WO2012133553A1 - Radiography system and radiography method - Google Patents

Abstract

The radiography system (10) generates a differential phase image of a subject by detecting a radiological image, which is formed by radiation that has passed through a first grating (31) and been modulated by passing through the subject, and is masked by a second grating (32). The relative positions of the first grating and the second grating are adjusted to generate, in the radiological image, a moire pattern with a period in a direction that intersects with the direction of the period of the first grating, and using a filter that acts on image contrast changes in the direction of the period of the moire pattern, nonuniformities in the phase differential image are removed or reduced.

Description

Radiographic system and radiographic method

The present invention relates to a radiation imaging system and a radiation imaging method.

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 an X-ray image, and a transmission image of the subject is captured. In this case, each X-ray radiated from the X-ray source toward the X-ray image detector has characteristics (atomic number, density, thickness) of the substance constituting the subject existing on the path to the X-ray image detector. ), The light is incident on the X-ray image detector. As a result, an X-ray transmission image of the subject is detected and imaged by the X-ray image detector. As an X-ray image detector, there is a flat panel detector (FPD: Flat Panel Detector) using a semiconductor circuit in addition to a combination of an X-ray intensifying screen and a film, a stimulable phosphor (accumulating phosphor), and so on. Widely used.

However, the X-ray absorptivity becomes lower as a substance composed of an element having a smaller atomic number, and the difference in the X-ray absorptivity is small in a soft tissue or soft material of a living body. Therefore, a sufficient image density as an X-ray transmission image is obtained. There is a problem that (contrast) cannot be obtained. 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 the difference in the amount of X-ray absorption between the two is small, so that it is difficult to obtain image contrast.

Against the background of such problems, in recent years, an X-ray phase for obtaining an image (hereinafter referred to as a phase contrast image) based on an X-ray phase change (angle change) by an object instead of an X-ray intensity change by an object. Imaging research is actively conducted. In general, it is known that when X-rays are incident on an object, the interaction is higher in phase than in X-ray intensity. For this reason, in the X-ray phase imaging using the phase difference, a high-contrast image can be obtained even for a weakly absorbing object having a low X-ray absorption capability. As a kind of such X-ray phase imaging, in recent years, an X-ray imaging system using an X-ray Talbot interferometer comprising two transmission diffraction gratings (phase grating and absorption grating) and an X-ray image detector has been proposed. It has been devised (for example, see Patent Document 1).

In the X-ray Talbot interferometer, a first diffraction grating (phase type grating or absorption type grating) is arranged behind a subject, and a specific distance (Talbot interference distance) determined by the grating pitch of the first diffraction grating and the X-ray wavelength. The second diffraction grating (absorption type grating) is disposed only downstream, and the X-ray image detector is disposed behind the second diffraction grating. The Talbot interference distance is a distance at which X-rays that have passed through the first diffraction grating form a self-image due to the Talbot interference effect, and this self-image is between the X-ray source and the first diffraction grating. It is modulated by the interaction (phase change) between the arranged subject and the X-ray. Then, the phase information of the subject is obtained by superimposing the second diffraction grating on the modulated self-image and analyzing the modulation of the self-image.

For example, a fringe scanning method is known as a method for analyzing the modulation of the self-image of the first diffraction grating. In the fringe scanning method, the second diffraction grating superimposed on the self-image is photographed a plurality of times while being translated at a scanning pitch obtained by equally dividing the grating pitch, and each pixel between the obtained images is obtained. An X-ray refraction angle distribution with respect to the periodic direction of the first diffraction grating is obtained from the change in signal value. This refraction angle distribution corresponds to the differentiation of the phase shift distribution of the subject, and a phase contrast image of the subject can be obtained based on this refraction angle distribution.

By the way, due to an error in the grating pitch of the first and second diffraction gratings, a deviation from a predetermined relative positional relationship, and the like, the superposition of the self-image of the first diffraction grating and the second diffraction grating is Usually, moire fringes. Depending on the period of the moire fringes, the influence may remain in the differential phase image (the distribution image of the refraction angle or the differential image of the phase shift distribution) and appear as unevenness. Such image unevenness is typically removed or reduced using a filter (see, for example, Patent Document 2).

International Publication No. 2004/058070 Japanese Patent Laid-Open No. 3-12785

The above-described X-ray Talbot interferometer has sensitivity to the X-ray phase change in the periodic direction of the first diffraction grating. If a filter for removing or reducing unevenness is applied to the phase contrast image without considering the sensitivity direction, the phase information may be affected.

The present invention has been made in view of the above-described circumstances, and an object thereof is to obtain a phase contrast image in which unevenness is removed or reduced while suppressing the influence on phase information.

(1) a first grating, a second grating that masks a radiation image that is formed by the radiation that has passed through the first grating and that has been transmitted through the subject and is modulated, and the second grating. A radiation image detector that detects the masked radiation image; and an arithmetic processing unit that generates a differential phase image of the subject based on the radiation image acquired by the radiation image detector. The second grating and the second grating are arranged in a relative positional relationship in which a moire fringe having a period in a direction intersecting with a period direction of the first grating is generated in the radiation image, and the arithmetic processing unit includes: A radiation imaging system that performs a filtering process for removing or reducing unevenness of the differential phase image using a filter that affects a change in image contrast in the periodic direction of the moire fringes.
(2) Radiation that generates a differential phase image of the subject by detecting a radiation image that has passed through the first grating and that has been modulated by passing through the subject and is masked by the second grating. An imaging method, wherein the relative positional relationship between the first grating and the second grating is adjusted, and moire fringes having a period in a direction intersecting with a period direction of the first grating are generated in the radiation image. A radiation imaging method for removing or reducing unevenness of the phase differential image using a filter that acts on a change in image contrast in the period direction of the moire fringes.

According to the present invention, the periodic direction of the moire fringes is set in a direction intersecting with the periodic direction of the first grating, and unevenness appearing in the differential phase image due to the moire fringes is also the period of the first grating. It occurs in a direction that intersects the direction. Then, by using a filter that affects the image contrast change in the periodic direction of the moire fringes, that is, the direction intersecting the periodic direction of the first grating, unevenness is removed or reduced without affecting the phase information. Can do.

