JP2006026425A - Digital phase contrasted x-ray imaging system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To obtain phase contrasted X-ray digital images by using digital X-ray image detectors such as a computed radiography (CR), a flat panel X-ray image detector (FPD) and the like. <P>SOLUTION: The pixel size of the digital X-ray image detector 2 is 0.5 or more times and 3 or less times of EB shown as the following simulation formula (1) or EP shown as the following simulation formula (2). EB=2.3(1+R2/R1)<SP>1/3</SP>äR2δ(2r)<SP>1/2</SP>}<SP>2/3</SP>+D×(R2/R1) (1). EP=2.3äR2δ(2r)<SP>1/2</SP>}<SP>2/3</SP>(2). In the above formulas (1) and (2), R1 is shown as the distance to the center of a cylinder when assuming the X-ray source 5 and the subject 1 to be cylinders, and R2 is shown as the distance from the center of the cylinder to the digital X-ray image detector when assuming the subject to be the cylinder, and r is shown as the radius of the cylinder when assuming the subject to be the cylinder, and δ is shown as the coefficient defined by each subject, and D is shown as the focal size of the X-ray source. <P>COPYRIGHT: (C)2006,JPO&NCIPI

Description

  The present invention relates to an X-ray imaging system for use in nondestructive inspection and medical image diagnosis, and relates to a digital phase contrast X-ray imaging system capable of obtaining a high-quality X-ray image with good sharpness.

  X-rays pass through a substance, and when passing through a substance, the X-rays are absorbed by photoelectric effect and Compton scattering by atoms constituting the substance. The amount of X-ray absorption depends on the atomic weight. Therefore, an X-ray image can be obtained by obtaining an X-ray intensity distribution two-dimensionally according to the distribution of atoms after X-rays pass through the object. The X-ray intensity difference generated as the image density obtained here is generally called density contrast. This is the principle that has been used as an X-ray image since the discovery of X-rays by Dr. Roentgen in 1895.

  On the other hand, in the 1990s, research on phase contrast X-ray images began to be widely conducted. Since X-rays are electromagnetic waves, they have the same properties as light. That is, refraction and interference occur. The X-ray intensity difference obtained by this property is called phase contrast. It has been found that the edge of a conventional absorption contrast image is sharply depicted by the phase contrast. However, since the refractive index of X-rays is extremely small compared to visible light, no phase contrast image was observed with the conventional X-ray imaging method. Since the wavelength of X-rays is as small as three orders of magnitude of visible light, it is very difficult to obtain a phase contrast image by controlling the interference, and it is currently in the research stage and not at a stage that can be widely used in hospitals.

  There has been reported a method for obtaining a clear X-ray image by obtaining a so-called edge-enhanced image that enhances the edge of an image by utilizing the property of X-ray refraction. That is, those using synchrotron X-rays are described in, for example, N.I. The magazine Medical Physics Vol. 26, No. 10, page 2190 (1999) by Yagi et al. And the one using a Coolidge X-ray tube (thermo-electron X-ray tube) are described in A.C. Journal of Optical Review Vol. 7, No. 6, page 566 (2000) by Ishisaka et al. Regarding a method for obtaining an edge-enhanced phase-contrast X-ray image using X-ray interference, a method using a special Coolidge X-ray tube with a small focal point size is described in S.A. W. Wilkins et al. Discloses a magazine Nature 384, 335 (1996), WO96 / 31098, and an application of the principle of the interferometer using the above-mentioned synchrotron X-ray, in JP-A-9-187455. ing.

  The present invention relates to an X-ray imaging system that uses X-ray refraction to obtain an X-ray image with excellent sharpness. In order to capture a clear X-ray image due to the refraction, a non-screen X-ray film, an X-ray intensifying screen and an X-ray film, or a very small CCD camera having a pixel size of about 10 μm have been used. It was. This is because it has been considered that a very high resolution is required to image edge enhancement due to X-ray refraction.

  Here, when using a non-screen X-ray film or an X-ray intensifying screen and an X-ray film, the resolution is very high. On the other hand, these are so-called analog X-ray image detectors, and their image processing and output image are performed. Cannot freely enlarge / reduce.

  In order to obtain a digital X-ray image, an X-ray detector can be used using a high-resolution CCD. However, in order to maintain sufficient sensitivity, the cooling of the CCD part and the apparatus become very expensive, and partly It is only used for special research.

  The present invention provides a digital phase contrast X-ray imaging system that can easily obtain a digital image of a phase contrast X-ray image. That is, sharpness can be easily achieved by using a digital X-ray image detector such as a computed radiography (CR) or a planar X-ray image detector (FPD) widely used in the medical field or nondestructive inspection. Provided is a digital phase contrast X-ray imaging system that obtains an excellent digital X-ray phase contrast image.

