JP2011206161A - Radiographic system and image generation method therefor - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide phase contrast images of uniform image quality regardless of the inclination of small gratings.SOLUTION: A first absorption type grating 21 comprises a plurality of small gratings 21a-21e arrayed along a virtual cylindrical surface whose center axis is a virtual line C passing through an X-ray focus 11a. A second absorption type grating 22 comprises a plurality of small gratings 22a-22e arrayed along a virtual cylindrical surface which is coaxial with the cylindrical surface and has a radius larger than that of the cylindrical surface. A scanning mechanism 23 scans the second absorption type grating 22 in a y direction orthogonal to the virtual line C. The radiographic system includes a pitch correction amount storage part 25 for storing the correction of a scanning pitch corresponding to the inclination angle of the small gratings 22a-22e to the scanning direction for each segment corresponding to the small gratings 22a-22e. A phase differential image generation part 24 corrects the scanning pitch using a correction amount for each segment, calculates the phase deviation of the intensity change of pixel data with the corrected scanning pitch as a reference, and thus generates the phase differential image of a subject.

Description

  The present invention relates to a radiation imaging system that images a subject with radiation such as X-rays and an image generation method thereof, and more particularly to a radiation imaging system that performs phase imaging using two gratings and an image generation method thereof.

  X-rays are used as a probe for seeing through the inside of a subject because they have characteristics such as attenuation depending on the atomic numbers of elements constituting the substance and the density and thickness of the substance. X-ray imaging is widely used in fields such as medical diagnosis and non-destructive inspection.

  In a general X-ray imaging system, a subject is placed between an X-ray source that emits X-rays and an X-ray image detector that detects X-rays, and a transmission image of the subject is taken. In this case, the X-rays emitted from the X-ray source toward the X-ray image detector correspond to the difference in the characteristics (atomic number, density, thickness) of the substances existing on the path to the X-ray image detector. After receiving a certain amount of attenuation (absorption), it enters each pixel of the X-ray image detector. As a result, an X-ray absorption image of the subject is detected and imaged by the X-ray image detector. As an X-ray image detector, in addition to a combination of an X-ray intensifying screen and a film and a stimulable phosphor, a flat panel detector (FPD) using a semiconductor circuit is widely used.

  However, since the X-ray absorption ability is lower as a substance composed of an element having a smaller atomic number, a sufficient softness or contrast (contrast) of an X-ray absorption image cannot be obtained in a soft body tissue or soft material. There is. For example, most of the components of the cartilage part constituting the joint of the human body and the joint fluid in the vicinity thereof are water, and there is little difference in the amount of X-ray absorption between them, so that it is difficult to obtain a difference in light and shade.

  Against this background, in recent years, instead of X-ray intensity changes by the subject, an image based on the X-ray phase change (angle change) by the subject (hereinafter referred to as a phase contrast image) is obtained. Research on line phase imaging has been actively conducted. In general, when X-rays are incident on an object, the phase exhibits a higher interaction than the intensity of the X-rays. Therefore, X-ray phase imaging using a phase difference is a weakly absorbing object with low X-ray absorption capability. Can also obtain a high-contrast image. As a kind of such X-ray phase imaging, an X-ray imaging system using an X-ray Talbot interferometer composed of two transmission diffraction gratings and an X-ray image detector has been devised (for example, Patent Document 1). reference).

  The X-ray Talbot interferometer has a first diffraction grating disposed behind the subject, and a second diffraction grating downstream by a specific distance (Talbot interference distance) determined by the grating pitch of the first diffraction grating and the X-ray wavelength. Are arranged, and an X-ray image detector is arranged behind them. The Talbot interference distance is a distance at which X-rays that have passed through the first diffraction grating form a self-image (stripe image) due to the Talbot interference effect. This self-image is the same as the X-ray source and the first diffraction grating. Are modulated by the interaction (phase change) between the subject and the X-rays arranged between the two.

  In the X-ray Talbot interferometer, the phase contrast image of the subject is obtained from the change (phase shift) caused by the subject of the fringe image whose intensity is modulated by superimposing the self-image of the first diffraction grating and the second diffraction grating. To do. In the fringe scanning method, the second diffraction grating is substantially parallel to the surface of the first diffraction grating with respect to the first diffraction grating and substantially in the grating direction (strip direction) of the first diffraction grating. In the vertical direction, a plurality of times of imaging are performed while translational movement (scanning) is performed at a scanning pitch obtained by equally dividing the lattice pitch, and the phase of intensity change with respect to the scanning of the pixel data of each pixel obtained by the X-ray image detector is detected. This is a method of obtaining a phase differential image (corresponding to the angular distribution of X-rays refracted by the subject) from the amount of deviation (phase deviation amount with and without the subject H). The phase contrast image of the subject is obtained by integrating along the fringe scanning direction. Since the intensity of the pixel data is periodically modulated with respect to the scan, the set of pixel data for the scan is hereinafter referred to as an “intensity modulation signal”.

  In an X-ray imaging system using an X-ray Talbot interferometer, an X-ray source that normally emits X-rays from a focal point in a cone beam shape is used. Therefore, in order to ensure a wide field of view without degrading the image quality, the first and second diffraction gratings are arranged in a concave shape (cylindrical surface shape) with a line passing through the X-ray focal point as a central axis. It is desirable. However, since it is not easy to integrally create such a large concave lattice, it is conceivable to dispose a plurality of small lattices in a concave shape as disclosed in Patent Document 1. Patent Document 1 describes that in order to perform the above scanning, the small lattice is individually moved by a driving device provided for each small lattice, and the entire lattice is moved by a common driving device. Yes.

International Publication No. 2004/058070 JP 2007-203061 A

  As described in Patent Document 1, when a concave lattice is formed by a plurality of small lattices, it is not practical to provide a driving device for each small lattice because the structure becomes complicated and the cost increases. . Therefore, from the viewpoint of structure and cost, it is preferable to provide a common driving device so as to linearly move the entire lattice in one direction.