It is a schematic diagram which shows the structure of an example of the radiography system for describing embodiment of this invention. It is a control block diagram of the radiography system of FIG. It is a schematic diagram which shows the structure of the radiographic image detector of the radiography system of FIG. It is a perspective view of the imaging part of the radiography system of FIG. It is a side view of the imaging part of the radiography system of FIG. It is a schematic diagram which shows the mechanism for changing the period of the moire fringe by superimposition of the 1st and 2nd grating | lattice. It is a schematic diagram for demonstrating the refraction | bending of the radiation by a to-be-photographed object. It is a schematic diagram for demonstrating the fringe scanning method. It is a graph which shows the change of the signal value of the pixel of the radiographic image detector accompanying a fringe scanning. An example of the nonuniformity which appears in a differential phase image by a moire fringe and this moire fringe is shown. It is a schematic diagram which shows an example of the spatial frequency spectrum of a differential phase image. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention.

FIG. 1 shows a configuration of an example of a radiation imaging system for explaining an embodiment of the present invention, and FIG. 2 shows a control block of the radiation imaging system of FIG.

The X-ray imaging system 10 is an X-ray diagnostic apparatus that images a subject (patient) H in a standing position, and is disposed opposite to the X-ray source 11 that emits X-rays to the subject H, and the X-ray source 11. An imaging unit 12 that detects X-rays transmitted through the subject H from the X-ray source 11 and generates image data, and controls the exposure operation of the X-ray source 11 and the imaging operation of the imaging unit 12 based on the operation of the operator. At the same time, it is roughly divided into a console 13 that generates a phase contrast image by calculating the image data acquired by the photographing unit 12.

The X-ray source 11 is held movably in the vertical direction (x direction) by an X-ray source holding device 14 suspended from the ceiling. The photographing unit 12 is held by a standing stand 15 installed on the floor so as to be movable in the vertical direction.

Based on the control of the X-ray source control unit 17, the X-ray source 11 is emitted from the X-ray tube 18 that generates X-rays according to the high voltage applied from the high voltage generator 16, and the X-ray tube 18. The X-ray includes a collimator unit 19 including a movable collimator 19a that limits an irradiation field so as to shield a portion that does not contribute to the inspection area of the subject H. The X-ray tube 18 is of an anode rotating type, and emits an electron beam from a filament (not shown) as an electron emission source (cathode) and collides with a rotating anode 18a rotating at a predetermined speed, thereby causing X-rays. Is generated. The colliding portion of the rotating anode 18a with the electron beam becomes the X-ray focal point 18b.

The X-ray source holding device 14 includes a carriage portion 14a configured to be movable in a horizontal direction (z direction) by a ceiling rail (not shown) installed on the ceiling, and a plurality of support column portions 14b connected in the vertical direction. It consists of. A motor (not shown) that changes the position of the X-ray source 11 in the vertical direction is provided on the carriage unit 14 a by expanding and contracting the column unit 14 b.

The standing stand 15 includes a main body 15a installed on the floor, and a holding portion 15b that holds the photographing unit 12 is attached to be movable in the vertical direction. The holding portion 15b is connected to an endless belt 15d that is suspended between two pulleys 15c that are spaced apart in the vertical direction, and is driven by a motor (not shown) that rotates the pulley 15c. The driving of the motor is controlled by the control device 20 of the console 13 described later based on the setting operation by the operator.

Further, the standing stand 15 is provided with a position sensor (not shown) such as a potentiometer that detects the position of the photographing unit 12 in the vertical direction by measuring the movement amount of the pulley 15c or the endless belt 15d. . The detection value of this position sensor is supplied to the X-ray source holding device 14 by a cable or the like. The X-ray source holding device 14 moves the X-ray source 11 so as to follow the vertical movement of the imaging unit 12 by expanding and contracting the support column 14 b based on the supplied detection value.

The console 13 is provided with a control device 20 comprising a CPU, ROM, RAM and the like. The control device 20 includes an input device 21 through which an operator inputs an imaging instruction and the content of the instruction, an arithmetic processing unit 22 that performs arithmetic processing on the image data acquired by the imaging unit 12 and generates an X-ray image, and X A storage unit 23 for storing line images, a monitor 24 for displaying X-ray images and the like, and an interface (I / F) 25 connected to each unit of the X-ray imaging system 10 are connected via a bus 26. .

As the input device 21, for example, a switch, a touch panel, a mouse, a keyboard, or the like can be used. By operating the input device 21, X-ray imaging conditions such as X-ray tube voltage and X-ray irradiation time, imaging timing, etc. Is entered. The monitor 24 includes a liquid crystal display or the like, and displays characters such as X-ray imaging conditions and X-ray images under the control of the control device 20.

The imaging unit 12 includes a flat panel detector (FPD) 30 made of a semiconductor circuit, a first absorption type grating 31 and a second absorption type for detecting phase change (angle change) of X-rays by the subject H and performing phase imaging. The absorption type grating 32 is provided.

The FPD 30 is arranged so that the detection surface is orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11. Although described in detail later, the first and second absorption gratings 31 and 32 are disposed between the FPD 30 and the X-ray source 11.

The imaging unit 12 changes the relative positional relationship of the second absorption type grating 32 with respect to the first absorption type grating 31 by translating the second absorption type grating 32 in the vertical direction (x direction). A scanning mechanism 33 is provided. The scanning mechanism 33 is configured by an actuator such as a piezoelectric element, for example.

FIG. 3 shows a configuration of a radiation image detector included in the radiation imaging system of FIG.

The FPD 30 as a radiological image detector includes an image receiving unit 41 in which a plurality of pixels 40 that convert X-rays into electric charges and store them in a two-dimensional array on an active matrix substrate, and an electric charge received from the image receiving unit 41. A scanning circuit 42 that controls the readout timing, a readout circuit 43 that reads out the charges accumulated in each pixel 40, converts the charges into image data and stores them, and performs arithmetic processing on the image data via the I / F 25 of the console 13. And a data transmission circuit 44 for transmission to the unit 22. The scanning circuit 42 and each pixel 40 are connected by a scanning line 45 for each row, and the readout circuit 43 and each pixel 40 are connected by a signal line 46 for each column.

