JP2011218147A - Radiographic system - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a radiographic system efficiently removing scattered light to obtain a phase contrast image of high quality.SOLUTION: In the radiographic system 10 for acquiring the phase contrast image of a specimen H from a plurality of image data obtained by performing radiographing while altering the relative position of a first absorption type grating 21 and a second absorption type grating 22, a scattered light removing grid 24, of which the grating pattern is arranged so as to cross the grating pattern of the first absorption type grating 21 at a right angle, is provided between the first and second absorption type gratings 21 and 22. The scattered light removing grid 24 performs Bucky operation in the direction crossing the grating pattern at a right angle by a reciprocating device 25.

Description

  The present invention relates to a radiation imaging system that images a subject with radiation such as X-rays, and more particularly to a radiation imaging system that enables phase imaging of a subject.

  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 one type of such X-ray phase imaging, an X-ray imaging system using an X-ray Talbot interferometer including two transmission diffraction gratings and an X-ray image detector is known (for example, Patent Document 1). reference).

  The X-ray Talbot interferometer arranges a first diffraction grating behind the subject, and arranges a second diffraction grating downstream by a Talbot interference distance determined by the grating pitch of the first diffraction grating and the X-ray wavelength, It is configured by placing an X-ray image detector behind it. The Talbot interference distance is a distance at which X-rays that have passed through the first diffraction grating form a self image (a fringe image) due to the Talbot interference effect. This self-image is modulated by the interaction (phase change) between the subject arranged between the X-ray source and the first diffraction grating and the X-ray.

  In this X-ray imaging system, the phase of the subject is detected by the fringe scanning method from the change (phase shift) caused by the subject in the fringe image whose intensity is modulated by the superposition of the self-image of the first diffraction grating and the second diffraction grating. A contrast image is acquired. In the fringe scanning method, the second diffraction grating is arranged with respect to the first diffraction grating in a direction substantially parallel to the plane of the first diffraction grating and substantially perpendicular to the grating line direction of the first diffraction grating. Imaging is performed a plurality of times while translationally moving (scanning) at a scanning pitch obtained by equally dividing the lattice pitch, and the phase shift amount of the intensity change with respect to the scanning of the pixel data of each pixel obtained by the X-ray image detector (depending on the subject) This is a method of obtaining a phase differential image from the phase shift amount). This phase differential image corresponds to the angular distribution of X-rays refracted by the subject. A phase contrast image of the subject can be obtained by integrating the phase differential image along the fringe scanning direction. Note that the intensity of the pixel data is periodically modulated by the scanning. A set of a plurality of pixel data for the scanning is hereinafter referred to as “intensity modulation signal”.

  As described above, the X-ray imaging system using the X-ray Talbot interferometer images the amount of refraction of X-rays by the subject. In addition to refraction according to the refractive index distribution, there are scattering phenomena such as Compton scattering and Rayleigh scattering that occur in the subject. Such a scattering phenomenon is known to deteriorate the image quality of the phase contrast image because it gives diffuse noise to the X-ray image detector (see Patent Document 2). Patent Document 2 describes that the second diffraction grating arranged immediately before the X-ray image detector acts as a scattered radiation removal grid (scattering prevention grating).

JP 2008-200399 A JP-T-2008-545981 (page 13)

  However, even if the second diffraction grating is caused to act as a scattered radiation removal grid as described in Patent Document 2, the second diffraction grating is effective in removing scattered radiation in the direction of the array pattern of the grating lines. However, there is a problem that the effect of removing scattered radiation cannot be obtained in the direction (lattice line direction) orthogonal to the array pattern direction of the grid lines.

  The present invention has been made in view of the above problems, and in a radiation imaging system that performs phase imaging using a grating, the radiation imaging system can efficiently remove scattered radiation and obtain a high-quality phase contrast image. The purpose is to provide.

  In order to achieve the above object, a radiation imaging system of the present invention includes a first grating that generates a first periodic pattern image by passing the radiation emitted from a radiation source, and the first periodic pattern image. Intensity modulating means for applying intensity modulation to generate a second periodic pattern image, a radiation image detector for detecting the second periodic pattern image and generating image data, the radiation source, and the radiation A scattered radiation elimination grid arranged between the image detector and a grating pattern so as to be orthogonal to the grating pattern of the first grating, and a phase differential image generating means for generating a phase differential image based on the image data And comprising.

