plate, and the photostimulated luminescence emitted by this excitation light scan is transmitted from both sides of the photostimulable phosphor plate. A double-sided condensing reading system such as Japanese Patent Laid-Open No. 55-87970 has been proposed that scans a light guide to collect light and converts it into image information.

  In double-sided scanning, it is well known that noise in radiographic images is reduced by superimposing image signals SA and SB obtained by reading radiographic images of the same subject from both sides of the photostimulable phosphor plate. It has been. SA is an image signal from the surface (upper side) close to the radiation source at the time of imaging, and SB is an image signal obtained from the opposite surface (lower side). The noise in the radiographic image makes it difficult to understand the difference in radiation absorption between the structure and its surroundings, which reduces the ability to detect lesions, but the overlay process takes place from both sides of the stimulable phosphor plate. By superimposing the read image signals, it is possible to obtain an image signal with good S / N ratio that averages noise and is inconspicuous.

  However, the image signal SB obtained from the surface away from the radiation source contains image information in the low frequency band as in the SA, but the image information in the high frequency band is less than that in the SA, and the scattered light. There are many noise components due to such effects. Therefore, when an image is produced by adding SA and SB as they are, the S / N ratio in the low frequency band is improved, but in the high frequency band, the influence of the noise component contained in SB becomes large and the image quality is deteriorated. Connected.

  The types of noise in the photostimulable phosphor plate include the following. (1) X-ray quantum noise, (2) Noise due to non-uniformity of phosphor coating distribution or phosphor particle distribution on the photostimulable phosphor plate, and (3) Image recorded on the photostimulable phosphor plate Noise of excitation light to be emitted, (4) noise of stimulated emission light emitted, collected and detected from the stimulable phosphor plate, and (5) electrical noise in a system for amplifying and processing an electric signal. . Since the noises (1) to (5) often show different distributions for each image of the stimulable phosphor plate, the noises are averaged by superimposing these images of the stimulable phosphor plate. The noise is not noticeable in the superimposed image. That is, an image with good graininess can be obtained.

Radiation image superimposition processing in the double-sided condensing reading method is described in, for example, Japanese Patent Laid-Open No. 8-76302. By superimposing the radiation image information read from both sides of the photostimulable phosphor plate, each noise is averaged, and the noise becomes inconspicuous in the image subjected to the superimposition processing. That is, an image signal with a good S / N can be obtained. This noise includes many noises that can be approximated by Poisson statistics, and radiation quantum noise, which is one of the dominant factors in the noise of radiographic images, is one example. Here, assuming that noise can be approximated by Poisson statistics, when two radiation images are considered to have signals S1 and S2 of the same magnitude and noises N1 and N2, respectively, two images are superimposed. The magnitudes of the signal and noise are S1 + S2 for the signal and √N1 ² + N2 ² for the noise. On the other hand, when considering S / N, which is one index representing the detection ability of radiographic images, the S / N of each image before superposition is S1 / N1 and S2 / N2, respectively. By doing so, S / N becomes (S1 + S2) / √N1 ² + N2 ² , and S / N is improved. In addition, when performing the superimposition processing, the S / N improvement can be optimized by weighting each signal.

  On the other hand, since radiation is also a kind of electromagnetic wave, it also has wave characteristics, and therefore, diffraction and refraction due to a phase shift occur when passing through a subject. In recent years, an apparatus for capturing a radiographic image with high subject contrast such as Japanese Patent Application Laid-Open No. 2001-238871 using this property has been proposed and is called a phase contrast radiographic image capturing apparatus. This imaging device is characterized by a smaller focal spot size and a longer distance between the subject table and the radiographic image detector than a conventional radiographic imaging device such as a CR. Contrast is increased and the sharpness of the entire image is dramatically improved.

  Furthermore, by reducing the image size of the radiographic image that has been digitally magnified in the phase contrast radiographic image capturing apparatus to the actual size and outputting it, the degradation of graininess caused by the magnified imaging is at the same level as the original image. You can go back up. Because of the above reasons, the radiographic image detectability is improved, and application to the medical field using radiation and the industrial nondestructive inspection field is expected.