Each pixel 40 is a direct conversion type in which X-rays are directly converted into electric charges by a conversion layer (not shown) such as amorphous selenium and the converted electric charges are stored in a capacitor (not shown) connected to the lower electrode. It can be configured as an X-ray detection element. Each pixel 40 is connected to a thin film transistor (TFT) switch (not shown), and the gate electrode of the TFT switch is connected to the scanning line 45, the source electrode is connected to the capacitor, and the drain electrode is connected to the signal line 46. 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 46.

Each pixel 40 once converts X-rays into visible light by a scintillator (not shown) made of terbium activated gadolinium oxide (Gd ₂ O ₂ S: Tb), thallium activated cesium iodide (CsI: Tl), or the like. It is also possible to configure as an indirect conversion type X-ray detection element that converts the converted visible light into a charge by a photodiode (not shown) and accumulates it. The X-ray image detector is not limited to an FPD based on a TFT panel, and various X-ray image detectors based on a solid-state imaging device such as a CCD sensor or a CMOS sensor can also be used.

The readout circuit 43 includes an integration amplifier circuit, an A / D converter, a correction circuit, and an image memory. The integrating amplifier circuit integrates the charges output from each pixel 40 via the signal line 46, converts them into a voltage signal (image signal), and inputs it to the A / D converter. The A / D converter converts the input image signal into digital image data and inputs the digital image data to the correction circuit. The correction circuit performs offset correction, gain correction, and linearity correction on the image data, and stores the corrected image data in the image memory. As correction processing by the correction circuit, correction of X-ray exposure amount and exposure distribution (so-called shading) and pattern noise depending on FPD 30 control conditions (drive frequency and readout period) (for example, leak signal of TFT switch) May be included.

4 and 5 show an imaging unit of the radiation imaging system of FIG.

The first absorption-type grating 31 includes a substrate 31a and a plurality of X-ray shielding portions 31b arranged on the substrate 31a. Similarly, the second absorption type grating 32 includes a substrate 32a and a plurality of X-ray shielding portions 32b arranged on the substrate 32a. The substrates 31a and 32a are both made of an X-ray transparent member such as glass that transmits X-rays.

Each of the X-ray shielding portions 31b and 32b is in one direction in a plane orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11 (in the illustrated example, the y direction orthogonal to the x direction and the z direction). It is comprised by the linear member extended | stretched. As a material of each X-ray shielding part 31b, 32b, a material excellent in X-ray absorption is preferable, and for example, a heavy metal such as gold or platinum is preferable. These X-ray shielding portions 31b and 32b can be formed by a metal plating method or a vapor deposition method.

X-ray shielding portion 31b is in a plane perpendicular to the optical axis A of the X-ray, at a predetermined period p ₁ in a direction (x-direction) orthogonal to the one direction, are arranged at a predetermined interval d ₁ from each other ing. Similarly, X-ray shielding portion 32b, in the plane orthogonal to the optical axis A of the X-ray, at a predetermined period p ₂ in a direction (x-direction) orthogonal to the one direction, at a predetermined interval d ₂ from each other Are arranged. Since the first and second absorption gratings 31 and 32 do not give a phase difference to incident X-rays but give an intensity difference, they are also called amplitude gratings. Note that the slit portions (regions having the distances d ₁ and d ₂ ) may not be voids, and the voids may be filled with an X-ray low-absorbing material such as a polymer or a light metal.

The first and second absorption gratings 31 and 32 are configured to geometrically project the X-rays that have passed through the slit portion regardless of the presence or absence of the Talbot interference effect. 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 rotary anode 18a described above 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 geometrically projected without being diffracted at the slit portion.

The X-ray emitted from the X-ray source 11 is not a parallel beam but a cone beam having the X-ray focal point 18b as a light emission point, and therefore a projected image projected through the first absorption grating 31 (hereinafter referred to as a projection image). The projection image is referred to as a G1 image) and is enlarged in proportion to the distance from the X-ray focal point 18b. The grating pitch p ₂ of the second absorption type grating 32 is determined so that the slit portion substantially coincides with the periodic pattern of the bright part of the G1 image at the position of the second absorption type grating 32. That is, when the distance from the X-ray focal point 18b to the first absorption grating 31 is L ₁ and the distance from the first absorption grating 31 to the second absorption grating 32 is L ₂ , the grating pitch p ₂ is determined so as to satisfy the relationship of the following formula (1).

In the Talbot interferometer, the distance L ₂ from the first absorption type grating 31 to the second absorption type grating 32 is limited to the Talbot interference distance determined by the grating pitch of the first diffraction grating and the X-ray wavelength. The imaging unit 12 of the present X-ray imaging system 10 has a configuration in which the first absorption grating 31 projects incident X-rays without diffracting, and the G1 image of the first absorption grating 31 is the first. because at every position of the rear absorption type grating 31 similarly obtained, the distance L _2, can be set independently of the Talbot distance.

As described above, the imaging unit 12 does not constitute a Talbot interferometer, but the Talbot interference distance Z when it is assumed that X-rays are diffracted by the first absorption type grating 31 is the first absorption type grating. the grating pitch _{p 1} of 31, the grating pitch _p 2, X-ray wavelength of the second absorption-type grating 32 (peak wavelength) lambda, and using the positive integer m, is expressed by the following equation (2).

Formula (3) is a formula that represents the Talbot interference distance when the X-rays emitted from the X-ray source 11 are cone beams. “Atsushi Momose, et al., Japan Journal of Applied Physics, Vol. 47, No. 10, October 2008, page 8077 ”.

In the present X-ray imaging system 10, the imaging unit 12 is thinned by setting the distance L ₂ to a value shorter than the minimum Talbot interference distance Z when m = 1. That is, the distance L ₂ is set to a value in the range satisfying the following equation (3).

Incidentally, Talbot distance Z by the following equation (4) and in the case of X-rays emitted from the X-ray source 11 can be regarded as substantially parallel beams, the distance L _2, the value of the range that satisfies the following equation (5) Set to.

However, not always the distance L ₂ need not satisfy equation (3) through (5), for example, when there is no demand for thinning of the imaging unit 12, from the equation (3) to (5) Values outside the range can also be taken.