  The radiation imaging system of the present invention further includes reciprocating means for reciprocating the scattered radiation removing grid in a direction orthogonal to the lattice pattern of the first grating.

  The scattered radiation removal grid is a focusing grid having directivity with respect to the focal point of the radiation source.

  The radiation imaging system of the present invention further includes phase contrast image generation means for generating a phase contrast image by performing integration processing on the image data generated by the radiation image detector.

  The intensity modulation means applies intensity modulation at a plurality of relative positions different in phase with respect to the first periodic pattern image to generate a plurality of second periodic pattern images, and the radiological image detector includes The second periodic pattern image is detected to generate a plurality of image data, and the phase differential image generating means is based on the plurality of image data, and the phase of the intensity modulation signal representing the intensity change of the pixel data with respect to the relative position The phase differential image is generated by calculating the amount of deviation.

  The intensity modulation means includes: a second grating having a periodic pattern in the same direction as the first periodic pattern image; and scanning means for moving one of the first and second gratings at a predetermined pitch. Become.

  The scattered radiation removal grid is disposed between the first grating and the second grating.

  The first grating is a phase grating, and the first grating forms radiation from the radiation source as the first periodic pattern image at the position of the second grating by the Talbot interference effect.

  The radiological image detector is a radiological image detector including a conversion layer for converting radiation into electric charges and a charge collection electrode for collecting electric charges converted in the conversion layer for each pixel. A plurality of linear electrode groups having a periodic pattern in the same direction as the first periodic pattern image are arranged so that their phases are different from each other, and the intensity modulation means is constituted by the charge collecting electrode. It may be a thing.

  Since the radiation imaging system of the present invention includes the scattered radiation removal grid arranged so that the grating pattern is orthogonal to the grating pattern of the first grating, the scattered radiation is efficiently removed and a high-quality phase contrast image is obtained. be able to. Further, by arranging the scattered radiation removal grid between the first grating and the second grating, it is possible to install the scattered radiation removal grid without increasing the size of the imaging unit.

It is a schematic diagram which shows the structure of the X-ray imaging system of this invention. It is a schematic sectional drawing which shows the structure of a scattered radiation removal grid. It is a schematic diagram which shows the structure of a flat panel detector. It is a schematic sectional drawing of the imaging | photography part cut | disconnected so that it might pass along the X-ray transmissive part of a scattered radiation removal grid. It is a schematic side view shown. It is explanatory drawing for demonstrating a fringe scanning method. It is a graph which illustrates an intensity modulation signal with and without a subject. It is a schematic sectional drawing along yz plane of an imaging part. It is a schematic diagram which shows the modification of an X-ray image detector.

  In FIG. 1, an X-ray imaging system 10 according to the first embodiment of the present invention is arranged so as to be opposed to an X-ray source 11 that irradiates a subject H with X-rays, and from the X-ray source 11 An imaging unit 12 that detects X-rays transmitted through the specimen H and generates 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 Image processing unit 14 for processing to generate a phase contrast image, image recording unit 15 for recording the phase contrast image generated by the image processing unit 14, and imaging control for controlling the X-ray source 11 and the imaging unit 12 Unit 16, a console 17 including an operation unit and a monitor, 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. That.

  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 first absorption type grating 21 has a plurality of X-ray shielding portions 21a extending in one direction (hereinafter referred to as y direction) in a plane orthogonal to the z direction. These are arranged at a predetermined pitch p ₁ in a direction perpendicular to the direction and the y direction (hereinafter referred to as the x direction). Similarly, the second absorption grating 22, in which a plurality of X-ray shielding portions 22a which extend in the y direction, are arranged at a predetermined pitch p ₂ in the x-direction to the X-ray transmissive substrate (not shown) is there. The grating pitches p ₁ and p ₂ are about ₂ to 20 μm. In addition, as a material of X-ray shielding part 21a, 22a, the metal which is excellent in X-ray absorptivity is preferable, for example, gold | metal | money, lead, etc. are preferable.