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

The invention according to claim 1 includes a digital X-ray image detector that obtains a digital image of a phase contrast X-ray image irradiated from an X-ray source and transmitted through a subject,
The pixel size of the digital X-ray image detector is not less than 0.5 times and not more than 3 times the EB represented by the following simulation formula (1) or the EP represented by the following simulation formula (2). A digital phase contrast X-ray imaging system.
EB = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3 + D × (R2 / R1)
(1)
EP = 2.3 {R2δ (2r) ^1/2 } ^2/3 (2)
In the above equations (1) and (2), R1 represents the distance to the center of the cylinder when the X-ray source and the object are assumed to be a cylinder, and R2 represents the cylinder when the object is assumed to be a cylinder. This represents the distance from the center to the digital X-ray image detector, r represents the radius of the cylinder when the object is assumed to be a cylinder, δ represents a coefficient determined for each object, and D represents the focal point size of the X-ray source. To express. ].

  The invention according to claim 2 is as follows: “The X-ray source is a Coolidge X-ray tube. The digital phase contrast X-ray imaging system according to claim 1, wherein: ].

  According to a third aspect of the present invention, the pixel size of the digital X-ray image detector is 0.5 to 3 times the EB. Digital phase contrast X-ray imaging system. ].

  The invention according to claim 4 is as follows: “E obtained by the following simulation formula (3) is not less than 12 μm and not more than 300 μm. The digital phase contrast X-ray imaging system according to claim 3.

E = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3 (3) ”.

  5. The digital phase contrast X-ray imaging system according to claim 1, wherein the X-ray source emits synchrotron radiation X-rays. ].

  The invention according to claim 6 is: "Digital phase contrast X according to claim 5, wherein the pixel size of the digital X-ray image detector is not less than 0.5 times and not more than 3 times the EP." Line imaging system. ].

  7. The digital phase contrast X-ray imaging system according to any one of claims 1 to 6, wherein the pixel size is 20 μm or more and 300 μm or less. ].

  The invention according to claim 8 is as follows: “The detection surface of the digital X-ray image detector has a size of 15 cm × 15 cm or more and 90 cm × 90 cm or less. 2. The digital phase contrast X-ray imaging system according to item 1. ].

  According to the first to eighth aspects of the present invention, it is possible to easily obtain a digital image of a phase contrast X-ray image, that is, a computed radiography widely used in a medical field or a nondestructive inspection. A digital X-ray phase contrast image having excellent sharpness can be easily obtained by using a digital X-ray image detector such as (CR) or a planar X-ray image detector (FPD).

  The configuration of the present invention will be described in detail below.

  The principle that phase contrast edge enhancement of a subject image occurs due to refraction of X-rays is shown in FIG. Since the refractive index of X-rays is smaller than 1, when the object 1 is a cylindrical or spherical object, the cylindrical or spherical object acts like a concave lens. That is, X-rays that have passed through the edge of the object overlap with X-rays that have passed beside the object on the digital X-ray image detector 2, and the X-ray intensity is increased. Since the X-rays pass through the object edge and protrude outside the object, the X-ray intensity is weak in the vicinity of the object edge. In this way, the X-ray intensity has a mountain on the outside and a valley on the inside with the object edge as a boundary. Such edge enhancement is sometimes called an edge effect. When phase contrast edge enhancement due to X-ray refraction occurs in this way, a sharp X-ray image with sharp edges can be obtained.

At this time, if the X-ray source is regarded as a point light source, as shown in FIG. 2, the half width E of phase contrast edge enhancement can be expressed by the following equation.
E = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3 (1)
Here, R1 is the distance from the X-ray point source 3 to the center of the cylinder of the subject 1, R2 is the distance from the center of the cylinder assumed as the subject 1 to the digital X-ray image detector 2, and r is the subject 1 as a cylinder. It corresponds to the radius when Since the human body is an X-ray image constituent material mainly composed of blood vessels and bones, the shape may be approximated to a cylinder. For δ, the refractive index n of X-rays is expressed as follows: n = 1−δ (2)
δ relates to the phase change.

  Here, in medical sites and non-destructive inspection facilities, the Coolidge X-ray tube 5 (also referred to as a thermoelectron X-ray tube) is widely used. The case where the Coolidge X-ray tube 5 is used is shown in FIG. In the Coolidge X-ray tube 5, thermoelectrons collide with a metal anode such as tungsten to emit X-rays, and X-rays are emitted radially from a substantially square window called a focal point. The length of one side of this almost square focus is called the focus size. When the Coolidge X-ray tube 5 is used, the X-ray source cannot be regarded as an ideal point light source. That is, the focus as an X-ray source having a finite size widens the half-width of phase contrast edge enhancement and reduces the intensity due to so-called geometrical sharpness. At this time, the half-value width of the phase contrast edge enhancement can be expressed by the following equation (3).