The X-ray shielding portions 31b and 32b preferably completely shield (absorb) X-rays in order to generate a periodic pattern image with high contrast, but the above-described materials (gold, platinum) having excellent X-ray absorption properties Etc.), there are not a few X-rays that are transmitted without being absorbed. Therefore, in order to enhance the shielding of the X-rays, the X-ray shielding portion 31b, the respective thicknesses _h 1, _{h 2} of 32b, it is preferable to increase the thickness much as possible. For example, when the tube voltage of the X-ray tube 18 is 50 kV, it is preferable to shield 90% or more of the irradiated X-rays. In this case, the thicknesses h ₁ and h ₂ are 30 μm or more in terms of gold (Au). It is preferable that

However, when the X-rays irradiated from the X-ray source 11 are cone beams, if the thicknesses h ₁ and h ₂ of the X-ray shielding portions 31b and 32b are too thick, the X-rays incident obliquely enter the slit portion. There is a problem that it becomes difficult to pass, so-called vignetting occurs, and the effective visual field in the direction (x direction) perpendicular to the extending direction (strand direction) of the X-ray shielding portions 31b and 32b becomes narrow. Therefore, in view of the field of view secured to define the upper limit of the thickness _h 1, _{h 2.} In order to secure the effective field length V in the x direction on the detection surface of the FPD 30, assuming that the distance from the X-ray focal point 18 b to the detection surface of the FPD 30 is L, the thicknesses h ₁ and h ₂ are shown in FIG. From the scientific relationship, it is necessary to set so as to satisfy the following expressions (6) and (7).

For example, when d ₁ = 2.5 μm and d ₂ = 3.0 μm, and assuming L = 2 m assuming normal hospital imaging, the effective visual field length V in the x direction is 10 cm. In order to ensure the length, the thickness h ₁ may be 100 μm or less and the thickness h ₂ may be 120 μm or less.

In the imaging unit 12 configured as described above, an intensity-modulated image is formed by superimposing the G1 image of the first absorption-type grating 31 and the second absorption-type grating 32 and is captured by the FPD 30. . The pattern period p ₁ ′ of the G1 image at the position of the second absorption grating 32 and the substantial grating pitch p ₂ ′ (substantial pitch after production) of the second absorption grating 32 are manufacturing errors. Some differences occur due to or placement errors. Among these, the arrangement error means that the substantial pitch in the x direction changes due to the relative inclination and rotation of the first and second absorption gratings 31 and 32 and the distance between the two changes. I mean.

Due to the minute difference between the pattern period p ₁ ′ of the G1 image and the grating pitch p ₂ ′, the image contrast becomes moire fringes. The period T of the moire fringes is expressed by the following equation (8).

In order to detect the moire fringes by the FPD 30, the arrangement pitch P in the x direction of the pixels 40 needs to be at least not an integral multiple of the moire period T, and it is necessary to satisfy the following equation (9) (where n Is a positive integer).

In addition, it is possible to detect moire fringes even if the arrangement pitch P is larger than the moire period T within a range satisfying the expression (9), but the arrangement pitch P is preferably smaller than the moire period T. 10) is preferably satisfied. This is because, in order to obtain a high-quality phase contrast image, moire fringes are preferably detected with high contrast in the phase contrast image generation process described later.

Since the arrangement pitch P of the pixels 40 of the FPD 30 is a value determined by design (generally about 100 μm) and is difficult to change, the magnitude relationship between the arrangement pitch P and the moire period T is adjusted. Adjusts the positions of the first and second absorption gratings 31 and 32 and changes the moire period T by changing at least one of the pattern period p ₁ ′ and the grating pitch p ₂ ′ of the G1 image. It is preferable to do.

FIG. 6 shows a method of changing the moire cycle T.

The moire period T can be changed by relatively rotating one of the first and second absorption gratings 31 and 32 around the optical axis A. For example, a relative rotation mechanism 50 that rotates the second absorption grating 32 relative to the first absorption grating 31 relative to the optical axis A is provided. When the second absorption type grating 32 is rotated by the angle θ by the relative rotation mechanism 50, the substantial grating pitch in the x direction changes from “p ₂ ′” → “p ₂ ′ / cos θ”. As a result, the moire cycle T changes (FIG. 6A).

As another example, the change of the moire period T is such that either one of the first and second absorption type gratings 31 and 32 is relatively centered about an axis perpendicular to the optical axis A and along the y direction. It can be performed by inclining. For example, a relative tilt mechanism 51 that tilts the second absorption type grating 32 relative to the first absorption type grating 31 about an axis perpendicular to the optical axis A and along the y direction is provided. Provide. When the second absorption type grating 32 is inclined by the angle α by the relative inclination mechanism 51, the substantial lattice pitch in the x direction changes from “p ₂ ′” → “p ₂ ′ × cos α”. As a result, the moire cycle T changes (FIG. 6B).

As another example, the moire period T can be changed by relatively moving one of the first and second absorption gratings 31 and 32 along the direction of the optical axis A. For example, with respect to the first absorption type grating 31, the second absorption type grating 32 is changed so as to change the distance L ₂ between the first absorption type grating 31 and the second absorption type grating 32. A relative movement mechanism 52 that relatively moves along the direction of the optical axis A is provided. When the second absorption type grating 32 is moved to the optical axis A by the movement amount δ by the relative movement mechanism 52, the G1 image of the first absorption type grating 31 projected onto the position of the second absorption type grating 32. The pattern period of “p ₁ ′” → “p ₁ ′ × (L ₁ + L ₂ + δ) / (L ₁ + L ₂ )” changes, and as a result, the moire period T changes (FIG. 6C).

In the X-ray imaging system 10, imaging unit 12 is not the Talbot interferometer as described above, since the distance L ₂ can be freely set, moire by changing the distance L ₂ as relative movement mechanism 52 A mechanism for changing the period T can be suitably employed. The change mechanism (relative rotation mechanism 50, relative tilt mechanism 51, and relative movement mechanism 52) of the first and second absorption gratings 31 and 32 for changing the moiré period T is constituted by an actuator such as a piezoelectric element. Is possible.

When the subject H is disposed between the X-ray source 11 and the first absorption type grating 31, the moire fringes detected by the FPD 30 are 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, the phase contrast image of the subject H can be generated by analyzing the moire fringes detected by the FPD 30.

Next, a method for analyzing moire fringes will be described.

FIG. 7 shows one X-ray refracted according to the phase shift distribution Φ (x) of the subject H in the x direction.

