JP2012035050A - Radiographic system and photographe processing method for the same - Google Patents

Abstract

PROBLEM TO BE SOLVED: To correct a noise generated due to variation of the positional relationship of grating between pre-photographing and real photographing.SOLUTION: A radiographic system 10 includes: a variation calculation section 32 for calculating the amount of variation based on a difference of the phase differential value corresponding to a transparent region 20b of FPD20 on the basis of a first phase differential image generated during real photographing performed in a state of disposing a subject by a phase differential image generation section 30, and a second phase differential image (offset data) generated during pre-photographing performed in a state without disposing the subject by the phase differential image generation section 30; a subtraction processing section 33 for subtracting the second phase differential image from the first phase differential image, and then subtracting the amount of variation of all pixels; and a phase contrast image generation section 34 for generating a phase contrast image from the phase differential image subjected to subtraction processing by the subtraction processing section 33.

Description

  The present invention relates to a radiation imaging system that images a subject with radiation such as X-rays and an image processing method thereof, and more particularly to a radiation imaging system that uses a fringe scanning method and an image processing method thereof.

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

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

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

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

  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 the distance between the X-ray source and the first diffraction grating. Are modulated by the interaction (phase change) between the subject placed between the two 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. Photographing is performed at each scanning position while translational movement (scanning) is performed at a scanning pitch obtained by equally dividing the lattice pitch, and from the amount of phase shift of intensity change with respect to the scanning position of pixel data of each pixel obtained by an X-ray image detector. Obtain a phase differential image. 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”. This fringe scanning method is also used in an imaging apparatus using laser light (see, for example, Non-Patent Document 2).

  In the fringe scanning method, the relative positional relationship between the first diffraction grating and the second diffraction grating is important for generating a phase contrast image. When distortion, fabrication error, arrangement error, etc. occur in the first diffraction grating or the second diffraction grating, an offset corresponding to the distortion or error is added to the phase differential image, and the image quality of the phase contrast image deteriorates. End up. In Patent Document 1, the above-described offset amount is acquired by performing pre-imaging without arranging a subject, and the offset amount is subtracted from the phase differential image obtained by main imaging performed by arranging the subject. Describes obtaining a phase differential image in which only the characteristics of the object are reflected.

Japanese Patent No. 4445397

C. David, et al., Applied Physics Letters, Vol. 81, No. 17, October 2002, p. 3287 Hector Canabal, et al., Applied Optics, Vol.37, No.26, September 1998, 6227

  However, in the offset correction method described in Patent Document 1, since pre-imaging and main imaging are premised on the same imaging except for the presence or absence of the subject, between pre-imaging and main imaging. There is a problem that noise generated when the scanning start position of the second diffraction grating relative to the first diffraction grating fluctuates cannot be corrected. This noise caused by fluctuations in the scanning start position displaces the entire phase differential image, and degrades the image quality of the resulting phase contrast image. In addition, the image quality deterioration is not limited to the change in the scanning start position between the pre-photographing and the main photographing, and there is some variation in the positional relationship between the first and second diffraction gratings between the pre-photographing and the main photographing. It also occurs when there is.

  The present invention has been made in view of the above problems, and provides a radiation imaging system and an image processing method thereof capable of correcting noise generated due to a change in the positional relationship of the lattice between pre-imaging and actual imaging. The purpose is to do.

  In order to achieve the above object, a radiation imaging system according to the present invention includes a first grating that generates a first periodic pattern image by passing radiation irradiated from a radiation source, and the first periodic pattern image. Intensity modulation means for applying intensity modulation to the second periodic pattern image, and the second periodic pattern image incident on the detection surface of the second periodic pattern image without passing through the subject. A radiation image detector having a blank area for detecting a component and detecting the second periodic pattern image to generate image data; and a phase differential image generating a phase differential image based on the image data A first phase differential image generated by the phase differential image generation unit in a state in which the subject is disposed; and a second phase generated by the phase differential image generation unit in a state in which the subject is not disposed. Based on differential image , A fluctuation amount calculating means for calculating a fluctuation amount by taking a difference between the phase differential values of the first and second phase differential images in the blank region, and the second phase from the first phase differential image. Subtracting processing means for subtracting the phase differential image and further subtracting the fluctuation amount for all pixels.