  In addition, the imaging unit 12 translates the second absorption type grating 22 in the direction (x direction) orthogonal to the lattice line direction (y direction), so that the second absorption with respect to the first absorption type grating 21 is performed. A scanning mechanism 23 for changing the relative position with respect to the mold grating 22 is provided. 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 second absorption type grating 22 and the scanning mechanism 23 constitute the intensity modulation means described in the claims.

Further, in the imaging unit 12, a scattered radiation removal grid 24 is disposed between the first absorption type grating 21 and the second absorption type grating 22. Scatter grid 24 is for a plurality of X-ray absorbing portion 24a which extends in the x direction are arranged at a predetermined pitch p ₃ in the y-direction, the X-ray absorption unit 24a, such as lead which absorbs X-rays It is formed with. That is, the lattice pattern of the scattered radiation removal grid 24 is orthogonal to the lattice patterns of the first and second absorption type gratings 21 and 22.

As shown in FIG. 2, the scattered radiation removal grid 24 is a converging grid in which the X-ray absorbing portion 24a is inclined toward the X-ray focal point 11a, and has directivity with respect to the X-ray focal point 11a. ing. Between the adjacent X-ray absorbing portion 24a, the X-ray transmitting portions 24b made of resin or the like is interposed, the pitch p ₃ is a typical anti-scatter grid comparable (about 160 .mu.m).

  Furthermore, the imaging unit 12 is provided with a reciprocating device 25 that reciprocates (oscillates) the scattered radiation removal grid 24 in the y direction independently of the first and second absorption gratings 21 and 22. ing. The reciprocating device 25 causes the scattered radiation removal grid 24 to perform a so-called bucky operation, and a fixed X-ray non-irradiation region (fixed shadow) is generated on the detection surface of the FPD 20 by the scattered radiation removal grid 24. To prevent that.

  The image processing unit 14 includes a phase differential image generation unit 26 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 step of fringe scanning, and a phase differential image. Is integrated along the x direction to provide a phase contrast image generation unit 27 that generates a phase contrast image. The phase contrast image generated by the phase contrast image generation unit 27 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. By operating the input device, tube voltage of the X-ray tube, tube current, X-ray imaging conditions such as X-ray irradiation time, imaging timing, and the like are input. 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 offset correction, gain correction, linearity correction, and the like on the image data, and inputs the corrected image data to the memory 13.

In FIG. 4, X-ray shielding portion 21a of the first absorption grating 21, at a predetermined pitch p ₁ in the x direction, and are arranged at a predetermined interval d ₁ from each other. Similarly, X-rays shielding portions 22a of the second absorption-type grating 22, at a predetermined pitch p ₂ in the x direction, and are arranged at a predetermined interval d ₂ from each other. The first and second absorption gratings 21 and 22 do not give a phase difference to incident X-rays but give an intensity difference, and are also called amplitude gratings. Incidentally, (a region of the interval d _1, d ₂₎ slit portion may not be a void, or may be low X-ray absorbing material such as a polymer or light metal is interposed. This figure shows a cross section passing through the X-ray transmission part 24 b of the scattered radiation removal grid 24.

The first and second absorption type gratings 21 and 22 are configured to project X-rays that have passed through the slit portion linearly (geometrically). 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 to 10 μm, most of the X-rays are projected linearly without being diffracted by the slit portion.

The X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam having the X-ray focal point 11a as a light emitting point, and therefore the first X-ray formed by passing through the first absorption type grating 21 is used. The periodic pattern image (hereinafter referred to as G1 image) is enlarged 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 grating 21 projects the incident X-rays without diffracting, and the G1 image of the first absorption 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 portions 21a and 22a preferably completely shield (absorb) X-rays in order to generate a periodic pattern image with high contrast, but the above-described materials (gold, lead) having excellent X-ray absorption properties 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 21a and 22a (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 thicknesses of the X-ray shielding portions 21a and 22a are 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, the G1 image generated by the first absorption type grating 21 is partially overlapped with the second absorption type grating 22. A second periodic pattern image (hereinafter referred to as G2 image) is generated by being shielded and intensity-modulated. This G2 image is picked up 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 G2 image.

  When the subject H is disposed between the X-ray source 11 and the first absorption type grating 21, the G2 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 G2 image detected by the FPD 20.