Reference numeral 55 indicates an X-ray path that travels straight when the subject H is not present. The X-ray that travels along the path 55 passes through the first and second absorption gratings 31 and 32 and enters the FPD 30. To do. Reference numeral 56 indicates an X-ray path refracted and deflected by the subject H when the subject H exists. X-rays traveling along this path 56 are shielded by the second absorption type grating 32 after passing through the first absorption type grating 31.

The phase shift distribution Φ (x) of the subject H is expressed by the following equation (11), 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 absorptive grating 31 to the position of the second absorptive grating 32 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. become. This amount of displacement Δx is approximately expressed by the following equation (12) based on the small X-ray refraction angle φ.

Here, the refraction angle φ is expressed by Expression (13) 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 the refraction of X-rays at the subject H is related to the phase shift distribution Φ (x) of the subject H. This displacement amount Δx is the signal of each pixel of the image data output from the FPD 30, that is, the phase shift amount ψ of the signal of each pixel 40 (the signal of each pixel 40 with and without the subject H) (Phase shift amount) is related to the following equation (14).

Therefore, by obtaining the phase shift amount ψ of the signal of each pixel 40, the refraction angle φ is obtained from the equation (14), and the differential amount of the phase shift distribution Φ (x) is obtained using the equation (13). Is integrated with respect to x, a phase shift distribution Φ (x) of the subject H, that is, a phase contrast image of the subject H can be generated. In the present X-ray imaging system 10, the phase shift amount ψ is calculated using a fringe scanning method described below.

In the fringe scanning method, imaging is performed while one of the first and second absorption type gratings 31 and 32 is translated in a stepwise manner in the x direction relative to the other (that is, the phase of both grating periods is changed). Shoot while changing). In the X-ray imaging system 10, the second absorption type grating 32 is moved by the scanning mechanism 33 described above, but the first absorption type grating 31 may be moved.

As the second absorption type grating 32 moves, the moire fringes move, and the translation distance (the amount of movement in the x direction) is one period of the grating period of the second absorption type grating 32 (grating pitch p ₂ ). (Ie, when the phase change reaches 2π), the moire fringes return to their original positions. A fringe image is photographed by the FPD 30 while moving the second absorption grating 32 by an integer of the grating pitch p ₂ , and signals of each pixel 40 are acquired from the photographed plural fringe images. By performing arithmetic processing, the phase shift amount ψ of the signal of each pixel 40 is obtained.

FIG. 8 schematically shows how the second absorption grating 32 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 33 translates the second absorption type grating 32 in order to M scanning positions of k = 0, 1, 2,..., M−1. In the same figure, the initial position of the second absorption grating 32 is the same as the dark part of the G1 image at the position of the second absorption grating 32 when the subject H is not present. The initial position is k = 0, 1, 2,..., M−1.

First, at the position of k = 0, X-rays that are not refracted by the subject H mainly pass through the second absorption type grating 32. Next, when the second absorption grating 32 is moved in order of k = 1, 2,..., The X-rays passing through the second absorption grating 32 are not refracted by the subject H. While the line component decreases, the X-ray component refracted by the subject H increases. In particular, at k = M / 2, mainly only the X-rays refracted by the subject H pass through the second absorption type grating 32. When k = M / 2 is exceeded, on the contrary, the X-ray component that is refracted by the subject H decreases in the X-rays that pass through the second absorption grating 32, while the X-ray that is not refracted by the subject H. The line component increases.

When shooting is performed by the FPD 30 at each position of k = 0, 1, 2,..., M−1, M signal values are obtained for each pixel 40. Hereinafter, a method of calculating the phase shift amount ψ of the signal of each pixel 40 from the M signal values will be described. When the signal value of each pixel 40 at the position k of the second absorption type grating 32 is denoted as I _k (x), I _k (x) is expressed by the following equation (15).

Here, x is a coordinate in the x direction of the pixel 40, A ₀ is the intensity of the incident X-ray, and _An is a value corresponding to the contrast of the signal value of the pixel 40 (where n is a positive value). Is an integer). 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 (16), the refraction angle φ (x) is expressed as the following expression (17).

Here, arg [] means the extraction of the declination, and corresponds to the phase shift amount ψ of the signal of each pixel 40. Accordingly, the refraction angle φ (x) is obtained by calculating the phase shift amount ψ of the signal of each pixel 40 from the M signal values obtained at each pixel 40 based on the equation (17).

FIG. 9 shows a change in the signal value of the pixel 40 due to the fringe scanning.

The M signal values obtained in each pixel 40 periodically change with a period of the grating pitch p _{2 with} respect to the position k of the second absorption grating 32. A broken line in FIG. 9 indicates a change in signal value when the subject H does not exist, and a solid line in FIG. 9 indicates a change in signal value when the subject H exists. The phase difference between the two waveforms corresponds to the phase shift amount ψ of the signal of each pixel 40.

The refraction angle φ (x) is obtained from the phase shift amount ψ of the signal of each pixel 40 based on the equation (17), and the refraction angle φ (x) is a phase shift as shown in the equation (13). Since the value corresponds to the differential amount of the distribution Φ (x), the phase shift distribution Φ (x) can be obtained by integrating the refraction angle φ (x) along the x-axis. 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 Φ (x , Y).

The phase shift distribution Φ (x, y) has been described as a phase contrast image. However, the differential amount of the phase shift distribution Φ and the refraction angle φ corresponding to the phase shift distribution Φ are also the X-ray phase of the subject. Related to change. Therefore, the differential amount of the phase shift distribution Φ and the two-dimensional distribution of the refraction angle φ may be used as the phase contrast image. Hereinafter, an image obtained by using a differential amount of the phase shift distribution Φ and a two-dimensional distribution of the refraction angle φ as an image is referred to as a differential phase image.

FIG. 10 shows an example of moire fringes and unevenness appearing in the differential phase image due to the moire fringes.

In the differential phase image, unevenness may occur due to the influence of moire fringes generated by the superposition of the G1 image of the first absorption type grating 31 and the second absorption type grating 32. The uneven gradient almost coincides with the periodic direction of the moire fringes.