  In order to improve the detectability of a radiographic image, a radiographic image having excellent sharpness and graininess is required. Especially in mammography images used for breast cancer diagnosis etc., it is important to simultaneously detect a mass with a faint shadow at the same time as a microcalcification group of several hundred microns or less in the early stage of cancer. In order to improve the detectability, various radiographic imaging apparatuses have been improved. Both the microcalcification group and the mass can improve the detectability by increasing the sharpness and decreasing the granularity.For example, if the contrast of the film is increased, the sharpness can be increased, but the granularity Also rises and causes deterioration of graininess.

When a phase contrast radiographic image capturing apparatus is used, sharpness is greatly improved and graininess is not deteriorated. Therefore, the degree of improvement in graininess is small with respect to the degree of improvement in sharpness, which is about the same as conventional CR. Images with high lesion detectability, such as microcalcifications, are images that have an excellent balance of graininess and sharpness. By improving the granularity of the phase contrast radiation image, the lesion detectability is further improved. High radiation image.
JP-A-55-87970 JP-A-8-76302 Japanese Patent Laid-Open No. 2001-238871

  As described above, the double-sided condensing reading method of the stimulable phosphor plate can improve the graininess but cannot improve the sharpness. In contrast, the phase contrast radiographic imaging device can dramatically improve sharpness without deteriorating the graininess compared to conventional imaging techniques. An image with high detectability can be obtained. For example, mammography images, which are medical radiographic images, are desired to improve the lesion detection performance beyond the conventional level, and an image with low granularity and high sharpness is required.

  The present invention has been made in view of the above circumstances, and by reading an image obtained by phase contrast radiographic imaging using a double-sided condensing reading method, noise is not noticeable and high sharpness is obtained. An object of the present invention is to provide a radiographic imaging system that can obtain a highly detectable image.

  In order to solve the above-described problems and achieve the object, the present invention is configured as follows.

According to the first aspect of the present invention, a breast is imaged with radiation from a radiation source, and a radiographic image of the breast is recorded on the photostimulable phosphor plate, and excitation is performed on the photostimulable phosphor plate. A radiation image double-sided reader that irradiates light and reads the photostimulated luminescent light on both sides of the photostimulable phosphor plate;
A radiographic imaging system characterized by comprising:

According to a second aspect of the present invention, there is provided a phase contrast image photographing apparatus for photographing a subject with radiation from a radiation source and recording a radiation image of the subject on a stimulable phosphor plate, and excitation light on the stimulable phosphor plate. A radiation image double-sided reader that reads the photostimulated luminescent light on both sides of the photostimulable phosphor plate,
A radiographic imaging system characterized by comprising:

  According to a third aspect of the present invention, in the radiographic image capturing system according to the first or second aspect, the phase-contrast radiographic image captured by digital enlargement can be output in an actual size.

  According to the invention of claim 4, image signals SA and SB obtained by reading a radiation image of the same subject from both sides of the stimulable phosphor plate are used between the radiation source and the stimulable phosphor plate at the time of photographing. 4. The radiographic imaging system according to claim 1, wherein weighting is performed using a weighting coefficient obtained on the basis of the distance information and the tube voltage information of the radiation source. 5.

  With the above configuration, the present invention has the following effects.

  According to the first aspect of the present invention, a radiation image is photographed by the phase contrast breast image photographing device, and the photostimulable phosphor plate on which the photographed radiation image is recorded is read by the radiation image double-side reading device, and the phase contrast is obtained. A high-definition breast image obtained by mammographic imaging can be read with a radiological image double-sided scanning device to obtain an image with a high S / N ratio, low graininess, and high image quality with high sharpness. Breast images can be obtained.

  The invention according to claim 2 is configured such that a radiation image is captured by a phase contrast image capturing device, and the photostimulable phosphor plate on which the captured radiation image is recorded is read by a radiation image double-side scanning device. It is possible to obtain a high S / N ratio image by reading a contrast-enhanced radiographic image of the chest, finger bone, etc. with a radiological image double-sided reader, with low graininess, high sharpness, and high image quality. Images can be obtained.

