JP2014012029A - Radiation imaging system and image processing method - Google Patents

Abstract

A phase differential image with high accuracy can be obtained even when a positional shift occurs in the relative position of a grating during fringe scanning.
First and second absorption gratings 21 and 22 are arranged between an X-ray source 11 and an X-ray image detector (FPD) 20, and the second absorption grating 22 is a first absorption. In the X-ray imaging system 10 that performs imaging at a plurality of relative positions while relatively moving in a direction (x direction) perpendicular to the grid line with respect to the mold grid 21, the first and second absorption type gratings 21 and 22 are used. A position sensor 24 for detecting the position in the x direction is provided. The phase differential image generation 14b of the image processing unit 14 uses image data obtained at each relative position and a detection value by the position sensor 24, and an object of an intensity modulation signal (change in pixel data with respect to the relative position) for each pixel. A phase differential image is generated by calculating the phase shift amount due to.
[Selection] Figure 1

Description

  The present invention relates to a radiation imaging system and an image processing method for imaging a subject with radiation such as X-rays, and more particularly to a radiation imaging system and an image processing method using a fringe scanning 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 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 the background of such problems, in recent years, instead of X-ray intensity change 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 produces a high-contrast image even for weakly absorbing objects with low X-ray absorption ability. Can be obtained. As a kind of such X-ray phase imaging, an X-ray imaging system using an X-ray Talbot interferometer composed of two transmission diffraction gratings and an X-ray image detector has been devised (for example, Patent Document 1). Non-Patent Document 1).

  The X-ray Talbot interferometer has a first diffraction grating disposed behind the subject, and is second downstream from the first diffraction grating by a Talbot interference distance determined by the grating pitch of the first diffraction grating and the X-ray wavelength. Are arranged, and an X-ray image detector is arranged behind the 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 of the first diffraction grating due to the Talbot interference effect. This self-image is modulated by the interaction (phase change) between the subject arranged between the X-ray source and the first diffraction grating and the X-ray.

  In the X-ray Talbot interferometer, the moiré fringes generated by the superposition (intensity modulation) of the first image of the first diffraction grating and the second diffraction grating are detected by the fringe scanning method. A phase contrast image of the subject is acquired. In the fringe scanning method, imaging is performed at a plurality of relative positions while changing the relative positions of the first and second diffraction gratings in a direction perpendicular to the grating lines. A phase differential image is acquired from the amount of phase shift of the intensity change of the pixel data of each pixel obtained by the X-ray image detector at each relative position (phase shift amount with and without the subject). Since this phase differential image is an image representing the angular distribution of X-rays refracted by the subject, a phase contrast image of the subject can be obtained by integrating this. Since the pixel data of each pixel is a signal whose intensity is periodically modulated with respect to the change in the relative position, the set of pixel data for each relative position is hereinafter referred to as an “intensity modulation signal”. This fringe scanning method is also used in an imaging apparatus that uses laser light instead of X-rays (see, for example, Non-Patent Document 2).

  In the radiographic system using the fringe scanning method as described above, if a positional shift occurs in the relative positions of the first and second diffraction gratings during fringe scanning, an error occurs in the calculation of the phase differential image, and the phase contrast image There is a problem that the image quality deteriorates. To solve this problem, a sensor for detecting the relative positions of the first and second diffraction gratings is provided, and when the detected value is outside a predetermined range, the detection result is displayed to warn the operator. It has been proposed (see Patent Document 2).

Japanese Patent No. 4445397 JP 2008-200320 A

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 radiation imaging system described in Patent Document 2, when a positional deviation occurs in the relative position, the operator is only warned and the phase differential image calculation error itself is not improved. In order to obtain an image with good image quality, it is necessary to take a picture again after eliminating the cause of the positional deviation. When re-imaging is performed in this way, there is a problem that the subject is unnecessarily exposed. Therefore, it is desired to prevent the deterioration of image quality without performing re-shooting even when a positional shift occurs in the relative position.

  The present invention has been made in view of the above problems, and even when a positional shift occurs in the relative position of the grating, it is possible to obtain a highly accurate differential phase image and prevent deterioration in image quality. An object is to provide a radiation imaging system and an image processing method.