The present X-ray imaging system 10 has sensitivity to the phase change of X-rays in the periodic direction (x direction) of the first absorption type grating 31 so that the periodic direction of moire fringes intersects the x direction. The relative positional relationship between the first and second absorption type gratings is adjusted. By setting the periodic direction of the moire fringes in a direction that intersects the x direction, a gradient of unevenness that appears in the differential phase image also occurs in the direction that intersects the x direction.

In the example shown in the figure, the periodic direction of the moire fringes is set to a substantially y direction that is substantially orthogonal to the x direction (FIG. 10A), and a nonuniformity gradient also occurs in the substantially y direction (FIG. 10B). The periodic direction of the moire fringe is rotated with respect to the first absorption type grating 31 around the optical axis A by using, for example, the relative rotation mechanism 50 (see FIG. 6) described above. Can be changed as appropriate.

FIG. 11 shows a spatial frequency spectrum of the differential phase image of FIG.

In the present X-ray imaging system 10, unevenness appearing in the differential phase image is removed or reduced using a frequency filter. First, the spatial frequency spectrum is acquired by performing Fourier transform on the differential phase image. The spatial frequency spectrum of the differential phase image includes a frequency component 60 derived from phase information (phase shift distribution of the subject H) and a frequency component 61 derived from unevenness. As described above, by generating a nonuniformity gradient in a direction crossing the x direction, the frequency component 60 of the phase information and the nonuniformity frequency component 61 can be separated in the frequency space.

The frequency filter F that masks the separated uneven frequency component 61 is applied to the spatial frequency spectrum of the differential phase image to remove or reduce the uneven frequency component 61. Such a frequency filter also acts on a change in image contrast in the periodic direction of moire fringes, but since the frequency component 60 of phase information and the frequency component 61 of unevenness are separated, the frequency filter described above is The influence on the frequency component 60 is suppressed.

In the illustrated example, the periodic direction of the moire fringes is set to a substantially y direction that is substantially orthogonal to the x direction, and the uneven gradient is also generated in the substantially y direction. In the frequency space, the periodic change in the y direction occurs. The uneven frequency component 61 is located on or near the corresponding v-axis. On the other hand, since it does not have sensitivity to the X-ray phase change in the y direction, the frequency component 60 of the phase information is located off the v-axis. Thus, by setting the periodic direction of the moire fringes in a direction orthogonal to the phase sensitivity direction, the frequency component 60 of the phase information and the frequency component 61 of the unevenness can be reliably separated, and the phase information The uneven frequency component 61 can be removed or reduced without affecting the frequency component 60.

Then, an inverse Fourier transform is performed on the spatial frequency spectrum from which the uneven frequency component 60 has been removed or reduced to restore the differential phase image. Thereby, a differential phase image from which unevenness is removed or reduced is obtained.

The above calculation is performed by the calculation processing unit 22, and the calculation processing unit 22 causes the storage unit 23 to store the phase contrast image from which unevenness has been removed or reduced. The above-described fringe scanning and phase contrast image generation processing is automatically performed after the imaging instruction is given by the operator from the input device 21, and the respective units are linked and operated based on the control of the control device 20. The phase contrast image of the subject H is displayed on the monitor 24.

As described above, the periodic direction of the moire fringes is set in a direction intersecting with the periodic direction of the first absorption type grating 31, and unevenness appearing in the differential phase image due to the moire fringes is also the first. This occurs in a direction crossing the periodic direction of the absorption type grating 31. Then, by using a filter that affects the image contrast change in the periodic direction of the moire fringes, that is, in the direction intersecting with the periodic direction of the first absorption type grating 31, unevenness can be removed without affecting the phase information. Can be reduced.

Further, since most of the X-rays are not diffracted by the first absorption type grating 31 and geometrically projected onto the second absorption type grating 32, high spatial coherence is required for the irradiated X-rays. Instead, a general X-ray source used in the medical field can be used as the X-ray source 11. The distance L ₂ from the first absorption type grating 31 to the second absorption type grating 32 can be set to an arbitrary value, and the distance L ₂ is smaller than the minimum Talbot interference distance in the Talbot interferometer. Since it can be set, the photographing unit 12 can be downsized (thinned). Furthermore, in this X-ray imaging system, almost all wavelength components of irradiated X-rays contribute to the projection image (G1 image) from the first absorption type grating 31 and the contrast of moire fringes is improved. Contrast image detection sensitivity can be improved.

In the X-ray imaging system 10 described above, the frequency filter is applied to the spatial frequency spectrum of the differential phase image in the frequency space to eliminate or reduce unevenness appearing in the differential phase image. You may make it remove or reduce the nonuniformity which appears in a differential phase image in real space using the convolution filter corresponding to a filter.

The X-ray imaging system 10 described above calculates the refraction angle φ by performing fringe scanning on the projection image of the first grating, and therefore the first and second gratings absorb both. Although described as a mold lattice, the present invention is not limited to this. As described above, the present invention is also useful when the refraction angle φ is calculated by performing fringe scanning on the Talbot interference image. Therefore, the first grating is not limited to the absorption type grating but may be a phase type grating.

Also, a differential phase image may be created from a radiographic image acquired by imaging (pre-imaging) in the absence of a subject. This differential phase image reflects the phase unevenness of the detection system (including phase shift due to moire, grid nonuniformity, etc.). Then, a differential phase image is created from a radiographic image acquired by imaging (main imaging) in the presence of a subject, and the differential phase image obtained by pre-imaging is subtracted from this, thereby correcting the phase unevenness of the detection system. A differential phase image can be obtained.

FIG. 12 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

A mammography apparatus 80 shown in FIG. 12 is an apparatus that captures an X-ray image (phase contrast image) of the breast B as a subject. The mammography apparatus 80 is disposed at one end of an arm member 81 that is pivotally connected to a base (not shown), and disposed at the other end of the arm member 81. An imaging table 83 and a compression plate 84 configured to be movable in the vertical direction with respect to the imaging table 83 are provided.

The X-ray source storage unit 82 stores the X-ray source 11, and the imaging table 83 stores the imaging unit 12. The X-ray source 11 and the imaging unit 12 are arranged to face each other. The compression plate 84 is moved by a moving mechanism (not shown), and the breast B is sandwiched between the imaging table 83 and compressed. The X-ray imaging described above is performed in this compressed state.