  However, when the concave grating as described above is linearly moved, the inclination angle of each small grating is different with respect to the linear movement direction, so that the effective scanning pitch is different for each small grating. For example, an effective scanning pitch changes from an actual scanning pitch because the small lattice in the peripheral portion has a larger inclination angle than the small lattice in the central portion of the lattice. That is, in this case, since a phase shift occurs for each small lattice, the phase shift caused by the tilt of the small lattice is added to the phase differential image obtained by actually placing the subject H as an offset. Become. For this reason, the phase contrast image has a problem that the image quality is lowered in the peripheral portion where the inclination of the small lattice is larger.

  The present invention has been made in view of the above problems, and an object of the present invention is to provide a radiation imaging system capable of obtaining a phase contrast image with uniform image quality regardless of the inclination of a small lattice and an image generation method thereof. To do.

  In order to achieve the above object, a radiation imaging system of the present invention includes a radiation source that emits radiation and a plurality of arrays arranged along a virtual cylindrical surface with a virtual line passing through the focal point of the radiation source as a central axis. A plurality of small lattices arranged along a virtual cylindrical surface that is coaxial with the cylindrical surface and has a larger radius than the cylindrical surface. A second grating having a lattice line in the imaginary line direction, and the second grating relative to the first grating at a predetermined scanning pitch in one direction orthogonal to the imaginary line. Scanning means for detecting the radiation that has passed through the first grating and the second grating at each position relatively moved by the scanning means to obtain image data, and for the one direction Inclination of each small lattice of the second lattice A pitch correction amount storage means for storing a correction amount of a scanning pitch corresponding to each degree for each segment corresponding to the small lattice, and correcting the scanning pitch using the correction amount for each segment, and scanning after correction A phase differential image generating means for generating a phase differential image of the subject by calculating the phase shift amount of the intensity change of the pixel data on the basis of the pitch, and a phase contrast image by generating the phase differential image are integrated. And a phase contrast image generation means.

  The correction amount is expressed as cos θ when the tilt angle is θ, and the phase differential image generation unit preferably corrects the scanning pitch by multiplying the scanning pitch by cos θ.

  Each of the small gratings of the first grating is preferably an absorption grating, and the radiation from the radiation source is projected as a fringe image onto the corresponding small grating of the second grating.

  Each of the small gratings of the first grating is a phase type grating, and the radiation from the radiation source is projected as a fringe image onto the corresponding small grating of the second grating by the Talbot interference effect. Is also preferable.

  Further, the radiation imaging system of the present invention comprises a radiation source that emits radiation and a plurality of small lattices arranged along a virtual cylindrical surface with a virtual line passing through the focal point of the radiation source as a central axis, Each pixel has a lattice in which the lattice line is in the imaginary line direction and a charge collection electrode that collects the charges converted by the radiation conversion layer, and each charge collection electrode is in phase with each other in a direction perpendicular to the imaginary line. A radiation image detector comprising a plurality of linear electrode groups arranged differently, and a correction amount of a scanning pitch corresponding to an inclination angle of each small lattice of the lattice with respect to a direction orthogonal to the virtual line, A pitch correction amount storage means for storing each segment corresponding to the above, and correcting the scanning pitch using the correction amount for each segment, and each linear electrode group based on the corrected scanning pitch. Phase differential image generating means for generating a phase differential image of the subject by calculating the amount of phase shift of intensity change of the pixel data obtained from the phase, and phase contrast generating a phase contrast image by integrating the phase differential image And an image generation means.

  Further, the image generation method of the present invention comprises a radiation source that emits radiation and a plurality of small lattices arranged along a virtual cylindrical surface with a virtual line passing through the focal point of the radiation source as a central axis, A first grid in which the grid line is in the virtual line direction, and a plurality of small grids arranged along a virtual cylindrical surface that is coaxial with the cylindrical surface and has a larger radius than the cylindrical surface. A second grating having the imaginary line direction, and a scanning means for moving the second grating relative to the first grating at a predetermined scanning pitch in one direction orthogonal to the imaginary line; An image used in a radiation imaging system comprising: a radiation image detector that detects image radiation that has passed through the first grating and the second grating at each position relatively moved by a scanning unit; A generation method, in the one direction The correction amount of the scanning pitch corresponding to the inclination angle of each small lattice of the second lattice is stored for each segment corresponding to the small lattice, and the scanning pitch is calculated using the correction amount for each segment. While correcting, the phase contrast image of the subject is generated by calculating the phase shift amount of the intensity change of the pixel data on the basis of the corrected scanning pitch, and the phase contrast image is integrated by integrating the phase differential image. It is characterized by generating.

  In the present invention, the correction amount of the scanning pitch corresponding to the inclination angle of each small lattice of the second lattice with respect to the scanning direction is stored for each segment corresponding to the small lattice, and scanning is performed using the correction amount for each segment. The phase differential image of the subject is generated by correcting the pitch and calculating the phase shift amount of the intensity change of the pixel data on the basis of the corrected scanning pitch, so that it is uniform regardless of the inclination of the small lattice. A phase contrast image with high image quality can be obtained.

1 is a schematic diagram illustrating a configuration of an X-ray imaging system according to a first embodiment of the present invention. It is the top view seen from the optical axis direction of the 1st and 2nd absorption type grating | lattice. It is a schematic diagram which shows the structure of a flat panel detector. It is a figure which shows the cross-section of the 1st and 2nd absorption type grating | lattice. It is explanatory drawing for demonstrating a fringe scanning method. It is a graph which illustrates pixel data (intensity modulation signal) modulated with respect to scanning. It is a schematic diagram which shows the effective grating period of the small grating | lattice inclined with respect to the scanning direction. It is a graph which shows the effective grating period of the small grating | lattice inclined with respect to the scanning direction. It is a figure which shows each segment which divided the detection surface of FPD. It is a schematic diagram which shows the structure of the X-ray image detector of 2nd Embodiment of this invention.