  To achieve the above object, a radiation imaging system of the present invention includes a radiation image detector that detects radiation emitted from a radiation source, and a first disposed between the radiation source and the radiation image detector. And a second grating, scanning means for changing the relative positions of the first and second gratings in a direction perpendicular to the grating lines, and first position detection for detecting each relative position changed by the scanning means The phase shift amount of the intensity modulation signal is calculated for each pixel using the means, the image data obtained by the radiation image detector at each relative position, and the detection value by the first position detection means. Phase differential image generation means for generating a differential image.

  The phase differential image generating means further includes second position detecting means for detecting the position of the radiation source, the phase differential image generating means, image data obtained by the radiation image detector at each relative position, and the first and first It is preferable to calculate the phase shift amount of the intensity modulation signal for each pixel using the detection value obtained by the position detection unit 2.

  Further, it is preferable that the phase differential image generating means calculates a phase shift amount of the intensity modulation signal using a calculation formula based on least squares.

  The phase differential image generated by the phase differential image generating means is preferably provided with a phase contrast image generating means for generating a phase contrast image by integrating the phase differential image along the change direction of the relative position.

  The first and second gratings are preferably absorption gratings, and the first grating projects the radiation from the radiation source onto the second grating as a fringe image.

  The first grating is preferably a phase-type grating, and it is preferable that the radiation from the radiation source is projected onto the second grating as a fringe image by a Talbot interference effect.

  A radiation source grid may be provided on the emission side of the radiation source. In this case, it further comprises second position detecting means for detecting the position of the source grating, and the phase differential image generating means includes image data obtained by the radiation image detector at each relative position, It is preferable to calculate the phase shift amount of the intensity modulation signal for each pixel using the detection values obtained by the first and second position detection means.

  Furthermore, the image processing method of the present invention includes a radiation image detector that detects radiation emitted from a radiation source, and first and second gratings disposed between the radiation source and the radiation image detector. A radiation imaging system comprising: scanning means for changing the relative positions of the first and second gratings in a direction perpendicular to the grating lines; and position detection means for detecting the relative positions changed by the scanning means. In the image processing method used for calculating the phase shift amount of the intensity modulation signal for each pixel, using the image data obtained by the radiation image detector at each relative position and the detection value by the position detection means. To generate a differential phase image.

  The present invention includes position detection means for detecting the relative positions of the first and second gratings changed by the scanning means, image data obtained by the radiation image detector at each relative position, and detection by the position detection means. Since the phase differential image is generated by calculating the phase shift amount of the intensity modulation signal for each pixel using the value, it is possible to obtain a highly accurate phase differential image even if a positional shift occurs in the relative position And deterioration of image quality can be prevented.

1 is a block diagram showing a configuration of an X-ray imaging system according to a first embodiment of the present invention. 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. It is a graph which illustrates the change of the pixel data obtained at a scanning step, (a) illustrates the case where there is no position shift, and (b) illustrates the case where a position shift occurs. It is a block diagram which shows the structure of the X-ray imaging system which concerns on 2nd Embodiment of this invention. It is a schematic diagram which shows the moving amount | distance of the G1 image accompanying the position shift | offset | difference of a X-ray focus. It is a schematic perspective view which shows the multi slit of 3rd Embodiment of this invention.

(First embodiment)
In FIG. 1, an X-ray imaging system 10 according to the first embodiment of the present invention detects an X-ray that is disposed so as to face an X-ray source 11 and the X-ray source 11 and has transmitted through a subject H from the X-ray source 11. Then, the image capturing unit 12 that generates image data, the memory 13 that stores the image data read from the image capturing unit 12, and the plurality of image data stored in the memory 13 are processed to generate a phase contrast image. The image processing unit 14 includes an image recording unit 15 that records the phase contrast image generated by the image processing unit 14, an imaging control unit 16 that controls the X-ray source 11 and the imaging unit 12, and an operation unit and a monitor. A console 17 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 are provided.

  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. The X-ray tube is a rotary anode type, which emits an electron beam from a filament in response to a voltage from a high voltage generator, and generates an X-ray by colliding the electron beam with a rotating anode rotating at a predetermined speed. To do. The rotating anode rotates to reduce deterioration due to the electron beam continuously hitting the fixed position. The colliding part of the electron beam is an X-ray focal point that emits X-rays. The collimator limits the irradiation field so as to shield a portion of the X-rays emitted from the X-ray tube that does not contribute to the examination region of the subject H.