  The radiation imaging system of the present invention includes: a storage unit that stores the second differential phase image; and an intensity modulation unit, the radiation image detector, and the pre-imaging instruction in a state where no subject is disposed. Control means for operating the phase differential image generation means and storing the phase differential image generated by the phase differential image generation means in the storage means as the second phase differential image.

  The radiation imaging system of the present invention further includes a housing that holds the first grating, the intensity modulation unit, and the radiation image detector, and has an incident surface formed so as to be able to identify the blank region.

  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 phase differential image generated as a result of the subtraction processing performed by the subtraction processing means.

  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 first and second gratings are absorption gratings, and the first grating linearly projects the radiation from the radiation source onto the second grating as the first periodic pattern image.

  The first and second gratings are formed by alternately arranging lattice-lined high absorption parts having a high radiation absorption rate and lattice line-like low absorption parts having a low radiation absorption rate. The high absorption portion is made of Au or Pt. The high absorption part has a thickness of 10 μm to 200 μm. The low absorption portion is made of silicon or polymer.

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

  In this case, the first grating is formed by alternately arranging a lattice-line high absorption portion having a high radiation absorption rate and a lattice line-shaped low absorption portion having a low radiation absorption rate, and the high absorption portion. A phase difference of π or 0.5π is given to the radiation between the part and the low absorption part. The high absorption portion is made of Au or Pt. The low absorption portion is made of silicon or polymer.

  The radiation imaging system of the present invention preferably includes a radiation source grid on the emission side of the radiation source.

  An image processing method according to the present invention includes a first grating that generates a first periodic pattern image by passing radiation emitted from a radiation source, and intensity modulation is applied to the first periodic pattern image. Intensity modulation means for generating two periodic pattern images, and a blank for detecting a component of the second periodic pattern image incident on the detection surface of the second periodic pattern image without passing through the subject Radiation comprising a radiation image detector that has a region and detects the second periodic pattern image to generate image data, and phase differential image generation means for generating a phase differential image based on the image data An image processing method for an imaging system, wherein the first differential phase image generated by the phase differential image generation unit in a state in which the subject is arranged and the phase differential image generation unit in a state in which the subject is not arranged Second to be Calculating a variation amount by taking a difference between the phase differential values of the first and second phase differential images in the blank region based on the phase differential image; and Subtracting the phase differential image of 2 and subtracting the fluctuation amount for all pixels.

  In the present invention, a first phase differential image generated in a state where a subject is placed (main imaging) and a second phase differential image (offset data) generated in a state where no subject is placed (pre-imaging). And calculating the amount of variation by taking the difference of the phase differential value corresponding to the undetected region of the radiation detector, subtracting the second phase differential image from the first phase differential image, Since the above-described variation amount is subtracted for the pixel, it is possible to correct noise generated due to variation in the positional relationship of the grid between the pre-photographing and the main photographing.

1 is a schematic diagram illustrating a configuration of an X-ray imaging system according to a first embodiment of the present invention. It is a schematic perspective view which shows the housing | casing of an imaging | photography part. It is a schematic diagram which shows the structure of a flat panel detector. It is a schematic side view which shows the structure of the 1st and 2nd absorption type grating | lattice. It is explanatory drawing for demonstrating a fringe scanning method. (A) is a graph illustrating the intensity modulation signal of the subject detection region, and (b) is a graph illustrating the intensity modulation signal of the background region. It is a flowchart which shows the effect | action at the time of the pre imaging | photography of an X-ray imaging system. It is a flowchart which shows the effect | action at the time of this imaging | photography of an X-ray imaging system. It is a figure which shows the multi slit used in 2nd Embodiment of this invention. It is a schematic diagram which shows the structure of the X-ray image detector used in 4th Embodiment of this invention.