(First embodiment)
In FIG. 1, an X-ray imaging system 10 according to the present invention is disposed so as to face an X-ray source 11 that irradiates a subject H with X-rays, and transmits the subject H from the X-ray source 11. An imaging unit 12 that detects X-rays to generate image data, a memory 13 that stores image data read from the imaging unit 12, and a plurality of image data stored in the memory 13 are subjected to image processing and phase contrast An image processing unit 14 that generates an image, an image recording unit 15 that records a phase contrast image generated by the image processing unit 14, an imaging control unit 16 that controls the X-ray source 11 and the imaging unit 12, and an operation And a system control unit 18 that comprehensively controls the entire X-ray imaging system 10 based on an operation signal input from the console 17.

  The X-ray source 11 includes a high voltage generator, an X-ray tube, a collimator (all not shown), and the like, and irradiates the subject H with X-rays based on the control of the imaging control unit 16. For example, the X-ray tube is of a rotating anode type, emits an electron beam from a filament in accordance with a voltage from a high voltage generator, and collides the electron beam with a rotating anode rotating at a predetermined speed, thereby causing an X-ray. Is generated. The rotating anode rotates in order to reduce deterioration due to the electron beam continuing to hit the fixed position, and the collision part of the electron beam becomes an X-ray focal point that emits X-rays. The collimator limits the irradiation field so as to shield a portion of the X-ray emitted from the X-ray tube that does not contribute to the examination region of the subject H.

  The imaging unit 12 includes a flat panel detector (FPD) 20 made of a semiconductor circuit, a first absorption grating 21 for detecting phase change (angle change) of X-rays by the subject H, and performing phase imaging. Two absorption gratings 22 are provided. The FPD 20 is arranged so that the detection surface is orthogonal to a direction along the optical axis A of X-rays irradiated from the X-ray source 11 (hereinafter referred to as z direction).

  The 1st absorption type | mold grating | lattice 21 is comprised by the 1st-5th small grating | lattices 21a-21e. The first to fifth small gratings 21 a to 21 e are strip-like flat gratings extending in one direction (hereinafter, referred to as y direction) in a plane orthogonal to the z direction, and the X-ray focal point of the X-ray source 11. It is arranged along a virtual cylindrical surface that passes through 11a and has a virtual line C along the y direction as a central axis.

  The second absorption type grating 22 is configured by first to fifth small gratings 22 a to 22 e and is disposed between the first absorption type grating 21 and the FPD 20. Similarly, each of the first to fifth small lattices 22a to 22e is a strip-shaped flat lattice extending in the y direction, and is disposed along a virtual cylindrical surface with the imaginary line C as a central axis. Yes.

  Of the small lattices constituting the first absorption-type lattice 21, the first small lattice 21 a is arranged so as to be orthogonal to the optical axis A. Further, among the small lattices constituting the second absorption-type lattice 22, the second small lattice 22 a is arranged so as to be orthogonal to the optical axis A.

In FIG. 2A, the first to fifth small lattices 21a to 21e of the first absorption type lattice 21 are each X stretched in the y direction on an X-ray transparent substrate 30 such as a glass substrate. Line shielding portions (lattice lines) 31 are arranged in a direction orthogonal to the z direction and the y direction (hereinafter referred to as the x direction) at a predetermined pitch p _{1 and} with an interval d ₁ . Similarly, in FIG. 2B, the first to fifth small gratings 22a to 22e of the second absorption grating 22 are respectively X-rays extending in the y direction on the X-ray transparent substrate 32. shielding portion (grating lines) 33, at a predetermined pitch p ₂ in the x-direction, and in which they are arranged at intervals d _2.

As a material of the X-ray shielding portions 31 and 33, a metal excellent in X-ray absorption is preferable, and for example, gold or lead is preferable. The slit portions of the first and second absorption type gratings 21 and 22 (regions with the distances d ₁ and d ₂ ) need not be voids, and are filled with an X-ray low absorption material such as a polymer or a light metal. May be.

  The imaging unit 12 has a scanning mechanism 23 that changes the relative position of the first absorption type grating 21 with respect to the second absorption type grating 22 by moving the second absorption type grating 22 in the x direction. (Hereinafter, the movement of the second absorption type grating 22 by the scanning mechanism 23 is referred to as scanning, and this moving direction is referred to as the scanning direction). The scanning mechanism 23 is configured by an actuator such as a piezoelectric element, for example. The scanning mechanism 23 is driven based on the control of the imaging control unit 16 at the time of stripe scanning described later. As will be described in detail later, the memory 13 stores image data obtained by the photographing unit 12 at each scanning step of fringe scanning.

  The image processing unit 14 includes a phase differential image generation unit 24 that generates a phase differential image based on a plurality of image data captured by the imaging unit 12 and stored in the memory 13 at each scanning position of fringe scanning. A phase contrast image generation unit 25 that generates a phase contrast image by integrating the phase differential image along the x direction, and a scanning pitch corresponding to an inclination angle with respect to the scanning direction of the first to fifth small gratings 22a to 22e. And a pitch correction amount storage unit 26 for storing the correction amount. The phase contrast image generated by the phase contrast image generation unit 25 is recorded in the image recording unit 15 and then output to the console 17 and displayed on a monitor (not shown).

  In addition to the monitor, the console 17 includes an input device (not shown) through which an operator inputs a shooting instruction and the content of the instruction. As this input device, for example, a switch, a touch panel, a mouse, a keyboard or the like is used, and an X-ray imaging condition such as a tube voltage of the X-ray tube or an X-ray irradiation time, an imaging timing, etc. are input by operation of the input device. The The monitor is composed of a liquid crystal display or a CRT display, and displays characters such as X-ray imaging conditions and the phase contrast image.

  In FIG. 3, the FPD 20 includes an image receiving unit 41 in which a plurality of pixels 40 that convert X-rays into electric charges and are stored two-dimensionally on the active matrix substrate along the x direction and the y direction, and an image receiving unit 41. The scanning circuit 42 controls the readout timing of the charge from the pixel, and the readout circuit 43 reads the charge accumulated in each pixel 40, converts the charge into image data, and stores it. The scanning circuit 42 and each pixel 40 are connected by a scanning line 44 for each row, and the readout circuit 43 and each pixel 40 are connected by a signal line 45 for each column. The arrangement pitch of the pixels 40 is about 100 μm in each of the x direction and the y direction.