EB = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3 + D × (R2 / R1) (3)
Here, D represents the focal spot size of the Coolidge X-ray tube 5 to be used.

  In other words, this geometrical sharpness widens the half width of the phase contrast edge enhancement, so that even if the pixel size of the digital X-ray image detector 2 has a constant size, it can be detected as shown in FIG. Can be expanded in size. This is the basic principle of the present invention when the Coolidge X-ray tube 5 is used.

  Next, a powerful X-ray beam that is very close to parallel can be obtained with Spring-8 in Harima, Hyogo Prefecture, and the synchrotron radiation X-ray apparatus at the High Energy Laboratory in Tsukuba. Parallel X-rays can be regarded as moving a point light source or a finite-size X-ray focal point to infinity. In other words, when R1 is infinite in Formula 1 or Formula 3, the half-value width EP of phase contrast edge emphasis when using parallel X-rays can be expressed as Formula 4.

EP = 2.3 × {R2δ (2r) ^1/2 } ^2/3 (4)
In order to expand the EP so that phase contrast edge enhancement can be observed by the digital X-ray image detector 2 having a fixed-size pixel with a parallel X-ray source, as shown in FIG. 5, it is realized by increasing R2. When using parallel X-rays, the X-ray intensity does not decrease even if the digital X-ray image detector 2 is moved away from the subject 1. Therefore, by setting R2 having a sufficient length, it is possible to observe phase contrast edge enhancement even in the digital X-ray image detector 2 having a large pixel.

  In digital X-ray image capturing, the two-dimensional planar digital X-ray image detector 2 is constituted by so-called pixels of a square or a rectangle with sides of about 10 μm to 300 μm. Each pixel represents the minimum unit of the X-ray image (FIG. 6).

  The pixel size of the digital X-ray image detector 2 is preferably 20 μm or more and 300 μm or less, and more preferably 50 μm or more and 200 μm or less. The size of the detection surface of the digital X-ray image detector 2 is preferably 15 cm × 15 cm or more and 90 cm × 90 cm or less, which can be preferably used in medicine.

  Here, when the pixel size is larger than the EB or EP, it is possible that peaks and valleys due to the edge effect cancel each other and phase contrast edge enhancement disappears. Therefore, in order to detect the phase contrast edge enhancement by the digital X-ray image detector 2, it is desirable that the peaks and valleys of the phase contrast edge enhancement are separated by more than the pixel size. Alternatively, even when the pixel size is larger than the above EB or EP, the phase contrast edge enhancement is detected because the peaks or valleys of the phase contrast edge enhancement are in different pixels. When the pixel size is larger than the distance between the peaks and valleys, phase contrast edge enhancement can be detected by capturing within one pixel in some cases and canceling with another pixel in other cases. In other words, phase contrast edge enhancement is detected in an established manner, and phase contrast edge enhancement is more easily detected as the distance between the peaks and valleys is larger than the pixel size.

  When the Coolidge X-ray tube 5 is used, the gap between the peak and valley of the phase contrast edge emphasis is widened due to geometrical sharpness, while the difference between the peak height and the valley to the bottom, that is, the intensity of the phase contrast edge emphasis is low. Become. Thus, when the distance between the peaks and valleys is increased due to geometrical sharpness, the phase contrast edge enhancement is detected most strongly when the distance between the peaks and valleys approximately matches the pixel size.

  On the other hand, when parallel X-rays are used, the edge width increases as R2 increases, so the X-ray intensity detected by the digital X-ray image detector 2 is such that the distance between the peaks and valleys of phase contrast edge enhancement is greater than or equal to the pixel size. It can be detected strongly.

  Thus, the strength of phase contrast edge enhancement is determined by the relative relationship between the peak-to-valley distance of phase contrast edge enhancement and the pixel size of the digital X-ray image detector 2, and varies continuously. Actually, if the distance between the peak and the valley is 0.5 times or more with respect to the pixel size, it is a practical area. This is because other factors such as blur due to the detector itself exist in addition to the focal diameter of the X-ray tube as a factor that blurs and spreads the phase contrast edge enhancement. In the case of the Coolidge X-ray tube 5, phase contrast edge enhancement is clearly observed even when the X-ray intensity decreases until the distance between the peak and the valley is about three times the pixel size. If the edge width becomes too wide, the image will be difficult to see, and more preferably about 1.5 times.

  On the other hand, in the case of parallel X-rays, the preferable condition is that the distance between the peaks and valleys is about three times the pixel size, more preferably about 1.5 times. That is, “the pixel size of the digital X-ray image detector is substantially equal to the half-width of the phase contrast edge enhancement of the phase contrast X-ray image” refers to a pixel size of about 0.5 to 3 times EB or EP. The EB or EP here is, for example, about 1 mm of the blood vessel size to be observed in the case of a medical image, or 0.1 mm of microcalcification at the time of mammography. In the case of non-destructive inspection, it is a metal wire having a width of several μm, and can be obtained by calculation using formulas 3 and 4.