(First embodiment)
In FIG. 1, an X-ray imaging system 10 according to the first embodiment of the present invention is disposed so as to face an X-ray source 11 that irradiates a subject H with X-rays, and is opposed to 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 detection surface of the FPD 20 mainly includes an object detection region 20a in which X-rays that have passed through the object H enter through the first and second absorption gratings 21 and 22 and an X-ray that has passed around the object H. Is divided into a blank region 20 b that does not pass through the subject H and is incident only through the first and second absorption gratings 21 and 22.

The first absorption type grating 21 includes a plurality of X-ray shielding portions (X-ray high absorption portions) 21a extending in one direction (hereinafter referred to as y direction) in a plane orthogonal to the z direction. Are arranged at a predetermined pitch p ₁ in a direction orthogonal to (hereinafter referred to as the x direction). Similarly, the second absorption grating 22 has a plurality of X-ray shielding portion which is stretched in the y-direction (X-ray high-absorbing portion) 22a is arranged at a predetermined pitch p ₂ in the x-direction. As a material of the X-ray shielding portions 21a and 22a, a metal excellent in X-ray absorption is preferable, for example, gold (Au) or platinum (Pt) is 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.

  The imaging unit 12 configured as described above is held inside a rectangular casing 24 as shown in FIG. On the X-ray incident surface 24a of the casing 24, a rectangular frame indicating the center lines 25a and 25b indicating the center positions in the x direction and the y direction and the boundary between the subject detection area 20a and the blank area 20b of the FPD 20 is provided. A line 25c is formed by printing. The outside of the frame line 25c corresponds to the blank area 20b. It should be noted that other methods may be used as long as the blank region 20b can be recognized without being limited to the printing of the frame line 25c.

  The image processing unit 14 includes a phase differential image generation unit 30, an offset data storage unit 31, a fluctuation amount calculation unit 32, a subtraction processing unit 33, and a phase contrast image generation unit 34. The phase differential image generation unit 30 generates a phase differential image based on a plurality of image data captured by the imaging unit 12 at each scanning step of fringe scanning by the scanning mechanism 23 and stored in the memory 13.

  The offset data storage unit 31 uses, as offset data, a phase differential image generated by the phase differential image generation unit 30 at the time of imaging (pre-imaging) in a state in which the subject H is not disposed between the X-ray source 11 and the imaging unit 12. Remember. The offset data storage unit 31 is configured by a nonvolatile storage device such as a flash memory.

  The fluctuation amount calculation unit 32 performs imaging in a state in which the subject H is arranged between the X-ray source 11 and the imaging unit 12 so that X-rays transmitted through the subject H are projected into the frame 25c ( The phase differential image generated by the phase differential image generation unit 30 during the main photographing) and the offset data (phase differential image) stored in the offset data storage unit 31 are the predetermined pixels included in the blank region 20b of the FPD 20 The difference value of the phase differential value is calculated for. Since the unexposed region 20b is a region where X-rays are incident without passing through the subject H during both pre-imaging and main imaging, scanning of the second absorption grating 22 is started in pre-imaging and main imaging. If there is no change in the position (initial value), the phase differential value is the same and the difference value is 0. The fluctuation amount calculation unit 32 detects the fluctuation amount of the scanning start position of the second absorption type grating 22 by calculating the difference value of the phase differential value of the background region 20b between the pre-photographing and the main photographing. is there.

  The subtraction processing unit 33 subtracts the offset data stored in the offset data storage unit 31 from the phase differential image generated by the phase differential image generation unit 30 for the main photographing, and is further calculated by the fluctuation amount calculation unit 32. The difference value is subtracted for the entire pixel. The phase contrast image generation unit 34 integrates the phase differential image subjected to the subtraction processing by the subtraction processing unit 33 along the scanning direction (x direction) to generate a phase contrast image. The phase contrast image generated by the phase contrast image generation unit 34 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, X-ray imaging conditions such as X-ray tube voltage and 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 accumulate them two-dimensionally on the active matrix substrate along the x direction and the y direction. The scanning circuit 42 controls the readout timing of the charges, and the readout circuit 43 that reads the charges from the pixels 40, converts the charges into image data, and outputs the image data. 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. This is a conversion type X-ray detection element. Each pixel 40 is provided with a TFT switch (not shown). 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 use an indirect conversion type X-ray detection element that converts the charge into a charge and stores it. 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 dark current correction, gain correction, linearity correction, and the like on the image data, and inputs the corrected image data to the memory 13.