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

The first absorption type grating 21 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 a direction (x direction) orthogonal to the grid line, thereby relative position of the second absorption type grating 22 with respect to the first absorption type grating 21. A scanning mechanism 23 for changing the above 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.

  Further, the imaging unit 12 includes a position sensor 24 that detects the positions of the first and second absorption gratings 21 and 22 in the x direction. The position sensor 24 is configured by, for example, a laser displacement sensor. The detection value of the position sensor 24 is supplied to the image processing unit 14 via the imaging control unit 16 and the system control unit 18. The position sensor 24 is not limited to a laser displacement sensor, and may be configured by an encoder, a potentiometer, a Hall element, an ultrasonic sensor, an acceleration sensor, or the like.

  The image processing unit 14 includes a phase differential image generation unit 14a and a phase contrast image generation unit 14b. The phase differential image generation unit 14 a 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. As will be described in detail later, the phase differential image generation unit 14a uses the detected value of the relative position supplied from the position sensor 24 when generating the phase differential image.

  The phase contrast image generation unit 14b generates a phase contrast image by integrating the phase differential image generated by the phase differential image generation unit 14a along the scanning direction (x direction). The phase contrast image 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 includes a liquid crystal display and a CRT display, and performs display based on information such as X-ray imaging conditions and a phase contrast image.

  In FIG. 2, the FPD 20 includes an image receiving unit 31 in which a plurality of pixels 30 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 32 controls the readout timing of the charges, and the readout circuit 33 that reads the charges from the pixels 30, converts the charges into image data, and outputs the image data. The scanning circuit 32 and each pixel 30 are connected by a scanning line 34 for each row, and the readout circuit 33 and each pixel 30 are connected by a signal line 35 for each column. The arrangement pitch of the pixels 30 is about 100 μm in each of the x direction and the y direction.

  The pixel 30 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 30 is provided with a TFT switch (not shown). The gate electrode of the TFT switch is connected to the scanning line 34, the source electrode is connected to the capacitor, and the drain electrode is connected to the signal line 35. When the TFT switch is turned on by the drive pulse from the scanning circuit 32, the charge accumulated in the capacitor is read out to the signal line 35.

Note that the pixel 30 temporarily converts X-rays into visible light with 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 33 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 the pixels 30 via the signal line 35 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.

In FIG. 3, 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 transmission 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 transmission 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 transmissive 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 linearly project X-rays that have passed through the X-ray transmission parts 21b and 22b 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 not diffracted. In addition, it is configured to pass through the X-ray transmission parts 21b and 22b 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 set to about ₁ μm to 10 μm, most of the X-rays are linearly projected without being diffracted. The grating pitches p ₁ and p ₂ are about ₂ μm to 20 μm.

The X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam having the X-ray focal point 11a as a light emission point, and therefore a projected image projected through the first absorption grating 21 (hereinafter referred to as a projection image). The projection image is referred to as a G1 image or a fringe image) and is enlarged in proportion to the distance from the X-ray focal point 11a. The grating pitch p ₂ and the interval d ₂ of the second absorption type grating 22 are such that the pattern of the X-ray transmission part 22 b substantially matches the periodic pattern of the bright part of the G1 image at the position of the second absorption type grating 22. Has been determined. That is, when the distance from the X-ray focal point 11a to the first absorption-type grating 21 is L ₁ and the distance from the first absorption-type grating 21 to the second absorption-type grating 22 is L ₂ , the grating pitch p ₂ and the distance d ₂ are determined so as to satisfy the relationship of the following expressions (1) and (2).

In the case of a Talbot interferometer, the distance L ₂ from the first absorption type grating 21 to the second absorption type grating 22 is a Talbot interference distance determined by the grating pitch of the first absorption type grating 21 and the X-ray wavelength. However, in the present embodiment, the first absorption type grating 21 projects the incident X-rays without diffracting, and the G1 image of the first absorption type grating 21 is the first absorption type. because similarly obtained at all positions of the back of the mold grating 21, 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, the grating pitch p ₁ of the first absorption-type grating _21, the grating pitch p _2, X-ray wavelength of the second absorption-type grating 22 (peak wavelength) lambda, and using the positive integer m Is expressed by the following equation (3).