  According to the third aspect of the present invention, it is possible to output a phase-contrast radiation image that has been digitally enlarged and photographed in an actual size, and an actual size of the phase-contrast radiation image on the film that is output from an output device such as an imager. Therefore, there is an advantage that interpretation is easy.

  In the invention according to claim 4, when two image signals read by the radiation image double-side reading device are combined into one image by superimposing processing, the weight of both images is set to the radiation source and the photostimulable phosphor plate at the time of photographing. By performing according to the distance R and the weighting coefficient obtained from the tube voltage information, the image quality improvement effect by phase contrast imaging and double-sided condensing reading can be exhibited in the best state.

  The image signal SB obtained from the lower surface of the photostimulable phosphor plate contains image information in a low frequency band, similar to the image signal SA obtained from the upper surface near the radiation source, but has a higher frequency than the upper surface. The frequency dependence characteristic of the band is low, the image information is reduced in the high frequency band, and the noise component is increased due to the influence of scattering or the like due to transmission through the phosphor layer. Therefore, if the same weighting is applied to the image signals obtained from the upper surface and the lower surface, the image quality is improved in the low frequency band of the added image signal, but the noise component is also emphasized in the high frequency band. As a result, the image quality is degraded.

  In view of this, there has been proposed a radiographic image superimposing method capable of obtaining a higher quality image with less noise components in the superimposed image (Japanese Patent Laid-Open No. 8-76302).

  In this superposition method, image processing is performed after changing the frequency characteristic of the image signal for a desired image signal among a plurality of image signals so that the signal-to-noise ratio of the added signal is high. The added image signal and the other image signal are added to obtain an added signal. In addition, the image signal to be superimposed is subjected to Fourier transform, wavelet transform, etc., to obtain a transform coefficient signal for each of a plurality of frequency bands, and in a frequency band having a low signal-to-noise ratio according to the frequency characteristics of the image signal. The weighting coefficient is relatively small compared to the weighting coefficient of the frequency band with a high signal-to-noise ratio, and the coefficient signal for each frequency band is added, and the added coefficient signal is subjected to inverse Fourier transform, inverse wavelet transform, etc. A method is also well known in which a final added signal is obtained. By these methods, the added signal has a high signal-to-noise ratio, and a high-quality superimposed image can be obtained by reproducing the added signal.

  The phase-contrast imaged image is also input to the weighted superposition processing apparatus as described above, weighted, and output. Compared with conventional normal imaging, phase contrast imaging is performed with the subject table and the radiation source detector separated by a certain distance or more. As a result, the distance between the radiation source and the radiation image detector is larger than that for normal imaging. become longer. When the distance between the radiation source and the radiation image detector is increased, the energy distribution of the X-rays that reach the radiation image detector shifts to the higher energy side and easily reaches the lower surface of the radiation image detector. It is considered that the difference in image quality between the upper surface and the lower surface of the detector is reduced. Furthermore, in phase contrast imaging, an image with high sharpness is obtained by utilizing X-ray refraction and diffraction, so that the X-ray subject transmission power increases, that is, when the X-ray energy increases, the phase contrast image. Is difficult to obtain, and sharpness decreases.

  Therefore, the weighting when the upper and lower surfaces of the radiation image detector are overlapped is determined based on distance information regarding the distance (R1 + R2) between the radiation source and the radiation image detector and tube voltage information at the time of imaging. In the present invention, distance information on (R1 + R2) and tube voltage information at the time of photographing are input to the weighted superposition processing device, and a coefficient is selected from the weighting coefficient list in the processing device according to the information and weighted. Perform overlay processing.

  Hereinafter, although the embodiment of the radiographic imaging system of this invention is described, this invention is not limited to this embodiment. The embodiment of the present invention shows the most preferable mode of the present invention, and the terminology of the present invention is not limited to this.