  The scanning circuit 42 and the readout circuit 43 are controlled by the system control unit 18 via the imaging control unit 16. As described above, the image receiving unit 41 is divided into the subject detection region 20a and the blank region 20b. The pixels 40 included in the subject detection region 20a and the blank region 20b have the same configuration, and the system control unit 18 uses the addresses of the scanning lines 44 and the signal lines 45 to determine the pixels 40 included in the subject detection region 20a. And the pixel 40 included in the blank area 20b.

In FIG. 4, X-ray shielding portion 21a of the first absorption grating 21, at a predetermined pitch p ₁ in the x-direction, are arranged at a predetermined interval d ₁ from each other, in a portion of the distance d ₁ The X-ray low absorption part 21b is provided. Similarly, X-rays shielding portions 22a of the second absorption-type grating 22, at a predetermined pitch p ₂ in the x-direction, are arranged at a predetermined distance from each other d _2, in a portion of the distance d _2, An X-ray low absorption part 22b is provided. 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. The X-ray low absorption portions 21b and 22b are preferably made of silicon (Si) or a polymer, and may be voids.

The first and second absorption type gratings 21 and 22 are configured to project X-rays that have passed through the X-ray low absorption portions 21b and 22b linearly (geometrically). Specifically, by setting the distances d ₁ and d ₂ to a value sufficiently larger than the peak wavelength of the X-rays emitted from the X-ray source 11, most of the X-rays included in the irradiated X-rays are reduced to X-rays Without being diffracted by the absorbers 21b and 22b, it is configured to pass 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 linearly projected without being diffracted by the X-ray low absorption portions 21b and 22b. The grating pitches p ₁ and p ₂ are about ₂ to 20 μm.

The X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam having an X-ray focal point as a light emission point, and therefore the first period formed by passing through the first absorption type grating 21. The pattern image (hereinafter referred to as G1 image) is enlarged in proportion to the distance from the X-ray focal point 11a. The lattice pitch p ₂ and the interval d ₂ of the second absorption type grating 22 are substantially the same as the periodic pattern of the bright part of the G1 image at the position of the second absorption type grating 22 in the X-ray low absorption part 22 b pattern. Has been determined to be. 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).

It is not always necessary to satisfy the expression (2), and the intervals d ₁ and d ₂ may be set individually.

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 shield (absorb) X-rays completely in order to generate a periodic pattern image with high contrast. However, the materials having excellent X-ray absorption properties (gold, platinum) 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, and the thicknesses of the X-ray shielding portions 21a and 22a are preferably in the range of 10 μm to 200 μm.

  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. Further, when an error occurs in the lattice arrangement direction of the first and second absorption type gratings 21 and 22 and the arrangement directions are not the same, a so-called rotational moire occurs in the G2 image. However, even when moire fringes occur in the G2 image, there is no particular problem as long as the period of the moire fringes in the x direction or y direction is larger than the arrangement pitch of the pixels 40. Ideally, moire fringes are preferably not generated, but the moire fringes can be used for confirming the scanning amount of the fringe scanning (translation distance of the second absorption type grating 22), as will be described later.

  When the subject H is disposed between the X-ray source 11 and the first absorption type grating 21, the 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, the principle of the 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 travels straight when the subject H does not exist. The X-ray that travels along the path 50 passes through the first and second absorption gratings 21 and 22 to the FPD 20. Incident. Reference numeral 51 indicates an X-ray path refracted and deflected by the subject H when the subject H exists. X-rays traveling along the path 51 pass through the first absorption type grating 21 and are then shielded by the X-ray shielding part 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. . 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 during pre-photographing and main photographing, and a phase differential image is generated for each.

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. An intensity modulation signal of each pixel is obtained from a plurality of image data obtained by photographing, and is processed by the phase differential image generation unit 30 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.