  Equation (3) is an equation representing the Talbot interference distance when the X-rays emitted from the X-ray source 11 are cone beams. “Atsushi Momose, et al., Japanese Journal of Applied Physics, Vol. 47, No. 10, October 2008, p. 8077 ”.

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, intensity modulation is performed by superimposing the G1 image (stripe image) of the first absorption-type grating 21 and the second absorption-type grating 22. The striped image is captured by the FPD 20. There is a slight difference between the pattern period of the G1 image at the position of the second absorption grating 22 and the grating pitch p _{2 of} the second absorption grating 22 due to an arrangement error or the like. Due to this minute difference, moire fringes occur in the intensity-modulated fringe image. Further, when an error occurs in the lattice arrangement direction of the first and second absorption type gratings 21 and 22, and the arrangement directions are not the same, so-called rotational moire occurs. However, even when such moire fringes are generated in the fringe image, there is no particular problem as long as the period of the moire fringes in the x direction or y direction is larger than the arrangement pitch of the pixels 30.

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

  Next, the principle of the method for generating the phase contrast image 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 40 indicates an X-ray path that goes straight when the subject H does not exist. X-rays traveling along the path 40 pass through the first and second absorption gratings 21 and 22 and enter the FPD 20. Reference numeral 41 indicates an X-ray path refracted and deflected by the subject H when the subject H exists. X-rays traveling along the path 41 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) using the refractive index distribution n (x, z) of the subject H. Here, the y-coordinate is omitted for simplification of description.

  The G1 image projected from the first absorption grating 21 to the position of the second absorption grating 22 is displaced in the x direction by an amount corresponding to the refraction angle φ due to refraction of X-rays at the subject H. . 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 30 detected by the FPD 20 (the amount of phase shift of the intensity modulation signal of each pixel 30 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 30, 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, 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 embodiment, the phase shift amount ψ is calculated using the fringe scanning method shown below.

  In the fringe scanning method, an image is taken while one of the first and second absorption type gratings 21 and 22 is translated relative to the other in the x direction (that is, while changing the phase of both grating periods). Take a picture). In the present embodiment, the first absorption type grating 21 is fixed, and the second absorption type grating 22 is moved by the scanning mechanism 23.

As the second absorption type grating 22 moves, the moiré 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 22 (grating pitch p ₂ ). (Ie, when the phase change reaches 2π), the moire fringes return to their original positions. Thus, while moving the second absorption grating 22 by an integral fraction of the grating pitch p _2, taking a fringe image by FPD 20. At this time, the position sensor 24 detects the positions in the x direction by measuring the positions of the end faces of the first and second absorption gratings 21 and 22.

FIG. 4 schematically shows how the second absorption type grating 22 is moved by the scanning pitch (p ₂ / M) obtained by dividing the grating pitch p ₂ into M pieces. The scanning mechanism 23 translates the second absorption type grating 22 in order in M scanning steps of k = 0, 1, 2,..., M−1. M is an integer greater than or equal to 2, for example, M = 5.

  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 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 were not refracted by the subject H. While the X-ray component decreases, the X-ray component refracted by the subject H increases. In particular, at the position of k = M / 2, only the X-rays refracted by the subject H pass through the second absorption type grating 22. When the position of k = M / 2 is exceeded, the X-ray component passing through the second absorption grating 22 is refracted by the subject H while the X-ray component refracted by the subject H decreases. The missing X-ray component increases.

When photographing is performed by the FPD 20 at each scanning step k = 0, 1, 2,..., M−1, M pixel data are obtained for each pixel 30. FIGS. 5A and 5B illustrate changes (intensity modulation signals) of the pixel data I _k (x) obtained at each scanning step k when M = 5. Here, x is a coordinate in the x direction of the pixel 30. [delta] _k is the relative phase of the second absorption grating 22 relative to the first absorption grating 21 (hereinafter, referred to as relative phase) represents the, defined by the following equation (9).

Here, γ _k represents a relative phase shift amount due to a positional shift between the first and second absorption gratings 21 and 22. FIG. 5A illustrates pixel data I _k (x) when there is no positional deviation in the first and second absorption gratings 21 and 22 at each scanning step k, that is, when γ _k = 0. ing. On the other hand, FIG. 5B illustrates pixel data I _k (x) when a positional deviation occurs in the first and second absorption gratings 21 and 22.