  Next, a G2 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 goes straight when the subject H does not exist. X-rays traveling along the path 50 pass through the first and second absorption gratings 21 and 22 and enter the FPD 20. 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 22 a 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 type grating 21 to the position of the second absorption type 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).

  Accordingly, by obtaining the phase shift amount ψ of the intensity modulation signal of each pixel 40, the refraction angle φ is obtained from the equation (8), and the differential amount of the phase shift distribution Φ (x) is obtained using the equation (7). 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 one of the first and second absorption type gratings 21 and 22 is translated in the x direction relative to the other (that is, while changing the phase of the grating period of both). Take a picture). In the present embodiment, the second absorption type grating 22 is moved by the scanning mechanism 23 described above. The moire fringes generated in the G2 image move with the movement of the second absorption type grating 22, and the translational distance (the amount of movement in the x direction) is one period of the grating period of the second absorption type grating 22 (grating When the pitch (p ₂ ) is reached (that is, when the phase change reaches 2π), the original position is restored. Thus, while moving the second absorption grating 22 by an integral fraction of the grating pitch _{p 2,} taking a G2 image in FPD 20. The intensity modulation signal of each pixel is acquired from a plurality of image data obtained by photographing, and is processed by the phase differential image generation unit 26 in the image processing unit 14 described above, whereby the phase shift of the intensity modulation signal of each pixel. The quantity ψ is obtained. This two-dimensional distribution of the phase shift amount ψ corresponds to a phase differential image.

FIG. 5 schematically shows how the second absorption grating 22 is moved by the scanning pitch (p ₂ / M) obtained by dividing the grating pitch p ₂ into M (an integer of 2 or more). The scanning mechanism 23 sequentially translates the second absorption type grating 22 to each of M scanning positions k = 0, 1, 2,..., M−1. In this 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 22a. 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-ray component (non-refractive component) that has not been refracted by the subject H passes through the second absorption type grating 22. Next, when the second absorption type grating 22 is moved in order of k = 1, 2,..., X-rays passing through the second absorption type grating 22 are reduced in non-refractive components. The X-ray component (refractive component) refracted by the subject H increases. In particular, at the position of k = M / 2, only the refraction component mainly passes through the second absorption type grating 22. When the position exceeds k = M / 2, the X-ray passing through the second absorption grating 22 decreases the refractive component while increasing the non-refractive component.

When photographing is performed by the FPD 20 at each position of k = 0, 1, 2,..., M−1 to generate image data, M pixel data is 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 I _k (x) of each pixel 40 at the position k of the second absorption type lattice 22 is generally expressed by the following equation (9).

Here, x is the coordinate in the x direction of the pixel, A ₀ is the intensity of the incident X-ray, A _n is a value corresponding to the contrast of the intensity-modulated signal, n represents a positive integer, i is the imaginary unit. 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 based on the equation (11) from the intensity modulation signal represented by M pixel data obtained at each pixel 40, the refraction angle φ (x) is obtained, and the phase A differential amount of the shift distribution Φ (x) is obtained.

Specifically, as shown in FIG. 6, the intensity modulation signal periodically changes with the period of the grating pitch p _{2 with} respect to the position k of the second absorption type grating 22. The broken line in the figure shows the intensity modulation signal when the subject H does not exist, and the solid line in the figure shows the intensity modulation signal 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.

  This phase differential image is input to the phase contrast image generation unit 27. The phase contrast image generation unit 27 generates a phase shift distribution Φ (x, y) of the subject H by integrating the input phase differential image along the x axis, and records this as a phase contrast image. Part 15 is recorded.

Next, the bucky operation of the scattered radiation removal grid 24 by the reciprocating device 25 will be described. The reciprocating device 25 reciprocates the scattered radiation removal grid 24 in the y direction during each imaging period performed at each scanning position k by the scanning mechanism 23. Specifically, a reciprocating operation device 25, the anti-scatter grid 24 in one imaging period, moving a sufficiently large distance (e.g., 1 cm) with respect to the arrangement pitch p ₃ of the X-ray absorbing portion 24a. The moving distance, and more preferably an integral multiple of the arrangement pitch p _3.