  When the relational expression expressed by the following expression (10) is applied, the refraction angle φ (x) is expressed as 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 FIGS. 6A and 6B, the intensity modulation signal periodically changes with a period of the grating pitch p _{2 with} respect to the position k of the second absorption type grating 22. . The solid line in the figure exemplifies the intensity modulation signal obtained at the time of actual photographing, and has a phase shift amount ψ ₁ . The broken line exemplifies an intensity modulation signal obtained at the time of pre-photographing, and has a phase shift amount ψ ₀ . The phase shift amount ψ ₀ at the time of pre-photographing is caused by lattice distortion, creation error, arrangement error, and the like of the first and second absorption gratings 21 and 22.

FIG. 6A illustrates an intensity modulation signal of the pixel 40 in the subject detection area 20a, and shows a case where a phase shift occurs in the intensity modulation signal due to the influence of the subject H during the main imaging. . On the other hand, FIG. 6B shows an example of the intensity modulation signal of the pixel 40 in the blank region 20b, and the phase difference (Δψ = ψ ₁ −ψ ₀ ) of the intensity modulation signal is pre-photographed. This is due to the amount of change in the scanning start position (the position of k = 0 in FIG. 5) of the second absorption type grating 22 between the time and the actual photographing. When an actuator such as a piezoelectric element is used as the scanning mechanism 23, the scanning pitch (p ₂ / M) can be controlled with relatively high accuracy. After scanning from 0 to k = M−1, when returning to the initial position of k = 0, the original position is not accurately returned, and an error that cannot be ignored (about several μm) occurs. Sometimes. This error corresponds to the fluctuation amount.

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 on the y coordinate, a two-dimensional phase shift amount distribution ψ ₀ (x , Y) and ψ ₁ (x, y). ψ ₀ (x, y) corresponds to a phase differential image (offset data) input from the phase differential image generation unit 30 to the offset data storage unit 31 during pre-imaging. Further, ψ ₁ (x, y) corresponds to a phase differential image input from the phase differential image generation unit 30 to the fluctuation amount calculation unit 32 and the subtraction processing unit 33 at the time of actual imaging. The physical quantity constituting the phase differential image is not limited to the phase shift amount ψ, and any physical quantity may be used as long as it has a proportional relationship with the differential value of the phase shift distribution Φ (x) such as the refraction angle φ. Also good.

The fluctuation amount calculation unit 32 calculates ψ ₁ (x, y) −ψ ₀ (x, y) for the predetermined pixel 40 in the background region 20b, and supplies this to the subtraction processing unit 33 as the variation amount Δψ. To do. Subtraction processing unit 33, for all the pixels 40 in the analyte detection region 20a, subtracts ψ ₁ (x, y) from ψ ₀ (x, y), further, it subtracts the variation amount [Delta] [phi]. As a result, in addition to the distortion of the first and second absorption gratings 21 and 22, the creation error, the arrangement error, and the like, the influence of the variation in the scanning start position of the second absorption grating 22 is removed, and the subject is examined. A phase differential image reflecting only the characteristics of H is obtained. Note that the fluctuation amount calculation unit 32 calculates ψ ₁ (x, y) −ψ ₀ (x, y) for the plurality of pixels 40 in the background region 20b, and averages these values to obtain the variation amount Δψ. It is also preferable that

Next, the operation of the X-ray imaging system 10 configured as described above will be described. As shown in the flowchart of FIG. 7, when an instruction to start pre-imaging is given from the console 17 (YES in step S <b> 10), each unit of the X-ray imaging system 10 operates in cooperation to Along with the scanning, X-ray exposure by the X-ray source 11 and detection operation by the FPD 20 are performed (step S11), and based on the plurality of image data stored in the memory 13, the phase differential image generation unit 30 causes the phase differential image ψ. ₀ (x, y) is generated (step S12) and stored in the offset data storage unit 31 as offset data (step S13). The pre-photographing operation is completed as described above, and the second absorption grating 22 is returned to the scanning start position (initial position) (step S14). At the end of the pre-photographing, the operator is notified that the pre-photographing is finished by displaying a message on the monitor (step S15).

  This pre-imaging does not have to be performed every time the main imaging is performed, and may be performed as appropriate when the X-ray imaging system 10 is started up. The offset data stored in the offset data storage unit 31 is overwritten with newly obtained offset data when pre-photographing is performed again.