The relative phase shift amount γ _k is expressed by the following equation (10), where α _k is the positional shift amount of the first absorption type grating 21 in the x direction and β _k is the positional shift amount of the second absorption type lattice 22 in the x direction. ). The positional deviation amounts α _k and β _k are calculated based on the detection value of the position sensor 24.

Hereinafter, a method for calculating the phase shift amount ψ (x) of the intensity modulated signal by the subject H based on the M pieces of pixel data I _k (x) will be described. In particular, a calculation method in consideration of the relative phase shift amount γ _k will be described. The pixel data I _k (x) is generally represented by the following equation (11).

A ₀ corresponds to the intensity of the incident X-ray, A _n is a value corresponding to the contrast of the intensity-modulated signal. Here, n is a positive integer and i is an imaginary number.

In the equation (11), ignoring higher-order terms of n ≧ 2, the pixel data I _k (x) is expressed by the following equation (12) as a sine wave.

Pixel data I _k (x) that satisfies Expression (12) is a theoretical value. On the other hand, the pixel data I _k (x) actually obtained by the FPD 20 includes an error, and a deviation occurs from the equation (12) by the error.

In order to calculate the phase shift amount ψ (x) from the actual measurement value of the pixel data I _k (x), first, the equation (12) is modified as shown in the following equation (13).

Here, a ₀ , a ₁ and a ₂ are represented by the following equations (14) to (16), respectively.

If a ₀ , a ₁ , and a ₂ are determined so as to minimize the difference between the theoretical value and the actual measurement value of the pixel data I _k (x) using the least square method or the like, the equation (17) is obtained. Thus, the phase shift amount ψ (x) is obtained from a ₁ and a ₂ .

The method of phase shift using the least square method is disclosed in “Introduction to Applied Optical Measurement, Toyohiko Yadagai, Second Edition, published on February 15, 2005, Maruzen Co., Ltd. (pages 196 to 198)” Yes. In the present embodiment, a ₀ , a ₁ , and a ₂ are determined by solving the following equation (18) obtained by the least square method.

Here, a, A (δ _k ), and B (δ _k ) are expressed by the following equations (19) to (21).

  In the above description, the y coordinate of the pixel 30 is not taken into consideration, but by performing the same calculation in consideration of the y coordinate of the pixel 30, a two-dimensional phase shift distribution ψ ( x, y) is obtained. This distribution ψ (x, y) is a phase differential image.

  In the above description, higher-order terms of n ≧ 2 or more are ignored in equation (12). However, since terms of n ≧ 2 are also expressed by a linear combination, the terms of n ≧ 2 are included. Needless to say, the equations (17) to (21) also hold in this case.

For each pixel 30, the phase differential image generation unit 14a applies a ₁ and a ₂ obtained by calculations based on the equations (18) to (21) to the equation (17), so that the phase differential image ψ (x, y) is generated. The phase contrast image generation unit 14 b generates a phase contrast image by performing integration processing along the x direction on the phase differential image ψ (x, y).

Next, the operation of the X-ray imaging system 10 configured as described above will be described. When the photographing conditions are input from the console 17 and the photographing is started, the scanning mechanism 23 moves the second absorption grating 22 by a predetermined scanning pitch (p ₂ / M) while scanning each step k. , X-ray exposure by the X-ray source 11 and detection operation by the FPD 20 are performed. As a result, M pieces of image data are generated by the FPD 20 and stored in the memory 13. In each scanning step k, the position sensor 24 detects the positions of the first and second absorption gratings 21 and 22 in the x direction, and the detected values are input to the image processing unit 14.

Next, the M pieces of image data stored in the memory 13 are read out by the phase differential image generation unit 14a. The phase differential image generation unit 14a performs calculations based on the equations (18) to (21) using the pixel data of each pixel 30 of each image data, and the obtained a ₁ and a ₂ are expressed by the equation (17). As a result, a phase differential image ψ (x, y) is generated.