Since the X-ray source 11 and the imaging unit 12 have the same configuration as that of the X-ray imaging system 10 described above, the same reference numerals as those of the X-ray imaging system 10 are given to the respective components. Since other configurations and operations are the same as those of the X-ray imaging system 10 described above, description thereof will be omitted.

FIG. 13 shows a modification of the radiation imaging system of FIG.

A mammography apparatus 90 shown in FIG. 13 is different from the mammography apparatus 80 described above in that the first absorption grating 31 is disposed between the X-ray source 11 and the compression plate 84. The first absorption type lattice 31 is accommodated in a lattice accommodation portion 91 connected to the arm member 81. The imaging unit 92 includes an FPD 30, a second absorption type grating 32, and a scanning mechanism 33.

Thus, even when the subject (breast) B is located between the first absorption type grating 31 and the second absorption type grating 32, it is formed at the position of the second absorption type grating 32. The projection image (G1 image) of the first absorption type grating 31 is deformed by the subject B. Therefore, even in this case, the moiré fringes modulated due to the subject B can be detected by the FPD 30. That is, the mammography apparatus 90 can also obtain a phase contrast image of the subject B based on the principle described above.

In the present mammography apparatus 90, the X-ray whose dose is almost halved is irradiated to the subject B due to the shielding by the first absorption type grating 31. Therefore, the exposure amount of the subject B is determined as described above. It can be reduced to about half that of the device 80. Note that the arrangement of the subject between the first absorption type grating 31 and the second absorption type grating 32 as in the mammography apparatus 90 can also be applied to the X-ray imaging system 10 described above. Is possible.

FIG. 14 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

The X-ray imaging system 100 differs from the X-ray imaging system 10 of the first embodiment in that a multi-slit 103 is provided in the collimator unit 102 of the X-ray source 101. Since other configurations are the same as those of the X-ray imaging system 10 described above, description thereof will be omitted.

In the X-ray imaging system 10 described above, when the distance from the X-ray source 11 to the FPD 30 is a distance (1 m to 2 m) set in a general hospital imaging room, the focal point of the X-ray focal point 18b. The blur of the G1 image due to the size (generally about 0.1 mm to 1 mm) is affected, and there is a possibility that the image quality of the phase contrast image is lowered. Therefore, it is conceivable to install a pinhole immediately after the X-ray focal point 18b to effectively reduce the focal spot size. However, if the aperture area of the pinhole is reduced to reduce the effective focal spot size, the X-ray focal point is reduced. Strength will fall. In the present X-ray imaging system 100, in order to solve this problem, the multi-slit 103 is disposed immediately after the X-ray focal point 18b.

The multi-slit 103 is an absorption type grating (third absorption type grating) having a configuration similar to that of the first and second absorption type gratings 31 and 32 provided in the imaging unit 12, and is in one direction (y direction). The extended X-ray shielding portions are periodically arranged in the same direction (x direction) as the X-ray shielding portions 31b and 32b of the first and second absorption gratings 31 and 32. The multi-slit 103 is intended to form a large number of small-focus light sources (dispersed light sources) arranged at a predetermined pitch in the x direction by partially shielding the radiation emitted from the X-ray focal point 18b. .

The lattice pitch p ₃ of the multi-slit 103 needs to be set so as to satisfy the following formula (18), where L ₃ is the distance from the multi-slit 103 to the first absorption-type lattice 31.

Expression (18) indicates that the projection image (G1 image) of the X-rays emitted from the small-focus light sources dispersedly formed by the multi-slit 103 by the first absorption-type grating 31 is the position of the second absorption-type grating 32. This is a geometric condition for matching (overlapping).

In addition, since the position of the multi slit 103 is substantially the X-ray focal position, the grating pitch p2 of the _second absorption grating 32 is determined so as to satisfy the relationship of the following equation (19).

As described above, in the present X-ray imaging system, the G1 images based on the plurality of small focus light sources formed by the multi-slit 103 are superimposed, thereby improving the image quality of the phase contrast image without reducing the X-ray intensity. Can be made. The multi slit 103 described above can be applied to any of the X-ray imaging systems described above.

FIG. 15 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

The first and second absorption gratings 31 and 32 of the above-described X-ray imaging system 10 are configured such that the periodic arrangement direction of the X-ray shielding portions 31b and 32b is linear (that is, the grating surface is planar). However, instead of this, as shown in FIG. 15, first and second absorption type gratings 110 and 111 in which the grating surface is concaved into a curved surface may be used. In this case, it is preferable to use the FPD 112 having a cylindrical detection surface, and the detection surface of the FPD 112 is a cylindrical surface having a straight line passing through the X-ray focal point 18b and extending in the y direction as a central axis.

The first absorption grating 110, the X-ray permeable and curved surfaces of the substrate 110a, a plurality of X-ray shielding section 110b is periodically arranged at a predetermined pitch p _1. Each X-ray shielding part 110b extends linearly in the y direction, and the lattice plane of the first absorption grating 110 is centered on a straight line passing through the X-ray focal point 18b and extending in the extending direction of the X-ray shielding part 110b. It has a shape along a cylindrical surface as an axis. Similarly, the second absorption grating 111, the X-ray permeable and curved surfaces of the substrate 111a, a plurality of X-ray shielding section 111b is periodically arranged at a predetermined pitch p _2. Each X-ray shielding part 111b extends linearly in the y direction, and the lattice plane of the second absorption grating 111 is centered on a straight line passing through the X-ray focal point 18b and extending in the extending direction of the X-ray shielding part 111b. It has a shape along a cylindrical surface as an axis.

The distance from the X-ray focal point 18b to the first absorption grating 110 _{L 1,} the distance from the first absorption grating 110 to the second absorption grating 111 when the _{L 2,} the first absorption type the grating pitch _{p 2} of the grating pitch _{p 1} and the second absorption grating 111 of the grating 110 is determined to satisfy equation (1).

By making the grating surfaces of the first and second absorption gratings 110 and 111 cylindrical, the X-rays irradiated from the X-ray focal point 18b are all perpendicular to the grating surface when the subject H is not present. since made incident, the upper limit of the limitation of the thickness _{h 2} of the thickness _{h 1} and the X-ray shielding portion 111b of the X-ray shielding section 110b is reduced, it is not necessary to consider the equation (6) and (7).