Next, the main imaging is performed in a state where the X-rays transmitted through the subject H are arranged so as to be incident only on the frame line 25 c indicated on the X-ray incident surface 24 a of the housing 24. As shown in the flowchart of FIG. 8, when an instruction to start pre-imaging is given from the console 17 (YES in step S20), the X-ray source 11 is scanned along with the scanning of the second absorption grating 22 as in pre-imaging. The X-ray exposure and detection operation by the FPD 20 are performed (step S21), and the phase differential image ψ ₁ (x, y) is generated by the phase differential image generation unit 30 (step S22).

Next, the fluctuation amount calculation unit 32 subtracts the offset data ψ ₀ (x, y) of the offset data storage unit 31 from the phase differential image ψ ₁ (x, y) for the predetermined pixel 40 in the background region 20b. Then, the fluctuation amount Δψ is calculated (step S23). The subtraction processing unit 33 subtracts the offset data ψ ₀ (x, y) from the phase differential image ψ ₁ (x, y) for all the pixels 40 in the subject detection region 20a, and further detects the subject. The variation Δψ is subtracted for all the pixels 40 in the region 20a (step S24).

  The phase differential image subjected to the subtraction processing by the subtraction processing unit 33 is input to the phase contrast image generation unit 34, and a phase contrast image is generated by integration processing in a direction corresponding to the x direction (step S25). The phase contrast image is recorded in the image recording unit 15 and then output to the console 17 and displayed on the monitor (step S26). This is the end of the actual shooting.

In step S24, the offset data ψ ₀ (x, y) is obtained from the phase differential image ψ ₁ (x, y) not only for the pixels 40 in the subject detection region 20a but also for all the pixels 40 of the image receiving unit 41. Subtraction may be performed, and the variation Δψ may be subtracted for all the pixels 40 of the image receiving unit 41.

  As described above, the phase contrast image generated by the first embodiment of the present invention includes the second and second absorption gratings 21 and 22 in addition to the distortion, creation error, arrangement error, and the like. A high-quality X-ray image in which only the characteristics of the subject H are reflected is removed without being affected by the change in the scanning start position of the absorption grating 22. In the first embodiment of the present invention, the first absorption between the pre-photographing and the main photographing is not limited to the change in the scanning start position of the second absorption grating 22 between the pre-photographing and the main photographing. Needless to say, even when the position of the absorption grating 21 fluctuates in the x direction for some reason, the influence of the fluctuation is eliminated.

(Second Embodiment)
In the first 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 focal point 11a (generally about 0.1 mm to 1 mm) affects the phase. There is a risk of degrading the image quality of the contrast image. Therefore, as a second embodiment of the present invention, as shown in FIG. 9, a multi slit (ray source lattice) 60 is arranged on the emission side of the X-ray source 11. The X-ray imaging system of the second embodiment has the same configuration as that of the first embodiment except that the multi-slit 60 is provided.

  The multi-slit 60 is an absorption-type grating having the same configuration as the first and second absorption-type gratings 21 and 22, and a plurality of X-ray shielding portions 61 extending in the y direction are periodically arranged in the x direction. It is a thing. The multi-slit 60 partially shields the 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. , G1 image blur is suppressed. Similarly, an X-ray low absorption part (not shown) is provided between the X-ray shielding parts 61 adjacent in the x direction.

  In the present embodiment, even when the positions of the first absorption grating 21 and the multi-slit 60 change in the x direction for some reason between the pre-photographing and the main photographing, the influence due to the variation is removed from the phase contrast image. Is done.

(Third embodiment)
In the first and second embodiments, the first absorption type grating 21 is configured to linearly project the X-rays that have passed through the X-ray low absorption part 21b, but the present invention has this configuration. However, the present invention is not limited thereto, and a configuration described in Japanese Patent No. 4445397 in which a so-called Talbot interference effect is generated by diffracting X-rays by the first absorption type grating 21 may be employed. As a third embodiment of the present invention, the first absorption-type grating 21 is a diffraction grating, the distance L ₂ between the first and second absorption-type gratings 21 and 22 is set as the Talbot interference distance, and Talbot interference is performed. Configure the total. In the present embodiment, a G1 image (self-image) generated by the first grating 21 due to the Talbot interference effect is formed at the position of the second absorption grating 22.