At the time of this calculation, the phase differential image generation unit 14a, based on the detection value input from the position sensor 24, the amount of positional deviation between the first and second absorption gratings 21 and 22 in the x direction at each scanning step k. By calculating α _k and β _k and applying them to the equations (9) and (10), the relative positions δ _k of the first and second absorption gratings 21 and 22 at each scanning step _k are obtained, and the equations Applies to (20) and (21).

  Then, integration processing is performed by the phase contrast image generation unit 14b to generate a phase contrast image. The phase contrast image is recorded on the image recording unit 15 and then displayed on the monitor.

  In the present embodiment, the positions of the first and second absorption type gratings 21 and 22 are individually detected by the position sensor 24. However, instead of this, the first and second absorptions are detected. A position sensor may be fixed to one of the mold lattices 21 and 22, and the position of the other may be detected by detecting the other position with this position sensor.

  In the present embodiment, the initial position of the second absorption type grating 22 by the scanning mechanism 23 is a position where the dark part of the G1 image substantially coincides with the X-ray shielding part 22a when the subject H is not present (k = 0). However, the initial position is not limited to this, and any of k = 0, 1, 2,..., M−1 may be selected.

  In this embodiment, the phase contrast image is recorded in the image recording unit 15 and displayed on the monitor. In addition to the phase contrast image, the phase differential image is recorded in the image recording unit 15 and displayed on the monitor. It is also possible to configure as described above.

  In the present embodiment, the two-dimensional distribution of the phase shift of the intensity modulation signal is a phase differential image. However, if it has a proportional relationship with the differential value of the phase shift distribution Φ (x, y), the refraction angle φ A two-dimensional distribution of any physical quantity such as a phase differential image may be used.

  In this embodiment, 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 21 and the second absorption type. You may arrange | position between the grating | lattices 22. FIG. In this case as well, a phase differential image and a phase contrast image can be generated.

  In the following, other embodiments of the present invention will be described. In the following embodiments, the same reference numerals are used for the same configurations as those already described, and detailed description thereof is omitted.

(Second Embodiment)
In the first embodiment, the positions of the first and second absorption gratings 21 and 22 are detected. However, the position of the X-ray source 11 is further detected, and the image quality deteriorates due to the position shift of the X-ray source 11. You may comprise so that it may correct | amend.

  As shown in FIG. 6, the X-ray imaging system 50 of the present embodiment includes a first position sensor 51 that detects the positions of the first and second absorption gratings 21 and 22 in the x direction, and an X-ray source. 11 and a second position sensor 52 for detecting a position in the x direction. The first position sensor 51 has the same configuration as the position sensor 24 of the first embodiment.

As shown in FIG. 7, if the X-ray focal point 11a is moved by η _{k in the} x direction with the positional deviation of the X-ray source 11, the G1 image formed at the position of the second absorption grating 22 is The amount of movement ξ _k is expressed by the following equation (22) from the geometrical relationship.

That is, the movement of the X-ray focal point 11a by η _{k in the} x direction corresponds to the movement of the second absorption grating 22 by ξ _{k in the} same direction. When this movement amount ξ _k is a positional deviation amount from the reference position in each scanning step k, a relative phase deviation amount γ _k ′ of the G1 image with respect to the second absorption type grating 22 is expressed by the following equation (23). It is represented by

Therefore, in the present embodiment, the phase differential image generation unit 14a obtains the shift amount γ _k by the following equation (24) instead of the above equation (10), and performs the same calculation below, thereby obtaining the first and first The positional shift of the two absorption gratings 21 and 22 and the X-ray source 11 is corrected. Other configurations and operations are the same as those of the first embodiment.

  In the present embodiment, the positions of the first and second absorption gratings 21 and 22 and the X-ray source 11 are individually detected, but a position sensor is fixed to one of them. The position sensor may be configured to detect the other two positions.

(Third 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 focal spot size (generally about 0.1 mm to 1 mm) of the X-ray focal spot 11a affects the phase contrast. There is a risk of degrading the image quality. Therefore, as a second embodiment of the present invention, as shown in FIG. 8, a multi slit (ray source grid) 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 transmission part (not shown) is provided between the X-ray shielding parts 61 adjacent in the x direction.