  In addition, it is preferable that the bucky folding operation (the operation of turning from the forward path to the backward path or the reverse operation) of the scattered radiation removal grid 24 is basically not performed during one imaging period. This is because the scattered radiation removal grid 24 is in a localized state during acceleration / deceleration of the folding operation, which may affect the image quality. However, when the imaging time is long, in order not to perform the folding operation, the scattered radiation removal grid 24 must be moved at a very low speed. In this way, the scattered radiation removal grid 24 is moved at a low speed. Control is generally difficult and speed unevenness occurs. In this case, conversely, the speed unevenness affects the image quality. Therefore, when the shooting time is long, the influence on the image quality may be reduced by performing the folding operation without lowering the moving speed of the disturbance removal grid 24. In such a case, during the shooting period, It may be allowed to perform the folding operation.

  By the above-described bucky operation, the X-ray non-irradiation region (shadow) on the detection surface of the FPD 20 is uniformly moved on the detection surface during one imaging period by the scattered radiation removal grid 24. The intensity difference due to the scattered radiation removal grid 24 does not occur. Further, since the direction of the bucky operation is the y direction orthogonal to the scanning direction (x direction) of the second absorption type grating 22, it does not affect the detection of the phase shift amount ψ obtained as a function of x. .

  In this way, during one imaging period performed at each scanning position k, the scattered radiation removal grid 24 is operated as a bucky operation in the y direction, thereby causing a scattering phenomenon inside the subject H as shown in FIG. X-rays (scattered X-rays) deflected regardless of the phase shift distribution Φ (x, y) of the subject H are effectively removed (absorbed) by the X-ray absorber 24a.

  In the X-ray imaging system 10 configured as described above, when an imaging instruction is input from the console 17 by the operator with the subject H placed between the X-ray source 11 and the FPD 20, imaging control is performed. Each unit of the imaging unit 12 is controlled by the unit 16 and the second absorption type grating 22 is moved relative to the first absorption type grating 21 while the X-ray source 11 performs exposure and the FPD 20 detects at each scanning position. Operation is performed. At this time, the reciprocating device 25 performs a bucky operation of the scattered radiation removal grid 24.

  During each exposure period (during imaging period) by the X-ray source 11, a bucky operation is performed in a direction orthogonal to the scanning direction, so that the scattered X-rays can be removed without preventing the detection of the phase shift amount ψ. Is called. As a result, a high-quality phase contrast image is obtained by the image processing unit 14 and displayed on the monitor via the image recording unit 15.

  In the above embodiment, the scattered radiation removal grid 24 is provided between the first absorption type grating 21 and the second absorption type grating 22 that are arranged at a predetermined interval in order to reduce the thickness of the imaging unit 12. However, the present invention is not limited to this, and the scattered radiation removal grid 24 is disposed between the X-ray source 11 and the first absorption grating 21 or between the second absorption grating 22 and the FPD 20. It is also possible to do. From the viewpoint of the effect of scatter removal, it is preferable to dispose the scattered radiation removal grid 24 between the second absorption type grating 22 and the FPD 20, that is, immediately before the FPD 20.

In the above embodiment, the scattered radiation removal grid 24 is operated as a bucky by the reciprocating device 25. However, the scattered radiation removing grid 24 may be fixed without providing the reciprocating device 25. In this case, so as not to be completely blocked by even the X-ray absorbing portion 24a any pixel 40 in the FPD 20, the arrangement pitch p ₃ of the X-ray absorbing portion 24a, than the arrangement pitch in the y direction of the pixel 40 It needs to be shortened. Furthermore, in order to equalize the amount of incident X-ray to each pixel 40, that the integral multiple of the arrangement pitch p ₃ of the X-ray absorbing portion 24a is, so as to match the arrangement pitch in the y direction of the pixel 40 preferable.

  In the case where the reciprocating device 25 is not provided as described above, it is also preferable that the scattered radiation removing grid 24 is integrated with the second absorption type grating 22 to provide a square absorption type second absorption type grating.

  In the above embodiment, when the distance from the X-ray source 11 to the FPD 20 is increased, the blur of the G1 image due to the focus size of the X-ray focus 11a (generally about 0.1 mm to 1 mm) affects the phase. Since there is a risk of degrading the image quality of the contrast image, a multi slit (ray source grid) may be arranged immediately after the X-ray focal point 11a.