  FIG. 1 is a block diagram showing a first configuration example of a radiographic image capturing system. Although the radiographic imaging system of this embodiment shows an example of an embodiment of digital phase contrast X-ray imaging, the present invention is not limited to this. In the following description, radiation and X-ray are described as synonymous.

  This radiographic image capturing system includes a phase contrast image capturing device 1 and a radiographic image double-sided reading device 7. Since this radiographic image capturing system obtains a digital X-ray image, it is possible to arbitrarily reduce or enlarge the image. In addition, gradation processing and frequency processing of the output image can be performed. In addition, it is possible to easily transfer an image to a remote place using the Internet or the like.

  An image photographed by the phase contrast image photographing device 1 is recorded in the radiation image detector 4. The radiographic image recorded in the radiographic image detector 4 is read by the radiographic image double-side reading device 7 and output to an image output device 10 such as an imager or an image display device 11 such as a liquid crystal display. Alternatively, it is also possible to send the read image signal to another radiographic imaging system via the Internet 12.

  The phase contrast image capturing apparatus 1 includes a small-focus radiation source 2, a subject table 3, a radiation image detector 4, an information input unit 5 such as imaging conditions, and control for controlling the imaging apparatus based on information from the information input unit 5. Part 6. The phase contrast image capturing is an image capturing method for capturing a radiation image by separating a distance R1 between the small focus radiation source 2 and the subject H and a distance R2 between the subject table 3 and the radiation image detector 4 by a certain distance or more. , R2 are also controlled by the control unit 6. The captured radiation image is recorded in the radiation image detector 4. The radiation image detector 4 is used for imaging again after the image is read and erased by the radiation image double-side reading device 7. The control unit 6 controls the phase contrast image photographing device 1 at the time of photographing, and at the same time records the tube voltage information v and the distance information l at the time of photographing, and stores the information in the image processing unit of the radiation image double-sided reading device 7. input.

  The radiological image double-side reading device 7 includes a radiological image reading unit 8 and an image processing unit 9 that read the radiographic image detector 4 on which a radiographic image is recorded. The image signal read by the image reading unit 9 includes an image signal SB read from the surface (lower surface) opposite to the image signal SA read from the surface (upper surface) of the radiation image detector 4 on the radiation source side. The two image signals SA and SB are output to the image output device 10 or the image display device 11 as one image signal SA + SB after appropriate image processing and image superposition by the image processing unit 9.

  First, an embodiment of a phase contrast breast imaging apparatus will be described. FIG. 2 is an explanatory diagram showing the main configuration of the phase contrast mammography apparatus used in the present invention. This phase contrast breast imaging apparatus 1 is a radiographic imaging apparatus for mammography. The phase contrast mammography apparatus 1 is provided with a main body 13 that supports the small focus radiation source 2. The main body 13 includes a subject table 3 that can move in the irradiation direction and supports the subject H, a vertically movable pressing portion 14 that compresses the subject H supported by the subject table 3, and the main body 13. A support shaft 18 that hangs vertically and a power supply unit 20 are provided.

  At the lower end side of the support shaft 18, a radiation image detector 16 that faces the small focus radiation source 2 through the subject table 3 and detects the radiation transmitted through the subject H is provided movably along the support shaft 18. ing. For this reason, since the radiation image detector 16 is movable along the support shaft 18, that is, is movable in the irradiation direction of the small focus radiation source 2, the distance R 2 between the subject H and the radiation image detector 16, the small focus. The distance R1 between the radiation source 2 and the subject table 3 and the distance (R1 + R2) between the small focus radiation source 2 and the radiation source image detector 16 can be changed. The support shaft 18 also functions as an electrical resistor and detects a resistance value corresponding to the position of the radiation image detector 16. The detected resistance value is input to the control unit 6.

  Further, input of shooting conditions and the like is performed by the information input unit 5, and the input information is sent to the control unit 6. The control unit 6 controls the main body unit 13 based on the input information. Further, the distance information 1 regarding the distance between the small focus radiation source 2 and the radiation image detector 16 obtained based on the electrical resistance value inputted from the support shaft 18 and the tube voltage information v obtained based on the input conditions are obtained on both sides of the radiation image. Transmit to the reader 7. Further, information such as photographing conditions input by the information input unit 5 is displayed on the monitor 19.