  The pixel 40 directly converts X-rays into charges by a conversion layer (not shown) such as amorphous selenium, and directly stores the converted charges in a capacitor (not shown) connected to an electrode below the conversion layer. It can be configured as a conversion type X-ray detection element. A TFT switch (not shown) is connected to each pixel 40, and the gate electrode of the TFT switch is connected to the scanning line 44, the source electrode is connected to the capacitor, and the drain electrode is connected to the signal line 45. When the TFT switch is turned on by the drive pulse from the scanning circuit 42, the charge accumulated in the capacitor is read out to the signal line 45.

Note that the pixel 40 temporarily converts X-rays into visible light using a scintillator (not shown) made of gadolinium oxide (Gd ₂ O ₃ ), cesium iodide (CsI), or the like, and converts the converted visible light into a photodiode ( It is also possible to configure as an indirect conversion type X-ray detection element that converts into electric charge and accumulates it (not shown). In this embodiment, an FPD based on a TFT panel is used as a radiation image detector. However, the present invention is not limited to this, and various types of radiation image detectors based on a solid-state imaging device such as a CCD sensor or a CMOS sensor. It is also possible to use.

  The readout circuit 43 includes an integration amplifier, a correction circuit, an A / D converter (all not shown), and the like. The integrating amplifier integrates the charges output from each pixel 40 via the signal line 45 and converts them into a voltage signal (image signal). The A / D converter converts the image signal converted by the integrating amplifier into digital image data. The correction circuit performs gain correction, linearity correction, and the like on the image data, and inputs the corrected image data to the memory 13.

  FIG. 4 shows a cross-sectional structure of the first and second absorption type gratings 21 and 22. The first and second absorption gratings 21 and 22 are arranged so that the corresponding small gratings are parallel to each other, and each small grating faces the X-ray focal point 11a. Each small lattice of the second absorption type grating 22 is arranged such that X-rays that have passed through the corresponding small lattice of the first absorption type grating 21 are projected. That is, the second absorption type grating 22 has a shape in which the first absorption type grating 21 is similarly enlarged with the X-ray focal point 11a as the center.

The first and second absorption gratings 21 and 22 are configured to linearly project the X-rays that have passed through the slit portions (regions with the distances d ₁ and d ₂ ) regardless of the presence or absence of the Talbot interference effect. ing. Specifically, by setting the distances d ₁ and d ₂ to a value sufficiently larger than the peak wavelength of X-rays emitted from the X-ray source 11, most of the X-rays included in the irradiated X-rays are slit at the slit portion. It is configured to pass through without being diffracted while maintaining straightness. For example, when tungsten is used as the rotating anode of the aforementioned X-ray tube and the tube voltage is 50 kV, the peak wavelength of the X-ray is about 0.4 mm. In this case, if the distances d ₁ and d ₂ are about ₁ μm to 10 μm, most of the X-rays are projected linearly without being diffracted by the slit portion. In this case, the lattice pitches p ₁ and p ₂ are about ₂ μm to 20 μm.

Since the X-rays emitted from the X-ray source 11 have a cone beam shape with the X-ray focal point 11a as a light emitting point, a projection image projected through the first absorption grating 21 (hereinafter, this projection image). Is referred to as a G1 image or a fringe image) is magnified in proportion to the distance from the X-ray focal point 11a. The grating pitch p ₂ and the interval d ₂ of the second absorption type grating 22 are determined so that the slit portions thereof substantially coincide with the periodic pattern of the bright part of the G1 image at the position of the second absorption type grating 22. Yes. That is, when the distance from the X-ray focal point 11a to the first absorption-type grating 21 is L ₁ and the distance from the first absorption-type grating 21 to the second absorption-type grating 22 is L ₂ , the grating pitch p ₂ and the distance d ₂ are determined so as to satisfy the relationship of the following expressions (1) and (2).

In the case of a Talbot interferometer, the distance L ₂ from the first absorption type grating 21 to the second absorption type grating 22 is limited to the Talbot interference distance determined by the grating pitch of the first diffraction grating and the X-ray wavelength. However, in the imaging unit 12 of the present embodiment, the first absorption type grating 21 projects the incident X-rays without diffracting, and the G1 image of the first absorption type grating 21 is the first one. because at every position of the rear absorption gratings 21 of similarly obtained, the distance L _2, can be set independently of the Talbot distance.

As described above, the imaging unit 12 of the present embodiment does not constitute a Talbot interferometer, but it is assumed that X-ray diffraction occurs in the first absorption grating 21 and a Talbot interference effect is generated. The Talbot interference distance Z is expressed by the following equation (3) using the grating pitch p _{1 of} the first absorption grating 21, the X-ray wavelength (peak wavelength) λ, and a positive integer m.

In the present embodiment, it is possible to set the distance L ₂ as described above irrespective of the Talbot distance, for the purpose of thinning in the z-direction of the imaging unit 12, the distance L _2, the case of m = 1 Is set to a value shorter than the minimum Talbot interference distance Z. That is, the distance L ₂ is set to a value in the range satisfying the following equation (4).

  The X-ray shielding units 31 and 33 preferably completely shield (absorb) X-rays in order to generate a periodic pattern image with high contrast. However, the X-ray shielding units 31 and 33 are excellent in X-ray absorption (such as gold or lead). Etc.), there are not a few X-rays that are transmitted without being absorbed. For this reason, in order to improve the shielding property of X-rays, it is preferable to increase the thickness of each of the X-ray shielding portions 31 and 33 (thickness in the z direction) as much as possible (that is, increase the aspect ratio). For example, when the tube voltage of the X-ray tube is 50 kV, it is preferable to shield 90% or more of the irradiated X-rays. In this case, the thickness of the X-ray shielding portions 31 and 33 is calculated in terms of gold (Au). It is preferable that it is 30 micrometers or more.