  The multi-slit is an absorption type grating having a configuration similar to that of the first and second absorption type gratings 21 and 22, and a plurality of X-ray shielding portions extending in one direction (in this embodiment, y direction) The first and second absorption type gratings 21 and 22 are periodically arranged in the same direction (in the present embodiment, the x direction) as the X-ray shielding portions 21a and 22a. This multi-slit partially shields X-rays from the X-ray source 11 to reduce the effective focal size in the x direction, and forms a large number of point light sources (dispersed light sources) in the x direction. Suppresses the blur of the G1 image.

In the above embodiment, the first absorption grating 21 is configured to linearly project the X-rays that have passed through the slit portion as a G1 image, but the present invention is limited to this configuration. It is good also as a structure as described in international publication WO2004 / 058070 etc. which produce what is called a Talbot interference effect by diffracting X-rays by a slit part instead of a thing. However, in this case, it is necessary to set the distance L ₂ between the first and second absorption gratings 21 and 22 to the Talbot distance. In this case, a phase type grating (phase type diffraction grating) can be used instead of the first absorption type grating 21, and the phase type grating used instead of the first absorption type grating 21 can be used. Forms a G1 image (self-image) generated by the Talbot interference effect at the position of the second absorption type grating 22.

  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 30 μm in terms of Au in the case of the absorption type grating. On the other hand, in the case of a phase type grating, it is set to about 1 μm to 5 μm. The phase-type grating generates a Talbot interference effect by applying a predetermined amount (π or π / 2) of phase modulation to the X-rays irradiated from the X-ray source 11 by the X-ray high-absorbing material. An image (self-image) is generated.

  Further, in the above embodiment, the subject H is disposed between the X-ray source 11 and the first absorption type grating 21, but the subject H is arranged with 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.

  Next, a modified example of the X-ray image detector will be described. In each of the above embodiments, the second absorption type grating 22 and the FPD 20 are provided independently. However, instead of the FPD 20, an X-ray image detector having a configuration disclosed in Japanese Patent Laid-Open No. 2009-133823 is used. By using it, the second absorption type grating 22 can be eliminated. This X-ray image detector is a direct conversion type X-ray image detector including a conversion layer that converts X-rays into electric charges and a charge collection electrode that collects electric charges converted in the conversion layer. The charge collection electrode is configured by arranging a plurality of linear electrode groups formed by electrically connecting linear electrodes arranged at a constant period so that the phases are different from each other. Constitutes the intensity modulating means described in the claims.

  In FIG. 8 showing the configuration of the X-ray image detector of the present modification, a plurality of pixels 60 are two-dimensionally arranged at a constant pitch along the x direction and the y direction. A charge collection electrode 61 is formed for collecting the charges converted by the conversion layer that converts the charges. The charge collection electrode 61 is composed of first to sixth linear electrode groups 62 to 67, and the phase of the arrangement period of the linear electrodes of each linear electrode group is shifted by π / 3. Specifically, when the phase of the first linear electrode group 62 is 0, the phase of the second linear electrode group 63 is π / 3, the phase of the third linear electrode group 64 is 2π / 3, The phase of the fourth linear electrode group 65 is π, the phase of the fifth linear electrode group 66 is 4π / 3, and the phase of the sixth linear electrode group 67 is 5π / 3.

  Further, each pixel 60 is provided with a switch group 68 for reading out the charges collected by the charge collecting electrode 61. The switch group 68 includes TFT switches provided in the first to sixth linear electrode groups 62 to 67, respectively. The charges collected by the first to sixth linear electrode groups 62 to 67 are individually read out by controlling the switch group 68, thereby detecting six types of G2 images having different phases from each other by one imaging. can do. A phase differential image can be generated based on a plurality of image data corresponding to the six types of G2 images, and a phase contrast image can be generated.

  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 a plurality of G2 images that have been intensity-modulated with different phases can be detected by one imaging, physical scanning for fringe scanning becomes unnecessary, and the scanning mechanism 23 can be eliminated. Instead of the charge collection electrode 61 shown in the figure, it is also possible to use a charge collection electrode having another configuration described in JP-A-2009-133823.