  Below the subject table 3 and below the radiation image detector 16, radiation dose detectors 15 and 17 (phototimers) for detecting the radiation dose are provided. The radiation dose detectors 15 and 17 include a phosphor sheet (detector) that emits light instantaneously according to the irradiated radiation dose, and a phosphor sheet that includes a light detector that detects emitted light emitted from the phosphor. There are semiconductor detector types with semiconductor detectors that detect the type and radiation. In the case of the phosphor sheet type, it is desirable to cover the entire range of the radiation image detector 4. Here, the radiation dose detection unit 15 provided in the lower part of the subject table 3 can be retracted from the subject table 3. In addition, the structure which provides a radiation dose detection part in the upper part of the to-be-photographed base 3 may be sufficient.

  As the small-focus radiation source 2, an X-ray tube that emits X-rays having a radiation wavelength of around 0.1 to 1 mm is used. This X-ray tube emits X-rays by accelerating electrons generated by thermal excitation at a high voltage and colliding them with a cathode, thereby converting the kinetic energy into radiation energy. When an X-ray image is taken, the acceleration voltage is set as a tube voltage, the amount of generated electrons is set as an intercurrent, and the X-ray emission time is set as an exposure time. An anode made of copper, molybdenum, rhodium, tungsten, or the like is used as an anode with which electrons collide. The emitted X-ray energy spectrum can be changed by changing the type of the anode. When an anode such as copper, molybdenum, rhodium, etc. is used, a relatively low energy line spectrum with a narrow X-ray energy distribution is obtained, which is used for mammography to read a fine light image using its characteristics. It is done. When tungsten is used as the anode, it is used for X-rays of a relatively high energy with a wide spectrum and used for non-destructive inspection of the human chest, abdomen, head, and general industry. The medical or industrial use is characterized by a large X-ray dose. In this case, the anode melts in order to cause a large amount of electrons to collide with the anode at a high speed, and the anode may melt at a high temperature. Avoidance is done. That is, it is common to use a rotating anode. Since the imaging apparatus of this embodiment is an apparatus that is used for medical purposes and is intended for mammography, a rotating anode that uses molybdenum or rhodium as an anode is desirable.

  When using a thermionic X-ray tube, the window through which X-rays are emitted is called a focal point. This is almost a square, and the length of one side is called the focal spot size. There are a pinhole camera method, a slit camera method, a resolving power method, and the like as methods for measuring the focus size, which are described in JIS 4704-1994. In general commercially available thermoelectron X-ray tubes, it is common for manufacturers to measure the respective methods and increase the focal spot size as a product specification. This accuracy is about ± 15%, and this focus size can be understood as the actual focus size of the X-ray tube.

  When the focus size of X-rays is large, the emitted X-ray dose increases, but a so-called penumbra occurs. The penumbra is a phenomenon in which one point on the subject is detected as an image having a certain size on the radiation reading unit due to the size of the focus size, and is so-called blur.

  On the other hand, when an X-ray image is taken at a distance R2 between the subject H and the radiation image detector 16, an edge-enhanced (phase contrast-enhanced) image resulting from X-ray refraction can be obtained. This enhancement effect refracts when X-rays pass through the object, and the X-ray density inside the boundary of the object becomes sparse, and further, the outside of the subject H overlaps with X-rays that do not pass through the subject. Rise. In this way, the edge that is the boundary portion of the subject is enhanced as an image. In order to obtain an edge-enhanced image, it is necessary to set the interval R2 between the radiation image detector 16 and the subject H to a predetermined value or more. However, when R2 is increased, blur due to the penumbra increases. Therefore, the focus size is determined based on the balance between the X-ray dose, penumbra and edge enhancement.