In the present embodiment, the slit position of the multi-slit 60 becomes the X-ray focal point, and therefore the distance from the multi-slit 60 to the first absorption grating 21 corresponds to the distance L ₁ of the first embodiment. In the present embodiment, by detecting the position of the multi-slit 60 in the x direction with a position sensor, it is possible to correct image quality degradation due to the movement of the X-ray focal point as in the second embodiment. What is necessary is just to let the moving amount | distance regarding x direction be the moving amount (eta) _k of said Formula (22). If the arrangement pitch in the x direction of the X-ray shielding part 61 is p ₀ , it is necessary to satisfy the relationship of p ₀ = p ₂ · L ₁ / L ₂ from the geometrical relationship. Other configurations and operations are the same as those of the first embodiment.

(Fourth embodiment)
In each of the above embodiments, the first and second absorption type gratings 21 and 22 are configured to linearly project the X-rays that have passed through the X-ray transmission parts 21b and 22b. The configuration is not limited to the configuration, and a configuration described in Japanese Patent No. 4445397 in which a Talbot interference effect is generated by diffracting X-rays by the first absorption type grating 21 may be employed. In the present embodiment, the first absorption grating 21 and the diffraction grating constitute a Talbot interferometer to set the distance L ₂ between the first and second absorption gratings 21 and 22 to the Talbot distance. A fringe image (self-image) generated by the Talbot interference effect on the first absorption type grating 21 is projected onto the second absorption type grating 22.

  In this case, the first absorption type grating 21 can be a phase type grating. In order to use the first absorption type grating 21 as a phase type grating, a phase difference of “π” or “π / 2” is generated in the X-ray between the X-ray high absorption part 21 a and the X-ray transmission part 21 b. Thus, what is necessary is just to set thickness and material.

  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.

10 X-ray imaging system 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 transmission part 22 2nd absorption-type grating | lattice (2nd grating | lattice)
22a X-ray shielding part (X-ray high absorption part)
22b X-ray transmission part

Claims (9)

A radiation image detector for detecting radiation emitted from a radiation source;
First and second gratings disposed between the radiation source and the radiation image detector;
Scanning means for changing the relative positions of the first and second gratings in a direction perpendicular to the grating lines;
First position detecting means for detecting each relative position changed by the scanning means;
A phase differential image is obtained by calculating a phase shift amount of an intensity modulation signal for each pixel using image data obtained by the radiation image detector at each relative position and a detection value by the first position detection unit. A phase differential image generating means for generating;
A radiation imaging system comprising: A second position detecting means for detecting the position of the radiation source;
The phase differential image generation means uses the image data obtained by the radiation image detector at each relative position and the detection values by the first and second position detection means, and uses the phase of the intensity modulation signal for each pixel. The radiation imaging system according to claim 1, wherein a deviation amount is calculated.   The radiation imaging system according to claim 1, wherein the phase differential image generation unit calculates a phase shift amount of the intensity modulation signal using a calculation formula based on least squares.   The phase contrast image generation means which produces | generates a phase contrast image by integrating the phase differential image produced | generated by the said phase differential image production | generation means along the change direction of the said relative position is provided. The radiation imaging system according to any one of 3 to 3.   The first and second gratings are absorption gratings, and the first grating projects the radiation from the radiation source onto the second grating as a fringe image. 4. The radiographic system according to any one of 4 above.   The said 1st grating | lattice is a phase type | mold grating | lattice, and projects the radiation from the said radiation source on the said 2nd grating | lattice as a fringe image by the Talbot interference effect. The radiation imaging system described.   The radiation imaging system according to claim 1, further comprising a radiation source grid on an emission side of the radiation source. Further comprising second position detecting means for detecting the position of the source grid;
The phase differential image generation means uses the image data obtained by the radiation image detector at each relative position and the detection values by the first and second position detection means, and uses the phase of the intensity modulation signal for each pixel. The radiation imaging system according to claim 7, wherein a deviation amount is calculated. A radiation image detector for detecting radiation emitted from a radiation source; first and second gratings disposed between the radiation source and the radiation image detector; and the first and second gratings In an image processing method used in a radiation imaging system, comprising: a scanning unit that changes a relative position in a direction orthogonal to a lattice line; and a position detection unit that detects each relative position changed by the scanning unit.
A phase differential image is generated by calculating a phase shift amount of an intensity modulation signal for each pixel using image data obtained by the radiation image detector at each relative position and a detection value by the position detection unit. An image processing method characterized by the above.

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