  In the present embodiment, the first absorption type grating 21 may be a phase type grating (phase type diffraction grating). In this case, if the thickness and material are set such that a phase difference of “π” or “0.5π” is generated in the X-ray between the X-ray high absorption portion 21a and the X-ray low absorption portion 21a. good.

  In the first to third embodiments, the subject H is arranged 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. You may arrange | position between 2nd absorption type | mold grating | lattices 22. FIG. In this case as well, a phase contrast image can be similarly generated.

(Fourth embodiment)
In the first to third embodiments, the second absorption type grating 22 is provided independently of the FPD 20, but an X-ray image detector having a configuration disclosed in Japanese Patent Laid-Open No. 2009-133823. The second absorption type grating 22 can be eliminated by using. As a fourth embodiment of the present invention, an X-ray image detector having the following configuration is used without the second absorption grating 22.

  The X-ray image detector of the present embodiment is a direct conversion type X-ray image detector including a conversion layer that converts X-rays into charges, and a charge collection electrode that collects charges converted in the conversion layer. The charge collection electrode of each pixel is configured by arranging a plurality of linear electrode groups formed by electrically connecting linear electrodes arranged at a constant period so that their phases are different from each other. In this embodiment, the charge collection electrode constitutes the intensity modulation means described in the claims.

  10, in the FPD 70 of this embodiment, pixels 71 are two-dimensionally arranged at a constant pitch along the x direction and the y direction, and each pixel 71 is provided with a conversion layer that converts X-rays into charges. A charge collecting electrode 72 for collecting the converted charge is formed. The charge collection electrode 72 includes first to sixth linear electrode groups 72a to 72f, and the phase of the arrangement period of the linear electrodes of each linear electrode group is shifted by π / 3. Specifically, if the phase of the first linear electrode group 72a is 0, the phase of the second linear electrode group 72b is π / 3, the phase of the third linear electrode group 72c is 2π / 3, The phase of the fourth linear electrode group 72d is π, the phase of the fifth linear electrode group 72e is 4π / 3, and the phase of the sixth linear electrode group 72f is 5π / 3.

  Further, each pixel 71 is provided with a switch group 73 for reading out the charges collected by the charge collection electrode 72. The switch group 73 includes TFT switches provided in each of the first to sixth linear electrode groups 72a to 72f. The charges collected by the first to sixth linear electrode groups 72a to 72f are individually read out by controlling the switch group 73, so that six types of G2 images having different phases can be detected by one imaging. Is done. A phase contrast image is generated based on a plurality of image data corresponding to the six types of G2 images. Since other configurations are the same as those of the first embodiment, description thereof will be omitted.

  In the present embodiment, since the second absorption type grating 22 is not required from the imaging unit 12, 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. In place of the charge collecting electrode 72 having the above configuration, a charge collecting electrode having another configuration described in Japanese Patent Laid-Open No. 2009-133823 can be used.

  Furthermore, as another embodiment that makes it possible to eliminate the second absorption type grating 22, a G1 image is directly detected by an X-ray image detector, and periodically sampled while changing the phase by signal processing. It is also possible to generate image data corresponding to a plurality of G2 images having different phases.

(Fifth embodiment)
In the first to fourth embodiments, the phase differential image is obtained by the fringe scanning method, but the present invention is not limited to this, and the phase differential image is obtained by the Fourier transform method described in International Publication WO2010 / 050833. You may ask for. 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.

  Each embodiment described above is not limited to a radiographic system for medical diagnosis, but can be applied to other radiographic systems for industrial use. In addition to X-rays, gamma rays or the like can be used as radiation.