  The focal spot size of the X-ray tube used in the phase contrast breast imaging apparatus 1 of the present invention needs to be 30 μm to 300 μm, and more preferably 50 μm to 200 μm. The smaller the focal spot size, the clearer the composition of the subject can be drawn. However, if the focal spot size is too small, X-rays with sufficient intensity to pass through the human body cannot be obtained. On the other hand, if the focal spot size is too large, the blur due to the penumbra is predominantly less sharp than the edge effect due to the phase contrast. The distance (R1 + R2) from the small focus radiation source 2 to the radiation image detector 16 must be at least 75 cm or more because of the problem of blur due to edge enhancement and penumbra. Further, the relationship between the distance R1 between the small focus radiation source 2 and the subject table 3, the distance R2 between the subject table 3 and the radiation image detector 16, and the distance (R1 + R2) between the small focus radiation source 2 and the radiation image detector 16 is as follows. Small focus radiation source 2 such that 85 cm ≦ R1 + R2 ≦ 200 cm, 50 cm ≦ R1 ≦ 100 cm, 30 cm ≦ R2 ≦ 100 cm, preferably 90 cm ≦ R1 + R2 ≦ 165 cm, 60 cm ≦ R1 ≦ 75 cm, 40 cm ≦ R2 ≦ 90 cm, respectively. It is desirable to set the subject table 3 and the radiation image detector 16.

  Further, according to the present invention, when the radiographic image detector 16 is brought close to the subject table 3 and the distance R2 between the subject H and the radiographic image detector 16 is brought close to 0, an image with only the conventional absorption contrast is taken. (Normal shooting) is also possible. In this embodiment, an X-ray tube having a focal size of 300 μm and an X-ray tube having a focal size of 100 μm are provided as the small-focus radiation source 2 so as to be switchable. When the tube performs phase contrast imaging, an X-ray tube with a focal size of 100 μm is used.

  FIG. 3 is an explanatory diagram showing the main configuration of the phase contrast image photographing apparatus. The phase contrast imaging apparatus of this embodiment is configured for the purpose of imaging medical X-ray images of parts other than the breast, such as the chest and abdomen. The basic configuration is the same as that of the above-described phase contrast mammography apparatus. The main body 27 includes a radiation source 21, a subject position setting tool 22, a radiation image detector 23, a rail 24, and a control unit 25. Since the radiation image detector 23 for detecting the radiation transmitted through the subject H is provided to be movable along the rail 24, the radiation image detector 23 is movable along the rail 24, that is, irradiation of the radiation source 21. Since it is movable in the direction, the distance R2 between the subject H and the radiation image detector 23, the distance R1 between the radiation source 21 and the subject position setting tool 22, and the distance between the radiation source 21 and the radiation source image detector 23 ( R1 + R2) can be varied. The rail 24 also functions as an electric resistor and detects a resistance value corresponding to the position of the radiation image detector 23. The detected resistance value is input to the control unit 25.

  Further, input of shooting conditions and the like is performed by the information input unit 26, and the input information is sent to the control unit 25. The control unit 25 controls the main body unit 27 based on the input information. Further, the radiation image double-side reading device 7 converts the distance information 1 regarding the distance between the radiation source 21 and the radiation image detector 23 obtained based on the electrical resistance value inputted from the rail 24 and the tube voltage information v obtained based on the input conditions. Send to.

  In the phase contrast imaging apparatus of FIG. 3, a rotating anode using tungsten is used for the anode portion in order to use X-rays having high energy and a broad spectrum. In addition, since the phase contrast imaging apparatus of FIG. 3 performs X-ray imaging of a part other than the breast, the interval R1 between the radiation source 21 and the subject position setting tool 22 is 1 m or more, the subject H and the radiation image detector 23. The distance R2 is preferably 1.5 m or more. Further, the relationship between the radiation source 21 and the radiation image detector 23 (R1 + R2) needs to be within 10 m, preferably within 5 m, considering the X-ray dose and the size of the X-ray imaging room. In the phase contrast image photographing device, it is a preferable aspect to perform photographing using a subject position setting tool and a photographing device as disclosed in JP-A-2001-311701.