In the first and second absorption-type gratings 21 and 22 configured as described above, intensity modulation is performed by superimposing the G1 image (stripe image) of the first absorption-type grating 21 and the second absorption-type grating 22. The striped image is captured by the FPD 20. There is a slight difference between the pattern period of the G1 image at the position of the second absorption type grating 22 and the grating pitch p _{2 of} the second absorption type grating 22 due to manufacturing errors and arrangement errors. Thus, moire fringes are generated in the intensity-modulated fringe image. Further, when an error occurs in the lattice arrangement direction of the first and second absorption type gratings 21 and 22, and the arrangement directions are not the same, so-called rotational moire occurs. However, even when such moire fringes occur in the fringe image, there is no particular problem as long as the period of the moire fringes in the x direction or y direction is larger than the arrangement pitch of the pixels 40. Ideally, it is preferable not to generate moire fringes, but the moire fringes can be used to confirm the scanning amount during the fringe scanning, as will be described later.

  When the subject H is disposed between the X-ray source 11 and the first absorption type grating 21, the fringe image detected by the FPD 20 is modulated by the subject H. This modulation amount is proportional to the angle of the X-ray deflected by the refraction effect by the subject H. Therefore, a phase contrast image of the subject H can be generated by analyzing the fringe image detected by the FPD 20.

  Next, a fringe image analysis method will be described. In the figure, one X-ray refracted according to the phase shift distribution Φ (x) in the x direction of the subject H is illustrated. Reference numeral 50 indicates an X-ray path that travels straight when the subject H does not exist. The X-ray that travels along the path 50 passes through the first and second absorption gratings 21 and 22 to the FPD 20. Incident. Reference numeral 51 indicates an X-ray path refracted and deflected by the subject H when the subject H exists. X-rays traveling along the path 51 pass through the first absorption type grating 21 and are then shielded by the X-ray shielding part 33 of the second absorption type grating 22.

  The phase shift distribution Φ (x) of the subject H is expressed by the following equation (5), where n (x, z) is the refractive index distribution of the subject H and z is the direction in which the X-ray travels.

  The G1 image projected from the first absorption grating 21 to the position of the second absorption grating 22 is displaced in the x direction by an amount corresponding to the refraction angle φ due to refraction of X-rays at the subject H. It will be. This displacement amount Δx is approximately expressed by the following equation (6) based on the fact that the X-ray refraction angle φ is very small.

  Here, the refraction angle φ is expressed by the following equation (7) using the X-ray wavelength λ and the phase shift distribution Φ (x) of the subject H.

  Thus, the displacement amount Δx of the G1 image due to X-ray refraction at the subject H is related to the phase shift distribution Φ (x) of the subject H. This displacement amount Δx is the amount of phase shift ψ of the intensity modulation signal of each pixel 40 detected by the FPD 20 (the amount of phase shift of the intensity modulation signal of each pixel 40 with and without the subject H). ) Is related to the following equation (8).

  Precisely, the above equation (8) indicates that the displacement Δx is the phase of the intensity modulation signal of the pixel 40 corresponding to the first small grating 22 a that is not inclined with respect to the scanning direction of the second absorption grating 22. This is a relational expression for associating the deviation amount ψ with the displacement amount Δx.

  Therefore, the phase shift amount ψ of the intensity modulation signal of each pixel 40 is obtained and applied to the equation (8) to obtain the refraction angle φ, and further the equation (7) is applied to the phase shift distribution Φ (x). The differential amount is obtained. Then, by integrating this with respect to x, the phase shift distribution Φ (x) of the subject H, that is, the phase contrast image of the subject H can be generated. In the present embodiment, the phase shift amount ψ is calculated using the fringe scanning method shown below.

  In the fringe scanning method, imaging is performed while scanning one of the first and second absorption type gratings 21 and 22 in the x direction relative to the other (that is, imaging while changing the phase of both grating periods). I do). In the present embodiment, the second absorption type grating 22 is moved by the scanning mechanism 23 described above. As the second absorption type grating 22 moves, the moire fringe moves, and when the scanning amount reaches one period of the grating period (that is, when the phase change reaches 2π), the moire fringe returns to the original position. .

In the present embodiment, a fringe image is formed by the FPD 20 while moving the second absorption type grating 22 by an integral number of the grating pitch p ₂ with reference to the first small grating 22a that is not inclined with respect to the scanning direction. The intensity modulation signal of each pixel is acquired from the plurality of captured stripe images, and the phase differential image generation unit 24 in the image processing unit 14 performs arithmetic processing to thereby obtain the phase of the intensity modulation signal of each pixel 40. A shift amount ψ is obtained. In addition, since the other 2nd-5th small grating | lattices 22b-22e incline with respect to the scanning direction, each effective scanning pitch differs depending on inclination-angle (theta). As will be described in detail later, the above-described pitch correction amount storage unit 26 holds a correction amount (conversion coefficient) for converting the scanning pitch into an effective scanning pitch corresponding to the inclination angle θ. The quantity is supplied to the phase differential image generator 24.

FIG. 5 shows the G1 image and the first small grating 22a when the second absorption grating 22 is scanned by the scanning pitch (p ₂ / M) obtained by dividing the grating pitch p ₂ into M (an integer of 2 or more). The positional relationship is shown. The scanning mechanism 23 moves the second absorption type grating 22 to M scanning positions k = 0, 1, 2,..., M−1. In the figure, the initial position of the second absorption type grating 22 is set such that the dark part of the G1 image at the position of the second absorption type grating 22 when the subject H is not present is almost at the X-ray shielding part 33. Although the matching position (k = 0) is assumed, this initial position may be any position among k = 0, 1, 2,..., M−1.