DESCRIPTION OF SYMBOLS 10 X-ray imaging system 11 X-ray source 11a X-ray focus 12 Imaging part 14 Image processing part 16 Imaging control part 18 System control part 20 Flat panel detector (FPD)
21 First absorption type grating (first grating)
21a X-ray shielding part (X-ray high absorption part)
21b X-ray low absorption part 22 2nd absorption type | mold grating | lattice (2nd grating | lattice)
22a X-ray shielding part (X-ray high absorption part)
22b X-ray low absorption unit 23 Scanning mechanism 24 Case 30 Phase differential image generation unit 31 Offset data storage unit 32 Fluctuation amount calculation unit 33 Subtraction processing unit 34 Phase contrast image generation unit 40 Pixel 41 Image reception unit 60 Multi-slit (line source grating) )
70 FPD
71 pixels 72 charge collection electrodes 72a to 72f first to sixth linear electrode groups 73 switch groups

Claims (17)

A first grating for generating a first periodic pattern image by passing 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 detection area for detecting a component of the second periodic pattern image incident on the detection surface of the second periodic pattern image without passing through the subject, and the second periodic pattern image is A radiation image detector for detecting and generating image data;
Phase differential image generating means for generating a phase differential image based on the image data;
A first differential phase image generated by the phase differential image generation unit in a state where the subject is disposed, and a second differential phase image generated by the phase differential image generation unit in a state where the subject is not disposed. A fluctuation amount calculating means for calculating a fluctuation amount by taking a difference between the phase differential values of the first and second phase differential images in the background missing region,
Subtracting means for subtracting the second phase differential image from the first phase differential image, and further subtracting the fluctuation amount for all pixels;
A radiation imaging system comprising: Storage means for storing the second phase differential image;
A phase differential image generated by the phase differential image generation unit by operating the intensity modulation unit, the radiation image detector, and the phase differential image generation unit in response to a pre-imaging instruction in a state where no subject is arranged. Control means for storing in the storage means as the second differential phase image,
The radiation imaging system according to claim 1, further comprising:   3. A housing having an incident surface that holds the first grating, the intensity modulation unit, and the radiation image detector and is formed so as to identify the blank region. The radiation imaging system described in 1.   4. The method according to claim 1, further comprising phase contrast image generation means for generating a phase contrast image by performing an integration process on a phase differential image generated as a result of the subtraction processing performed by the subtraction processing means. The radiation imaging system according to claim 1. 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 5, wherein   The first and second gratings are absorption gratings, and the first grating projects the radiation from the radiation source linearly onto the second grating as the first periodic pattern image. The radiation imaging system according to claim 6.   The first and second gratings are characterized in that lattice-lined high absorption parts having a high radiation absorption rate and grid line-like low absorption parts having a low radiation absorption rate are alternately arranged. The radiation imaging system according to claim 7.   The radiation imaging system according to claim 8, wherein the high absorption portion is made of Au or Pt.   The radiation imaging system according to claim 9, wherein the high absorption portion has a thickness of 10 μm to 200 μm.   The radiation imaging system according to claim 8, wherein the low absorption portion is made of silicon or a polymer.   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.   The first grating is formed by alternately arranging a grid-line high absorption part having a high radiation absorption rate and a grid line-like low absorption part having a low radiation absorption rate, and the high absorption part and the The radiation imaging system according to claim 12, wherein a phase difference of π or 0.5π is given to the radiation with the low absorption part.   The radiation imaging system according to claim 13, wherein the high absorption portion is made of Au or Pt.   The radiation imaging system according to claim 13, wherein the low absorption portion is made of silicon or polymer.   The radiation imaging system according to claim 1, further comprising a source grid on an emission side of the radiation source. A first grating for generating a first periodic pattern image by passing 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 detection area for detecting a component of the second periodic pattern image incident on the detection surface of the second periodic pattern image without passing through the subject, and the second periodic pattern image is A radiation image detector for detecting and generating image data;
Phase differential image generating means for generating a phase differential image based on the image data;
An image processing method for a radiation imaging system comprising:
A first differential phase image generated by the phase differential image generation unit in a state where the subject is disposed, and a second differential phase image generated by the phase differential image generation unit in a state where the subject is not disposed. A step of calculating a variation amount by taking a difference between the phase differential values of the first and second phase differential images in the blank region,
Subtracting the second phase differential image from the first phase differential image, and further subtracting the variation amount for all pixels;
An image processing method comprising:

Images