  As described above, when the Coolidge X-ray tube is used, the gap between the peak and valley of the phase contrast edge emphasis is widened due to geometrical sharpness, while the difference between the peak height and the valley and the bottom, that is, the phase contrast edge. The intensity of emphasis is low. Therefore, it cannot be detected unless the original phase contrast edge enhancement is sufficient. That is, it is preferable that the phase contrast edge emphasis E1 before the geometrical unsharpness expressed by Formula 5 is 12 μm or more and 300 μm or less.

E = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3 (5)
On the other hand, in the case of parallel X-rays, there is no reduction in phase contrast edge enhancement due to geometrical insensitivity, so the above limitation does not occur.

  Based on the above configuration, a digital image of a phase contrast X-ray image can be easily obtained. That is, phase contrast with excellent sharpness using digital X-ray image detectors such as computed radiography (CR) and planar X-ray image detector (FPD) widely used in the medical field and non-destructive inspection. A digital image of an X-ray image can be obtained.

  That is, the phase contrast edge is obtained by the “digital phase contrast X-ray imaging system characterized in that the pixel size of the digital X-ray image detector is substantially equal to the half-width of the phase contrast edge enhancement of the phase contrast X-ray image”. An enhanced digital X-ray image with excellent sharpness can be obtained.

  Furthermore, by the “digital phase contrast X-ray imaging system of the present invention characterized in that the X-ray tube to be used is a Coolidge X-ray tube”, a medical diagnostic image having excellent sharpness can be obtained in a general hospital facility or the like. it can.

  When a Coolidge X-ray tube is used, “digital phase contrast X-ray imaging characterized by using a digital X-ray image detector having a pixel size 0.5 to 1.5 times the EB value obtained by the following simulation formula The “system” makes it possible to easily obtain medical diagnostic images with excellent sharpness in general hospital facilities.

EB = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3 + D × (R2 / R1)
In addition, an “X-ray image clearly depicting a fine structure can be obtained by the“ digital phase contrast X-ray imaging system of the present invention characterized in that the X-ray used is a synchrotron radiation X-ray ”.

In the case of using parallel X-rays, “digital phase contrast X-ray imaging characterized by using a digital X-ray image detector having a pixel size 0.5 to 1.5 times the EP value obtained by the following simulation formula. With the “system”, an X-ray image that clearly depicts a fine structure can be obtained.

EP = 2.3 {R2δ (2r) ^1/2 } ^{2/3 In} addition, “E value obtained by the following simulation formula is 12 μm.
A medical diagnostic image having excellent sharpness can be obtained by a digital phase contrast X-ray imaging system characterized by being m to 300 μm.

E = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3

Hereinafter, although the example of embodiment of the digital phase contrast X-ray imaging system of this invention is shown, this invention is not limited to this.

  As shown in FIGS. 1 to 6, the digital phase contrast X-ray imaging system according to the present invention includes an X-ray source such as a synchrotron X-ray source and a Coolidge X-ray tube, and two X-ray images such as CR and FPD. A digital X-ray image detector such as a two-dimensional planar image detector. The digital X-ray image obtained here is appropriately subjected to image processing 10 and CAD 11 as shown in FIG. 7, and is printed on the image print output 12 as a transmission image X-ray film by a laser imager or the like. Further, the image is displayed on an image display 13 such as a cathode ray tube or a liquid crystal display, or an image signal storage 14 is performed and used for purposes such as image diagnosis.

  In this digital phase contrast X-ray imaging system, since a digital X-ray image can be obtained, the image can be arbitrarily reduced or enlarged, and gradation processing and frequency processing of the output image can be performed. . Also, as shown in FIG. 7, image transfer to a remote place can be easily performed using the Internet 15 or the like.

  The reason why the Coolidge X-ray tube is also called a thermoelectron X-ray tube is an X-ray generation mechanism in which thermoelectrons are emitted from a heated filament and collide with a metal anode to generate X-rays having energy equivalent to an acceleration voltage. For the metal anode, molybdenum, rhodium, tungsten, copper, silver or the like is used. In order to eliminate inconveniences such as melting of the anode metal due to intense collision of thermoelectrons, the rotating anode thermoelectron X-ray tube that rotates the disk-shaped anode is widely used in medical fields Has been. In the present invention, a fixed anode X-ray tube and a rotary anode X-ray tube can be used. The type of metal of the anode and the acceleration voltage are not particularly limited. In general, anodes such as molybdenum, rhodium, and tungsten are used for nondestructive inspection and medical diagnostic imaging. The voltage for accelerating the thermoelectrons is preferably in the range of 10 kV to 200 kV. Particularly in medical image diagnosis, 20 kV to 150 kV is preferable. Since the set voltage is the highest kinetic energy component of X-rays, the X-rays of the set voltages are called 20 kVp (kilovolt peak) or 150 kVp X-rays.