  Further, as another embodiment in which the second absorption type grating 22 is not disposed, the G1 image is directly detected by an X-ray image detector, and the phase is mutually sampled by periodically sampling while changing the phase by signal processing. It is also possible to generate image data corresponding to a plurality of different G2 images.

  Further, in each of the above embodiments, the phase differential image is obtained by the fringe scanning method. However, the present invention is not limited to this, and the phase differential image may be obtained by the Fourier transform method described in International Publication WO2010 / 050833. Good. This Fourier transform method obtains a Fourier spectrum of moire fringes generated in image data by Fourier transforming one piece of image data obtained by an X-ray image detector, and a spectrum corresponding to the carrier frequency from this Fourier spectrum. This is a method of obtaining a phase differential image by performing inverse Fourier transform by separating. In this case, it is not necessary to move the first and second absorption gratings 21 and 22, and the scanning mechanism 23 is not necessary.

  The first and second embodiments described above can be applied to other radiation imaging systems for industrial use in addition to the radiation imaging system for medical diagnosis. In addition to X-rays, gamma rays or the like can be used as radiation.

10 X-ray imaging system 11 X-ray source (radiation source)
11a X-ray focus 12 Imaging unit 14 Image processing unit 16 Imaging control unit 18 System control unit 20 Flat panel detector (FPD)
DESCRIPTION OF SYMBOLS 21 1st absorption-type grating | lattice 21a X-ray shielding part 22 2nd absorption-type grating | lattice 22a X-ray shielding part 23 Scanning mechanism 24 Scattering ray removal grid 24a X-ray absorption part 24b X-ray transmission part 25 Reciprocating motion apparatus 26 Phase differential image Generation unit 27 Phase contrast image generation unit 40 pixels 60 pixels 61 charge collection electrodes 62 to 67 first to sixth linear electrode groups 68 switch groups

Claims (9)

A first grating that generates a first periodic pattern image by passing the radiation emitted from a radiation source;
Intensity modulation means for applying intensity modulation to the first periodic pattern image to generate a second periodic pattern image;
A radiation image detector for detecting the second periodic pattern image and generating image data;
A scattered radiation elimination grid disposed between the radiation source and the radiation image detector so that a grating pattern is orthogonal to the grating pattern of the first grating;
Phase differential image generating means for generating a phase differential image based on the image data;
A radiation imaging system comprising:   The radiation imaging system according to claim 1, further comprising reciprocating means for reciprocating the scattered radiation removing grid in a direction orthogonal to a lattice pattern of the first lattice.   The radiation imaging system according to claim 1, wherein the scattered radiation removal grid is a focusing grid having directivity with respect to a focal point of the radiation source.   The phase contrast image generation means which produces | generates a phase contrast image by performing an integration process with respect to the image data produced | generated by the said radiographic image detector, The any one of Claim 1 to 3 characterized by the above-mentioned. Radiography system. The intensity modulation means applies intensity modulation at a plurality of relative positions having different phases with respect to the first periodic pattern image to generate a plurality of second periodic pattern images;
The radiation image detector detects the second periodic pattern images to generate a plurality of image data;
The phase differential image generation unit generates the phase differential image by calculating a phase shift amount of an intensity modulation signal representing an intensity change of pixel data with respect to the relative position based on the plurality of image data. The radiation imaging system according to any one of claims 1 to 4.   The intensity modulation means includes: a second grating having a periodic pattern in the same direction as the first periodic pattern image; and scanning means for moving one of the first and second gratings at a predetermined pitch. The radiation imaging system according to claim 1, wherein:   The radiation imaging system according to claim 6, wherein the scattered radiation removal grid is disposed between the first grating and the second grating.   The first grating is a phase grating, and the first grating forms radiation from the radiation source as the first periodic pattern image at the position of the second grating by a Talbot interference effect. The radiation imaging system according to claim 6 or 7. The radiological image detector is a radiological image detector comprising a conversion layer for converting radiation into electric charge and a charge collecting electrode for collecting electric charge converted in the conversion layer for each pixel,
The charge collection electrode is formed by arranging a plurality of linear electrode groups having a periodic pattern in the same direction as the first periodic pattern image so that the phases thereof are different from each other.
The radiation imaging system according to claim 1, wherein the intensity modulation unit includes the charge collection electrode.

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