  First, at the position of k = 0, mainly the X-rays that are not refracted by the subject H pass through the second absorption type grating 22. Next, when the second absorption type grating 22 is moved in order of k = 1, 2,..., The X-rays passing through the second absorption type grating 22 were not refracted by the subject H. While the X-ray component decreases, the X-ray component refracted by the subject H increases. In particular, at the position of k = M / 2, mainly only the X-rays refracted by the subject H pass through the second absorption type grating 22. When the position of k = M / 2 is exceeded, the X-ray component passing through the second absorption grating 22 is refracted by the subject H while the X-ray component refracted by the subject H decreases. The X-ray component that has not been increased.

When photographing is performed by the FPD 20 at each position of k = 0, 1, 2,..., M−1, M pixel data are obtained for each pixel 40. Hereinafter, a method of calculating the phase shift amount ψ of the intensity modulation signal of each pixel 40 from the M pixel data will be described. The pixel data of each pixel 40 at the position k of the second absorption type grating 22 is denoted as I _k (x). I _k (x) is expressed by the following equation (9).

Here, x is a coordinate in the x direction of the pixel, I ₀ is the intensity of the incident X-ray, and _An is a value corresponding to the contrast of the intensity modulation signal (where n is a positive integer). is there). Φ (x) represents the refraction angle φ as a function of the coordinate x of the pixel 40.

  Next, using the relational expression of the following expression (10), the refraction angle φ (x) is expressed as the expression (11).

  Here, arg [] means extraction of the declination and corresponds to the phase shift amount ψ. Therefore, by calculating the phase shift amount ψ from the M pieces of pixel data (intensity modulation signal) obtained from each pixel 40 based on the equation (11), the refraction angle φ (x) is obtained, and the phase shift distribution is obtained. The differential amount of Φ (x) is obtained.

Specifically, as shown in FIG. 6, M pixel data obtained in each pixel 40 is periodically with a period of the grating pitch p _{2 with} respect to the position k of the second absorption type grating 22. Change. The broken line in the figure shows the change of the pixel data when the subject H does not exist, and the solid line in the figure shows the change of the pixel data when the subject H exists. The phase difference between the two waveforms corresponds to the phase shift amount ψ.

  In the above description, the y coordinate in the y direction of the pixel 40 is not taken into consideration. However, by performing the same calculation for each y coordinate, a two-dimensional phase shift distribution ψ (x, y) is obtained. This phase shift distribution ψ (x, y) corresponds to the phase differential image.

  Since the phase shift amount ψ (x, y) obtained by the above equation (11) does not take into account the inclination angle θ of the small lattice, in order to obtain the actual phase shift amount that depends only on the subject H, For each segment corresponding to the second to fifth small lattices 22b to 22e, it is necessary to calculate the phase shift amount with an effective scanning pitch according to the inclination angle θ. The phase differential image generation unit 24 corrects the scanning pitch to an effective scanning pitch for each segment based on the correction amount supplied from the pitch correction amount storage unit 26, and then the phase shift amount ψ (x, y) That is, a phase differential image is calculated.

FIG. 7 shows one small lattice among the second to fifth small lattices 22b to 22e. Although the small grating 60 has a grating period of p ₂ and has an inclination angle θ with respect to the scanning direction, the effective grating period p ₂ ′ is expressed as p ₂ ′ = p ₂ / cos θ. FIG. 8 shows the period p ₂ of the intensity modulation signal corresponding to the small grating without inclination (θ = 0) and the intensity modulation signal corresponding to the small grating with the inclination angle θ when the subject H is not arranged. Period p ₂ 'is shown.

Therefore, since the relationship of p ₂ / M = (p ₂ ′ / M) cos θ is established, the slope of the inclination that occurs when the second absorption type grating 22 is moved by the scanning pitch (p ₂ / M) by the scanning mechanism 23. If the phase modulation amount of a small lattice that is not present is δ, then the phase modulation that occurs in the small lattice with the tilt angle θ is δ cos θ. That is, the effective scanning pitch of the small grating with the inclination angle θ is cos θ times when there is no inclination. This “cos θ” is a scanning pitch correction amount (conversion coefficient) stored in the pitch correction amount storage unit 26.

  Next, the calculation principle of the phase shift amount ψ will be described. This calculation principle is based on the method disclosed in “Introduction to Applied Optical Measurement” by Toyohiko Yadagai, the second edition, published on February 15, 2005, Maruzen Co., Ltd. (pages 196 to 198). .

  First, assuming that the change (intensity modulation signal) of the pixel data with respect to the scanning position k when the subject H is not arranged is an ideal sine wave, the following expression (12) is expressed.

Here, I _k is the intensity at the scanning position k, and the x and y coordinates are omitted. Α is a parameter corresponding to contrast. Δ _k is a phase modulation amount corresponding to an effective scanning pitch in the case of the inclination angle θ, as represented by the following equation (13).

Next, the above expression (12) is expanded as shown in the following expression (14).

Here, a ₀ , a ₁ and a ₂ are as follows.

Next, assuming that the measured value of I _k when the subject H is arranged is I _k ′, the phase shift amount of the intensity modulation signal by the subject H is calculated by obtaining the most probable ψ from the measured value I _k ′. Is done. That is, the phase shift amount ψ can be obtained by determining a ₀ , a ₁ , and a ₂ so as to minimize the difference between I _k ′ and I _k . Specifically, first, an index W of the difference between I _k ′ and I _k is set as in the following equation (18).

In order to obtain a ₀ , a ₁ , and a ₂ that minimize the index W, the following equations (19) to (21) may be solved.

  The above formulas (19) to (21) are expressed as follows using a matrix expression.

ここで、

である。

here,

It is.

By substituting a ₁ and a ₂ obtained by solving the above equation (22) into the following equation (26), the phase shift amount ψ is obtained. The following equation (26) is based on the above equations (16) and (17).

  As shown in FIG. 9, the pitch correction amount storage unit 26 classifies the image receiving unit 41 (detection surface) of the FPD 20 so as to correspond to the first to fifth small gratings 22 a to 22 e of the second absorption type grating 22. For each of the segments SG1 to SG5, the value of cos θ (scanning pitch correction amount) included in the equation (13) is stored, and this value is supplied to the phase differential image generation unit 24. The phase differential image generation unit 24 calculates the phase shift amount ψ by using the above formula (13) and the above formulas (22) to (26) for each pixel data of each pixel 40. The value of cos θ for the segment to which it belongs is applied to the above equation (13).