  When using a thermionic X-ray tube, the window through which X-rays are emitted is called a focal point. It is almost 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.

  The focal spot size of the X-ray tube used in the present invention is preferably 10 μm to 500 μm, 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, geometrical sharpness increases and the edge effect disappears.

  The synchrotron X-ray is obtained by a so-called synchrotron. This accelerates the electrons to near the speed of light, and when the electrons are moved in an arc shape, powerful X-rays having a very small radiation angle in the tangential direction can be obtained. For example, the radiation angle is 100 μm radians or less, and the radiation angle varies depending on the speed of the accelerated electrons, such as 10 μm radians or several μm radians. The characteristics of synchrotron radiation X-rays are that the radiation angle is small, that is, the parallelism is high, and that the intensity is so high that monochromatic X-rays with sufficient intensity can be extracted using the Bragg reflection of silicon crystal. For example, high-energy monochromatic X-rays such as 17 keV and 50 keV that can be sufficiently transmitted through the human body can be obtained, and a clear image that can be used for medical image diagnosis with strong phase contrast edge enhancement due to refraction can be obtained. it can.

  Synchrotron X-rays are, for example, the synchrotron of the High Energy Research Institute in Tsukuba, Ibaraki, SPring-8 of Harima, Hyogo, and the microtron disclosed in Japanese Patent Application 2000-366836. Since the X-ray obtained here has extremely good parallelism, even if R2 is increased, the X-ray does not spread, so that the X-ray intensity does not decrease, and therefore, the half width EP of the edge effect is sufficiently wide. Can do.

  When using a synchrotron radiation X-ray source, the distance between the X-ray source and the subject need not be determined. However, a place where the subject is sufficiently safe is preferable. When a thermionic X-ray tube is used, the distance (R1) between the X-ray tube and the subject is about 0.1 m to 2 m. Preferably, it is about 0.3 m to 1 m.

  The distance (R2) from the subject to the digital X-ray image detector is preferably 0.15 m to 50 m when using synchrotron X-rays. If it is too close, a sufficient phase contrast edge enhancement effect cannot be obtained. If it is too far away, it is subject to the physical constraints of the room where the image is taken. When using a thermionic X-ray tube, the distance (R2) from the subject to the digital X-ray image detector is preferably 0.15 or more and 5 m. The reason for taking this range is the same as above.

  The minimum reading size of the subject in the present invention, that is, the minimum size of how much subject information is to be read, is about 30 μm to 10 mm as the diameter of the cylinder. For non-destructive inspection, about 30 μm is required. In the case of a mammogram, about 100 μm is required. Further, in the chest image or the like, about 1 mm to 5 mm is required. It can be set according to the purpose.

The subject is a metal such as iron or a human body. At this time, the value of δ is 10 ⁻⁸ to 10 ⁻⁶ .

The two-dimensional planar digital X-ray image detector 2 used in the present invention is a solid-state imaging such as computed radiography: CR, and flat panel detector: FPD (direct method, indirect method) using a stimulable phosphor plate. There are an element, a phosphor (Gd ₂ O ₂ S: Tb, CsI, etc.), a lens (or taper), and a CCD.

  When these digital X-ray image detectors 2 are used, since the phase contrast image is enlarged, it corresponds to the pixel size of the digital X-ray image detector 2 being reduced. In other words, this corresponds to high-definition reading and has the advantage of increasing image information.

  When the stimulable phosphor plate is irradiated with radiation (X-rays, α-rays, β-rays, γ-rays, electron beams, ultraviolet rays, etc.), a part of the radiation energy is accumulated, and then irradiated with excitation light such as visible light. Then, using a stimulable phosphor (stimulable phosphor) that exhibits stimulating luminescence according to the stored energy, radiation image information of a subject such as a human body is once recorded on the stimulable phosphor on the sheet, The stimulable phosphor sheet is scanned with excitation light such as laser light to generate stimulated emission light, and the obtained stimulated emission light is photoelectrically read to obtain an image signal. A radiation image can be output as a visible image on a recording material such as a photographic material, a CRT, etc. (Japanese Patent Laid-Open Nos. 55-124929, 56-163472, 56-104645, 55-116340, etc.) ).

  Further, as a solid-state imaging device such as a flat panel detector, for example, as described in JP-A-6-342098, a photoconductive layer that generates charges according to the intensity of irradiated X-rays is generated. A method is used in which the accumulated charges are stored in a plurality of capacitors arranged two-dimensionally. Further, as described in Japanese Patent Application Laid-Open No. 9-90045, a photodiode or the like in which X-rays are absorbed by a phosphor layer such as an intensifying screen to emit fluorescence and the intensity of the fluorescence is provided for each pixel. A method of detecting with a photodetector is also used.