In the X-ray imaging system using the first and second absorption type gratings 110 and 111, one of the first and second absorption type gratings 110 and 111 is arranged with the X-ray focal point 18b as the center. The above-described fringe scanning is performed by moving in the direction along the surface (cylindrical surface), the refraction angle φ is calculated, and the phase shift distribution Φ can be obtained based on the refraction angle φ.

In each of the above-described X-ray imaging systems, the case where general X-rays are used as radiation has been described. However, the radiation used in the present invention is not limited to X-rays, but other than X-rays such as α-rays and γ-rays. It is also possible to use other radiation.

As described above, the following (1) to (8) radiation imaging systems and the following (9) radiation imaging methods are disclosed in the present specification.

(1) a first grating, a second grating that masks a radiation image that is formed by the radiation that has passed through the first grating and that has been transmitted through the subject and is modulated, and the second grating. A radiation image detector that detects the masked radiation image; and an arithmetic processing unit that generates a differential phase image of the subject based on the radiation image acquired by the radiation image detector. The second grating and the second grating are arranged in a relative positional relationship in which a moire fringe having a period in a direction intersecting with a period direction of the first grating is generated in the radiation image, and the arithmetic processing unit includes: A radiation imaging system that performs a filtering process for removing or reducing unevenness of the differential phase image using a filter that affects a change in image contrast in the periodic direction of the moire fringes.
(2) The radiation imaging system according to (1), wherein the arithmetic processing unit performs the filtering process in a frequency space.
(3) In the radiographic system according to (2), the arithmetic processing unit acquires a spatial frequency spectrum of the differential phase image using Fourier transform, and uses the frequency filter to calculate the unevenness from the spatial frequency spectrum. The radiography system which removes or reduces the frequency component corresponding to.
(4) The radiation imaging system according to (1), wherein the arithmetic processing unit performs the filtering process in a real space.
(5) The radiation imaging system according to (4), wherein the arithmetic processing unit removes or reduces unevenness of the differential phase image using a convolution filter.
(6) The radiographic system according to any one of (1) to (5), wherein a periodic direction of the moire fringes is substantially orthogonal to a periodic direction of the first grating.
(7) The radiation imaging system according to any one of (1) to (6), wherein the first grating and the second grating intersect with each other at an angle of less than 90 ° in the periodic direction. The radiography system that is in place.
(8) The radiation imaging system according to (7), wherein the first grating and the second grating are relatively rotatable around an optical axis of radiation passing through both gratings.
(9) Radiation that generates a differential phase image of the subject by detecting a radiation image that has passed through the first grating and that has been modulated by passing through the subject and masked by the second grating. An imaging method, wherein the relative positional relationship between the first grating and the second grating is adjusted, and moire fringes having a period in a direction intersecting with a period direction of the first grating are generated in the radiation image. A radiation imaging method for removing or reducing unevenness of the phase differential image using a filter that acts on a change in image contrast in the period direction of the moire fringes.

According to the present invention, the periodic direction of the moire fringes is set in a direction intersecting with the periodic direction of the first grating, and unevenness appearing in the differential phase image due to the moire fringes is also the period of the first grating. It occurs in a direction that intersects the direction. Then, by using a filter that affects the image contrast change in the periodic direction of the moire fringes, that is, the direction intersecting the periodic direction of the first grating, unevenness is removed or reduced without affecting the phase information. Can do.

Although the present invention has been described in detail and with reference to specific embodiments, it will be apparent to those skilled in the art that various changes and modifications can be made without departing from the spirit and scope of the invention.
This application is based on a Japanese patent application filed on Mar. 29, 2011 (Japanese Patent Application No. 2011-072409), the contents of which are incorporated herein by reference.

10 X-ray imaging system 11 X-ray source 12 Imaging unit 13 Console 30 FPD
31 First Absorption Type Grating 32 Second Absorption Type Grating 33 Scanning Mechanism 40 Pixel 41 Image Receiver 43 Reading Circuit

Claims (9)

1. A first lattice;
   A second grating that masks a radiation image formed by the radiation that has passed through the first grating and is transmitted through the subject and modulated;
   A radiation image detector for detecting the radiation image masked by the second grating;
   An arithmetic processing unit that generates a differential phase image of the subject based on the radiographic image acquired by the radiological image detector;
   With
   The first grating and the second grating are arranged in a relative positional relationship that causes a moire fringe having a period in a direction intersecting a period direction of the first grating in the radiation image,
   The radiation processing system, wherein the arithmetic processing unit performs a filtering process for removing or reducing unevenness of the differential phase image using a filter that affects a change in image contrast in the periodic direction of the moire fringes.
2. The radiation imaging system according to claim 1,
   The said arithmetic processing part is a radiography system which performs the said filtering process in frequency space.
3. The radiographic system according to claim 2,
   The said arithmetic processing part acquires the spatial frequency spectrum of the said differential phase image using a Fourier transformation, and removes or reduces the frequency component corresponding to the said nonuniformity from the said spatial frequency spectrum using a frequency filter.
4. The radiation imaging system according to claim 1,
   The said arithmetic processing part is a radiography system which performs the said filtering process in real space.
5. The radiation imaging system according to claim 4,
   The said arithmetic processing part is a radiography system which removes or reduces the nonuniformity of the said differential phase image using a convolution filter.
6. The radiographic system according to any one of claims 1 to 5,
   The radiography system in which the periodic direction of the moire fringes is substantially orthogonal to the periodic direction of the first grating.
7. The radiographic system according to any one of claims 1 to 6,
   The radiation imaging system in which the first grating and the second grating are arranged so that their periodic directions intersect each other at an angle of less than 90 °.
8. The radiation imaging system according to claim 7,
   The radiation imaging system in which the first grating and the second grating are relatively rotatable around an optical axis of radiation passing through both gratings.
9. A radiographic method for generating a differential phase image of a subject by detecting a radiation image formed by radiation that has passed through a first grating and transmitted through a subject and is modulated and masked by a second grating. There,
   Adjusting the relative positional relationship between the first grating and the second grating, and generating a moiré fringe having a period in a direction intersecting with a period direction of the first grating in the radiation image;
   A radiation imaging method for removing or reducing unevenness of the phase differential image using a filter that affects a change in image contrast in the periodic direction of the moire fringes.

Images