  In the X-ray imaging system 10 configured as described above, when the subject H is placed between the X-ray source 11 and the imaging unit 12, an X-ray source is input from the console 17 by the operator. The X-ray irradiation is performed from 11 to the subject H, and the FPD 20 detects a fringe image whose intensity is modulated by passing through the subject H and passing through the first and second absorption gratings 21 and 22. The The fringe image detection operation is performed at each scanning position while moving the second absorption type grating 22 by a predetermined scanning pitch. The fringe image obtained by each detection operation is stored in the memory 13 as image data.

  Next, the phase differential image generation unit 24 generates the phase shift amount ψ (x, y) of the intensity modulation signal based on the plurality of image data stored in the memory 13, and at this time, the segments SG1 to SG5 Each time, the correction amount (cos θ) of the scanning pitch is supplied from the pitch correction amount storage unit 26, and the phase differential image generation unit 24 calculates the phase shift amount ψ (x, y) using the corrected effective scanning pitch. Calculate to generate a differential phase image.

  The phase differential image generated by the phase differential image generation unit 24 is subjected to integration processing by the phase contrast image generation unit 25 to be converted into a phase contrast image, recorded in the image recording unit 15, and then on the monitor of the console 17. Is displayed.

  Note that the pitch correction amount storage unit 26 is preferably configured with a data rewritable memory (flash memory or the like) so that the correction amount (cos θ) can be changed when the inclination angle θ changes.

  Moreover, in the said embodiment, although the 1st and 2nd absorption type | mold grating | lattices 21 and 22 are comprised by the several small grating | lattice divided | segmented into the x direction, these small grating | lattices are further divided | segmented into the y direction. May be. That is, the first and second absorption type gratings 21 and 22 may be configured by small gratings arranged in a matrix.

Moreover, in the said embodiment, although the 1st absorption-type grating | lattice 21 is comprised so that the X-ray which passed the slit part between the X-ray shielding parts 31 may be projected linearly, this invention is this structure. The present invention is not limited to this, and a configuration in which a so-called Talbot interference effect is generated by diffracting X-rays at the slit portion (a configuration described in International Publication No. 2004/058070) may be employed. However, in this case, it is necessary to set the first absorption grating 21 the distance L ₂ between the second absorption grating 22 to the Talbot distance. In this case, a phase type grating (phase type diffraction grating) can be used in place of the first absorption type grating 21. The phase type grating used as the first absorption type grating 21 projects a fringe image (self-image) generated by the Talbot interference effect onto the FPD 20.

  The difference between the phase type grating and the absorption type grating is only the thickness of the X-ray high absorption material (X-ray shielding part), and the thickness of the X-ray shielding part is about Au in the case of the absorption type grating. While it is set to 30 μm or more, in the case of a phase type grating, it is set to about 1 μm to 5 μm. In the phase-type grating, the X-ray high-absorbing material applies a predetermined amount (π or π / 2) of phase modulation to the X-rays irradiated from the X-ray source 11, thereby generating a Talbot interference effect and causing fringes. An image (self-image) is generated.

  In the above embodiment, the subject H is disposed between the X-ray source 11 and the first absorption type grating 21. However, the subject H is arranged between the first absorption type grating 21 and the second absorption type. A phase contrast image can be generated in the same manner when it is disposed between the grating 22 and the grating 22.

(Second Embodiment)
In the first embodiment, the second absorption type grating 22 is provided independently of the FPD 20. However, as the second embodiment, X-ray image detection having the configuration disclosed in Japanese Patent Laid-Open No. 2009-133823 is provided. By using the vessel, the second absorption type grating 22 can be eliminated. The X-ray image detector according to the second embodiment is a direct conversion type X-ray image detection including a conversion layer that converts X-rays into electric charges, and a charge collection electrode that collects electric charges converted in the conversion layers. In this device, the charge collection electrode of each pixel is configured by arranging a plurality of linear electrode groups formed by electrically connecting linear electrodes arranged at a constant period so that their phases are different from each other. ing.

  FIG. 10 illustrates the configuration of the X-ray image detector (FPD) of this embodiment. The pixels 70 are two-dimensionally arranged at a constant pitch along the x and y directions, and each pixel 70 has a charge collection for collecting the charges converted by the conversion layer that converts X-rays into charges. An electrode 71 is formed. The charge collection electrode 71 includes first to sixth linear electrode groups 72 to 77, and the phase of the arrangement period in the x direction of the linear electrodes of each linear electrode group is shifted by π / 3. Specifically, when the phase of the first linear electrode group 72 is 0, the phase of the second linear electrode group 73 is π / 3, the phase of the third linear electrode group 74 is 2π / 3, The phase of the fourth linear electrode group 75 is π, the phase of the fifth linear electrode group 76 is 4π / 3, and the phase of the sixth linear electrode group 77 is 5π / 3.

  Further, each pixel 70 is provided with a switch group 78 for reading out the charges collected by the charge collecting electrode 71. The switch group 78 includes TFT switches provided in each of the first to sixth linear electrode groups 72 to 77. The charges collected by the first to sixth linear electrode groups 72 to 77 are individually read out by controlling the switch group 78, so that six kinds of fringes having different phases with respect to the x direction are obtained by one photographing. An image can be acquired, and a phase contrast image can be generated based on the phase shift amount of the intensity modulation signal obtained for each pixel from the six types of stripe images.

  By using the X-ray image detector having the above configuration in place of the FPD 20, the second absorption type grating 22 is not required from the imaging unit 12, so that the cost can be reduced and the thickness can be further reduced. Further, in the present embodiment, since it is possible to acquire the intensity modulation signal by one shooting, physical scanning for fringe scanning becomes unnecessary, and the scanning mechanism 23 can be eliminated. Instead of the charge collection electrode 71, it is possible to use a charge collection electrode having another configuration described in Japanese Patent Laid-Open No. 2009-133823.