  As for the solid-state imaging device such as the flat panel detector, an organic flat panel detector in which a photodetector such as a photodiode or a switching element such as a TFT is manufactured using an organic semiconductor can also be used.

  Here, the configuration of the imaging panel 241 provided in the flat panel detector 240 will be described with reference to FIG. The imaging panel 241 has a substrate having a thickness sufficient to obtain a predetermined rigidity, and a detection element 2413- (1,1) that outputs an electrical signal on the substrate in accordance with the dose of irradiated radiation. ) To 2413- (m, n) are two-dimensionally arranged in a matrix. Further, the scanning lines 2411-1 to 2411-m and the signal lines 2412-1 to 2412-n are arranged so as to be orthogonal to each other, for example.

  The scanning lines 2411-1 to 2411 -m of the imaging panel 241 are connected to the scanning driving unit 2414. When the readout signal RS is supplied from the scan driver 2414 to one of the scanning lines 2411-1 to 2411-m, the scanning line 2411-p (p is any value of 1 to m), the scanning line 2411. Electric signals SV-l to SV-n corresponding to the radiation dose emitted from the detection element connected to −p are output, and are sent to the imaging data generation circuit 2415 via the signal lines 2412-l to 2412-n. Supplied.

  The detection element 2413 may be any element that outputs an electrical signal corresponding to the dose of irradiated radiation. For example, when a detection element is formed using a photoconductive layer in which electron-hole pairs are generated and change in resistance when irradiated with radiation, the detection element is in accordance with the amount of radiation generated in the photoconductive layer. An amount of charge is stored in the charge storage capacitor, and the charge stored in the charge storage capacitor is supplied to the imaging data generation circuit 2415 as an electrical signal. It is desirable that the photoconductive layer has the highest dark resistance value, such as amorphous selenium, lead oxide, cadmium sulfide, mercuric iodide, or an organic material exhibiting photoconductivity (light added with an X-ray absorption compound). Including conductive polymer), and amorphous selenium is particularly desirable.

  Further, when the detection element 2413 is formed by using, for example, a scintillator that generates fluorescence when irradiated with radiation, an electrical signal based on the fluorescence intensity generated by the scintillator is generated by a photodiode to generate imaging data. It may be supplied to the circuit 2415.

  The photographing data generation circuit 2415 sequentially selects the electrical signal SV supplied based on an output control signal SC from a reading circuit 242 described later, and converts it into digital photographing data DT. The photographing data DT is supplied to the reading circuit 242.

  The reading circuit 242 is connected to the controller 210, and generates the scanning control signal RC and the output control signal SC based on the control signal CTD supplied from the controller 210. The scan control signal RC is supplied to the scan driver 2414, and the read signal RS is supplied to the scan lines 2411-l to 2411-m based on the scan control signal RC.

  Further, the output control signal SC is supplied to the photographing data generation circuit 2415. For example, when the imaging panel 41 includes (m × n) detection elements 2413 as described above by the scanning control signal RC and the output control signal SC from the reading circuit 242, the detection element 2413- ( If the data based on the electrical signal SV from l, l) to 2413 (m, n) is data DP (l, l) to DP (m, n), data DP (1, 1), DP (1, 2 ),... DP (1, n), DP (2,1),..., DP (m, n) are generated in this order, and this shooting data DT is generated as a shooting data generation circuit. 2415 is supplied to the reading circuit 242. In addition, the reading circuit 242 also performs processing for sending the photographing data DT to the controller 210.

  Shooting data DT obtained by the flat panel detector 240 is supplied to the controller 210 via the reading circuit 242. If the image data obtained by logarithmic conversion processing is supplied when the image data obtained by the radiation image reader 240 is supplied to the controller 210, the processing of the radiation image data in the controller 210 can be simplified.

  Next, FIG. 9 shows a mechanical part of a configuration example of a flat panel detector 240 using a stimulable phosphor plate. First, the flat panel detector 240 will be described. The photostimulable phosphor plate 241B is fixed to the left side wall and is used repeatedly. The reading unit 243 is composed of a stepping motor or the like, and moves along the guide shaft 244B by driving the ball screw 244A by the sub-scanning motor 244M, and scans the scanning line (light beam) 245 in the sub-scanning direction.

  Scanning in the scanning direction is performed by the polygon scanning mechanism 243A. Polygon scanning mechanism 243A includes a polygon and a mechanism for rotating the polygon. The operation of the sub-scanning motor 244M is controlled by the sub-scanning motor control mechanism 244C. The fluorescence is condensed by a condenser 243B and converted into an electric signal by a photomultiplier 243C.