  In the present invention, photostimulable phosphor plates are used as the radiation image detectors 4, 16, and 23 described above. When the photostimulable phosphor plate is irradiated with radiation (X-rays, α-rays, β-rays, γ-rays, electron beams, ultraviolet rays, etc.), a part of this radiation energy is accumulated, and then excitation light such as visible light is emitted. A photostimulable phosphor that exhibits photostimulated luminescence according to the energy stored upon irradiation is used. Radiation image information of a subject such as a human body is temporarily recorded on a photostimulable phosphor on a sheet, and this stimulable phosphor plate is scanned with excitation light such as laser light to generate photostimulated emission light. The stimulated emission light is photoelectrically read to obtain an image signal.

  FIG. 4 shows a schematic diagram of the photostimulable phosphor plate used in the present invention. The photostimulable phosphor plate 31 includes, as a basic structure, a support 30 and a photostimulable phosphor layer 29 provided thereon. However, when the photostimulable phosphor layer 29 is self-supporting, a support is not necessarily required. Further, a protective layer 28 is provided on the upper surface of the photostimulable phosphor layer 29 (the surface not facing the support 30) to protect the phosphor layer from chemical alteration or physical impact. doing. The support 30 and the protective layer 28 need to be made of a material that transmits the photostimulated luminescence light in order to photoelectrically read the photostimulated luminescence light on both sides of the plate. It is also necessary that the light to be collected be a material that does not totally reflect the surface of the material, and that the surface be smooth. In the photostimulable phosphor plate 31, it is preferable to use a sharpness improving technique as disclosed in JP-A-2000-227500 and JP-A-6-130523.

  In the present invention, a double-sided light condensing type reader is used as the radiation image double-side reader 7 in the system of FIG. In the radiation image recording / reproducing method, reading of radiation image information is generally performed by irradiating excitation light from the front surface side of the radiation image conversion panel, and stimulating light emitted from the phosphor particles on the excitation light irradiation side. The light is collected by a light collecting guide provided in the above, and photoelectrically converted and read (single-sided light collecting method). However, when it is desired to read out as much of the stimulated emission light as possible, or the energy intensity of the radiation energy storage image formed in the photostimulable phosphor layer changes in the depth direction of the phosphor layer, and the energy intensity changes. When it is desired to obtain (intensity distribution) as image information, a method (double-sided condensing method) that collects and reads out the stimulated emission light from both the front and back sides of the panel is also used. This double-sided condensing type reader is described in, for example, Japanese Patent Laid-Open No. 55-87970.

  When this double-sided condensing type reader is used, one radiation image is accumulated and recorded on the photostimulable phosphor plate 31, and the photostimulated luminescent light is collected from both sides of the plate. There is an advantage that the light efficiency is improved and the S / N ratio is improved as compared with the conventional single-side light condensing method.

  FIG. 5 is a diagram showing the concept of the in-device radiation image reading method according to the double-sided condensing method. The photostimulable phosphor plate 31 is conveyed in the direction of the arrow by a conveying means such as a roller. Excitation light such as a laser beam (light having a wavelength range of ± 5% centered on the maximum peak wavelength in the stimulable excitation spectrum of the stimulable phosphor of the radiation image conversion panel to be used is selected as excitation light) 35 Stimulated emission light 36 emitted from the stimulable phosphor plate 31 and emitted from the stimulable phosphor plate 31 by irradiation of the excitation light travels to both surfaces of the stimulable phosphor plate 31. The stimulated emission light 36b that has traveled to the lower surface of the photostimulable phosphor plate 31 is collected by a condensing guide attached to the photoelectric conversion device 34b such as a photomultiplier provided below the photostimulable phosphor plate 31. The light is converted into an electric signal by the photoelectric conversion device 34b, amplified by the amplifier 37b, and finally sent to the image processing unit 9.