When the second absorption type grating 22 is eliminated in this way, the G1 image projected on the detection surface of the X-ray image detector becomes the first to fifth small gratings 21a of the first absorption type grating 21. Therefore, the effective pattern period of the G1 image on the detection surface is p ₂ ′ = p ₂ / cos θ. Here, _{p 2} is the arrangement pitch of the linear electrodes of each linear electrode group 72 to 77.

  Therefore, also in the present embodiment, as in the first embodiment, the phase differential image generation unit 24 sets the scanning pitch to an effective scanning pitch based on the correction amount supplied from the pitch correction amount storage unit 26. After correction, a phase differential image is generated. Specifically, the pitch correction amount storage unit 26 stores a correction amount (cos θ) corresponding to the inclination angle θ of the first to fifth small lattices 21 a to 21 e of the first absorption type lattice 21. The detection surface of the X-ray image detector is segmented (similar to FIG. 9), the correction amount corresponding to each segment is used to correct the scanning pitch to an effective scanning pitch, and the phase shift amount ψ (x, y ) Is calculated.

  Furthermore, as another embodiment when the second absorption type grating 22 is not disposed, the fringe image (G1 image) obtained by the X-ray image detector is periodically sampled while changing the phase by signal processing. Thus, it is also possible to apply intensity modulation to the fringe image.

  The present invention can be applied to other radiography systems for industrial use, in addition to radiography systems for medical diagnosis.

10 X-ray imaging system 11 X-ray source (radiation source)
11a X-ray focus 12 Imaging unit 14 Image processing unit 20 Flat panel detector (FPD)
DESCRIPTION OF SYMBOLS 21 1st absorption-type grating | lattice 21a-21e Small grating | lattice 22 2nd absorption-type grating | lattice 22a-22e Small grating | lattice 23 Scanning mechanism 24 Phase differential image generation part 25 Phase contrast image generation part 26 Pitch correction amount memory | storage part 30, 32 X-ray Translucent substrate 31, 33 X-ray shielding portion 40 pixels 41 image receiving portion 60 small lattice 70 pixels 71 charge collecting electrodes 72 to 77 first to sixth linear electrode groups 78 switch groups

Claims (6)

A radiation source that emits radiation; and
A plurality of small lattices arranged along a virtual cylindrical surface with a virtual line passing through the focal point of the radiation source as a central axis, wherein the lattice line is the virtual line direction;
A plurality of small lattices arranged along a virtual cylindrical surface coaxial with the cylindrical surface and having a larger radius than the cylindrical surface, and a second lattice in which the lattice line is in the imaginary line direction;
Scanning means for moving the second grating relative to the first grating at a predetermined scanning pitch in one direction orthogonal to the virtual line;
A radiation image detector for obtaining image data by detecting radiation that has passed through the first grating and the second grating at each position relatively moved by the scanning means;
Pitch correction amount storage means for storing, for each segment corresponding to the small lattice, a correction amount of a scanning pitch corresponding to an inclination angle of each small lattice of the second lattice with respect to the one direction;
A phase for generating a phase differential image of a subject by correcting a scanning pitch using the correction amount for each segment and calculating a phase shift amount of intensity change of pixel data on the basis of the corrected scanning pitch. Differential image generation means;
Phase contrast image generating means for generating a phase contrast image by integrating the phase differential image;
A radiation imaging system comprising:   The correction amount is represented as cos θ when the tilt angle is θ, and the phase differential image generation unit corrects the scanning pitch by multiplying the scanning pitch by cos θ. The radiation imaging system described.   2. Each of the small gratings of the first grating is an absorption grating, and the radiation from the radiation source is projected as a fringe image onto the corresponding small grating of the second grating. Or the radiography system of 2.   Each of the small gratings of the first grating is a phase type grating, and the radiation from the radiation source is projected as a fringe image onto the corresponding small grating of the second grating by the Talbot interference effect. The radiation imaging system according to claim 1 or 2. A radiation source that emits radiation; and
A plurality of small lattices arranged along a virtual cylindrical surface with a virtual line passing through the focal point of the radiation source as a central axis, and a lattice line in the virtual line direction;
A plurality of linear electrode groups each provided with a charge collection electrode for collecting charges converted by the radiation conversion layer, wherein each charge collection electrode is arranged in a direction orthogonal to the virtual line so as to have different phases from each other A radiation image detector comprising:
A pitch correction amount storage means for storing a correction amount of a scanning pitch corresponding to an inclination angle of each small lattice of the lattice with respect to a direction orthogonal to the virtual line, for each segment corresponding to the small lattice;
By correcting the scan pitch using the correction amount for each segment, and calculating the phase shift amount of the intensity change of the pixel data obtained by each linear electrode group on the basis of the corrected scan pitch Phase differential image generating means for generating a phase differential image of
Phase contrast image generating means for generating a phase contrast image by integrating the phase differential image;
A radiation imaging system comprising: A radiation source that emits radiation; and
A plurality of small lattices arranged along a virtual cylindrical surface with a virtual line passing through the focal point of the radiation source as a central axis, wherein the lattice line is the virtual line direction;
A plurality of small lattices arranged along a virtual cylindrical surface coaxial with the cylindrical surface and having a larger radius than the cylindrical surface, and a second lattice in which the lattice line is in the imaginary line direction;
Scanning means for moving the second grating relative to the first grating at a predetermined scanning pitch in one direction orthogonal to the virtual line;
Used in a radiation imaging system comprising: a radiation image detector that obtains image data by detecting radiation that has passed through the first grating and the second grating at each position relatively moved by the scanning means. An image generation method comprising:
A correction amount of a scanning pitch corresponding to an inclination angle of each small lattice of the second lattice with respect to the one direction is stored for each segment corresponding to the small lattice,
While correcting the scan pitch using the correction amount for each segment, and generating a phase differential image of the subject by calculating the phase shift amount of the intensity change of the pixel data on the basis of the corrected scan pitch,
An image generation method comprising: generating a phase contrast image by integrating the phase differential image.

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