  LD1 is a laser light source, PD1 is a photo sensor, and constitutes an origin position detection sensor. This origin position detection sensor detects the origin position of the reading unit 243 in the sub-scanning direction. The output of the photosensor PD1 is input to the sub-scanning motor control mechanism 244C, and the sub-scanning motor control mechanism 244C controls the stop position of the reading unit 243.

  In this example, the reading unit 243 is moved by driving the ball screw 244A, but the stimulable phosphor plate 241B may be moved in the sub-scanning direction.

  As another fluorescence detection means, there is a method using a CCD or a C-MOS sensor. A configuration in which an X-ray scintillator that emits visible light by X-ray irradiation, a lens array, and an area sensor corresponding to each lens is also used.

  Further, for example, radiographic images are usually taken using an X-ray film in a mass examination, and a laser digitizer is used to input these X-ray photographs to the system of this embodiment. This can be handled as digital image data by scanning the film with a laser beam, measuring the amount of transmitted light, and converting the value into analog to digital. The pixel size at this time corresponds to the sampling pitch of the laser digitizer.

  When a digital X-ray image is obtained with the various configurations described above, the effective pixel size of the image is preferably 200 μm or less, particularly 100 mm or less for a mammogram, depending on the imaging region and the diagnostic purpose. It is preferable. Further, the number of gradations of the image is preferably 10 bits or more, and particularly preferably 12 bits or more.

  In the digital X-ray image detector 2, in the case of CR, the minimum reading size of the imaging plate is the pixel size of the digital image, and is preferably 30 μm or more and 300 μm or less. If it is smaller than 30 μm, the number of pixels becomes enormous, which hinders rapid image processing and image display. The thickness is preferably 50 to 200 μm. In the case of FPD, the minimum size for detecting an X-ray image is the pixel size, and similarly to CR, 30 μm or more and 300 μm are similarly preferable ranges.

  The optimum pixel size varies depending on the subject. In the case of performing double-magnification imaging, a pixel size of about 100 μm is appropriate for a part including fine components such as a hand bone. Even if the pixel size becomes smaller than that, the amount of information hardly changes, and when the pixel size becomes larger than that, the image information deteriorates. A chest image or the like has fewer fine components than a hand bone image or the like, and therefore there is no deterioration of image information until the pixel size is about 200 μm. On the other hand, since a breast image or the like requires very fine information such as the shape of microcalcification, a pixel size of 100 μm or less is necessary, and a sufficiently good image can be obtained at about 50 μm.

  The resulting digital image signal of the phase contrast edge enhanced image can be displayed on a monitor after appropriate image processing, or a hard copy can be obtained with a printer. For example, in the case of medical use, it is very important to observe the actual size of the object to be observed. Therefore, it is a preferable aspect to display the actual size on the monitor or on the hard copy. On the other hand, it is also a preferable aspect to enlarge the image to an arbitrary size in order to examine the image. In addition, it is a preferable aspect that an image is stored as an electronic signal after being used for image diagnosis or the like.

  The digital phase contrast X-ray imaging system according to the present invention can be used for medical image diagnosis, medical specimen image diagnosis, industrial CI chip inspection, and the like.

  1. Simulation calculation when using Coolidge X-ray tube A 1 mm diameter plastic fiber is the subject. If the energy of the X-ray is 50 keV, δ = 8 × 10 −7. The edge profile on the digital X-ray image detector when R1 = 1 m and R2 = 1 m and the Coolidge X-ray tube was used was calculated using three equations. FIG. 10 shows the result of changing the focal spot size of the X-ray tube from 20 μm to 150 μm. As the focal spot size increases, the edge width at which the edge strength decreases becomes wider. In the 87.5 μm pixel size, when the focus size is 100 μm, the peaks and valleys of the phase contrast edge enhancement are outside the pixels, respectively. That is, when an X-ray tube having a focal size of 100 μm is used, it is understood that the phase contrast edge enhancement is obtained most strongly in the simulation calculation here.

  2. Results of photographing experiment using Coolidge tube An X-ray tube L6622-02 manufactured by Hamamatsu Photonics, whose anode is tungsten, was used. The focal spot size was 100 μm and the tube voltage was set to 50 keV. A polyester base plate having a thickness of 200 μm was used as a holder, and a cylindrical resin having a diameter of 1 mm was used as a subject. As the digital X-ray image detector, Konica REGIUS plate RP-1S (35 cm × 43 cm), which is a photostimulable plate, was used. Photographing was set at distances R1 and R2 as shown in the table. The photographing condition was 10 mAs when R1 = R2 = 1 m. When the distance was changed, the X-ray dose was adjusted so that the X-ray dose applied to the subject was constant, and an X-ray image was taken. The image information was read using a REGIUS150 manufactured by Konica for each photoshoot.