  On the other hand, the photostimulated luminescent light 36a that has traveled to the upper surface of the photostimulable phosphor plate 31 is directly by the condensing guide attached to the photoelectric conversion device 34a provided above the photostimulable phosphor plate 31 or the mirror 32. Then, the light is condensed, converted into an electric signal by a photoelectric conversion device 34a such as a photomultiplier, amplified by an amplifier 37a, and finally sent to the image processing unit 9. The image processing unit 9 performs image processing on the electrical signals SA and SB sent from the amplifier 37a and the amplifier 37b, and transmits an image signal corresponding to a target radiation image to the image display device 11 or the image output device 10. .

  When magnified photography is performed to obtain a phase contrast image, the obtained X-ray image is automatically displayed on a hard copy such as a monitor or photographic film by automatically returning it to the actual size (same size) of the subject. Is a preferred embodiment. The image enlargement ratio at the time of X-ray image capturing is calculated by the image processing unit 9 based on the distance information l transmitted from the control unit 6 or the control unit 25 to the image processing unit 9 of the radiographic image double-sided reading device 7. When X-ray image information is sent from the calculated value to the image display device 11 or the image output device 10, interpretation is facilitated by changing or displaying the image magnification rate to the actual size.

  FIG. 6 is a schematic diagram of the overlay processing apparatus. The radiographic image signals SA and SB read from both sides of the photostimulable phosphor plate 31 by the radiographic image reading unit 8 are converted into digital signals by the A / D converters 38 a and 38 b and input to the image processing unit 9. . In the image processing unit 9, the image signals SA and SB are subjected to image processing such as blur mask processing and edge enhancement processing by the image processing means 39 and then input to the weighted overlay processing device 40. In the weighted superposition processing device 40, the weighting factors K1 and K2 are selected from the weighting factor list 41 based on the tube voltage information v and the distance information l input from the control unit 6 of the phase contrast image capturing device, and the image signal SA ′. , SB ′ is weighted, and then the image signal SA + SB is output to the appropriate image output device 10 and image display device 11.

  This radiographic imaging system performs phase contrast radiographic imaging, and reads the photostimulable phosphor plate on which the radiographic image is recorded from both sides of the photostimulable phosphor plate, thereby sharpening the radiographic image, The present invention can be applied to an imaging system that provides a radiation image with improved granularity and high lesion detectability.

It is a block diagram which shows the 1st structural example of a radiographic imaging system. It is explanatory drawing which showed the main structures of the phase contrast mammography apparatus. It is explanatory drawing which showed the main structures of the phase contrast image imaging device. It is the schematic of a photostimulable phosphor plate. It is a figure which shows the concept of the radiographic image reading method in an apparatus according to a double-sided condensing system. It is the schematic of a superposition processing apparatus.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 Phase contrast imaging device 2,21 Small focus radiation source 3 Subject stand 4,16,23 Radiation image detector 5,26 Information input part 6,25 Control part 7 Radiation image double-sided reader 8 Radiation image reading part 9 Image processing Part 10 Image output device 11 Image display device 12 Internet 31 Stimulable phosphor plate

Claims (4)

A phase-contrast breast imaging apparatus that images a breast with radiation from a radiation source and records a radiation image of the breast on a stimulable phosphor plate;
Radiation image double-sided reading device that irradiates the stimulable phosphor plate with excitation light and reads the photostimulated luminescent light on both sides of the photostimulable phosphor plate;
A radiographic imaging system comprising: A phase-contrast image capturing device that images a subject with radiation from a radiation source and records a radiation image of the subject on a stimulable phosphor plate;
Radiation image double-sided reading device that irradiates the stimulable phosphor plate with excitation light and reads the photostimulated luminescent light on both sides of the photostimulable phosphor plate;
A radiographic imaging system comprising: The radiographic image capturing system according to claim 1, wherein the phase contrast radiographic image captured by digital enlargement can be output in an actual size. The image signals SA and SB obtained by reading the radiation image of the same subject from both sides of the photostimulable phosphor plate, the distance information between the radiation source and the photostimulable phosphor plate at the time of imaging, and the radiation source The radiographic imaging system according to any one of claims 1 to 3, wherein weighting is performed by a weighting coefficient obtained based on the tube voltage information.

Images