  The reading size at this time was 87.5 μm. The read image data was printed on a silver halide photographic film for recording with a Konica laser imager Li62P. After the photographic film was developed, the image was observed on a 8000 lx Schaukasten. The case where the edge enhancement was not visible at the edge of the subject, the case where it was too strong and uncomfortable, was marked as “X”, the case where it was visible as △, and the case where it looked good as ◯. The results are shown in Table 1.

  3. Simulation calculation using parallel X-rays A plastic fiber having a diameter of 1 mm is used as a subject. If the energy of the X-ray is 50 keV, δ = 8 × 10 −7. R2 for sufficiently observing the phase contrast edge enhancement is obtained using a CR of 87.5 μm pixel size. In the case of parallel X-rays, there is no reduction in edge strength due to geometrical sharpness of the focal spot diameter. Therefore, when the value of EP is 87.5 μm and the corresponding R2 is determined, 9.3 m is obtained. From this calculation, when the imaging plate is separated from the subject by about 5 m or more, the edge of the 1 mm fiber starts to be observed, and the edge effect due to the phase contrast is clearly obtained at 10 m or more.

  The present invention can be applied to a digital phase-contrast X-ray imaging system that can be used for nondestructive inspection and medical image diagnosis to obtain high-definition, high-quality X-ray images. That is, using a digital X-ray image detector such as a computed radiography (CR) and a planar X-ray image detector (FPD) widely used in the medical field or non-destructive inspection, A digital X-ray phase contrast image having excellent sharpness can be easily obtained.

It is a figure explaining the principle which phase contrast edge emphasis of a to-be-photographed image arises by refraction | bending of a X-ray. It is a figure which shows the half value width of phase contrast edge emphasis. It is a figure which shows the half value width of phase contrast edge emphasis in the case of using a Coolidge X-ray tube. It is a figure explaining that it can detect even if the pixel size of a digital X-ray image detector has a fixed size. It is a figure explaining extending EP so that phase contrast edge emphasis can be observed with a digital X-ray image detector of a fixed size pixel with a parallel X-ray source. It is a figure explaining comprising by what is called a square or rectangular so-called pixel of one side of a digital X-ray image detector. It is a figure explaining the output utilization of a digital phase contrast X-ray imaging system. It is a figure explaining the structure of the imaging panel provided in the flat panel detector. It is a figure which shows the mechanical part of the structural example of the flat panel detector using a stimulable fluorescent substance plate. It is a figure which shows the result of having changed the focus size of X-ray tube from 20 micrometers to 150 micrometers.

Explanation of symbols

1 Subject 2 Digital X-ray image detector 3 X-ray point light source 5 Coolidge X-ray tube

Claims (8)

A digital X-ray image detector for obtaining a digital image of a phase contrast X-ray image irradiated from an X-ray source and transmitted through a subject;
The pixel size of the digital X-ray image detector is not less than 0.5 times and not more than 3 times the EB represented by the following simulation formula (1) or the EP represented by the following simulation formula (2). A digital phase contrast X-ray imaging system.
EB = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3 + D × (R2 / R1) (1)
EP = 2.3 {R2δ (2r) ^1/2 } ^2/3 (2)
In the above equations (1) and (2), R1 represents the distance to the center of the cylinder when the X-ray source and the object are assumed to be a cylinder, and R2 represents the cylinder when the object is assumed to be a cylinder. This represents the distance from the center to the digital X-ray image detector, r represents the radius of the cylinder when the object is assumed to be a cylinder, δ represents a coefficient determined for each object, and D represents the focal point size of the X-ray source. To express.   The digital phase contrast X-ray imaging system according to claim 1, wherein the X-ray source is a Coolidge X-ray tube.   3. The digital phase contrast X-ray imaging system according to claim 1, wherein a pixel size of the digital X-ray image detector is not less than 0.5 times and not more than 3 times the EB. 4. The digital phase contrast X-ray imaging system according to claim 3, wherein E obtained by the following simulation formula (3) is not less than 12 μm and not more than 300 μm.
E = 2.3 (1 + R2 / R1) ^1/3 {R2δ (2r) ^1/2 } ^2/3 (3)   The digital phase contrast X-ray imaging system according to claim 1, wherein the X-ray source emits synchrotron radiation X-rays.   6. The digital phase contrast X-ray imaging system according to claim 5, wherein a pixel size of the digital X-ray image detector is 0.5 to 3 times the EP.   The digital phase contrast X-ray imaging system according to claim 1, wherein the pixel size is 20 μm or more and 300 μm or less. 8. The digital phase contrast X-ray according to claim 1, wherein a size of a detection surface of the digital X-ray image detector is 15 cm × 15 cm or more and 90 cm × 90 cm or less. Image shooting system.