WO2012169427A1 - Radiography system - Google Patents

Abstract

An x-ray imaging system (10) provided with an imaging unit (12) and a computing/processing unit (22), wherein: the imaging unit (12) has grids (31, 32) for using transmitting radioactive rays to generate an x-ray image including periodic intensity distribution, and an x-ray image detector (30) for detecting the x-ray image; and the grid (31) is configured by division into a plurality of modules (33). The computing/processing unit (22): divides the x-ray image into a plurality of sectional x-ray images (Img_{m, n}), according to the boundaries between the modules in the grid (31); for each of these sectional x-ray images, generates a sectional phase contrast image by analyzing the period pattern included in the sectional x-ray image using a Fourier transform and an inverse Fourier transform; and generates a phase contrast image of the imaging subject by bonding the obtained plurality of sectional phase contrast images.

Description

Radiography system

The present invention relates to a radiation imaging system.

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

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

However, the X-ray absorptivity becomes lower as a substance composed of an element having a smaller atomic number, and the difference in the X-ray absorptivity is small in a soft tissue or soft material of a living body. Therefore, a sufficient image density as an X-ray transmission image is obtained. There is a problem that (contrast) cannot be obtained. For example, most of the components of the cartilage part constituting the joint of the human body and the joint fluid in the vicinity thereof are water, and the difference in the amount of X-ray absorption between the two is small, so that it is difficult to obtain image contrast.

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

In the above X-ray Talbot interferometer, a first diffraction grating (phase type grating or absorption type grating) is arranged behind the subject, and a specific distance (Talbot interference determined by the grating pitch of the first diffraction grating and the X-ray wavelength is set. A second diffraction grating (absorption type grating) is arranged downstream by a distance), and an X-ray image detector is arranged behind the second diffraction grating. The Talbot interference distance is a distance at which the X-rays that have passed through the first diffraction grating form a self-image that exhibits a periodic intensity distribution due to the Talbot interference effect. Are modulated by the interaction (phase change) between the subject and the X-rays arranged between the diffraction gratings.

In the above X-ray Talbot interferometer, the moire generated by the superposition of the self-image of the first diffraction grating and the second diffraction grating is detected, and the periodic pattern appearing in the image corresponding to the moire is modulated by the subject. By analyzing, the phase information of the object is acquired. For example, a fringe scanning method is known as a method for analyzing a periodic pattern appearing in an image. According to this fringe scanning method, the second diffraction grating is substantially parallel to the surface of the first diffraction grating with respect to the first diffraction grating and substantially in the grating direction (strip direction) of the first diffraction grating. X-rays refracted by the subject from a change in signal value for each corresponding pixel between a plurality of image data obtained by performing a plurality of times of imaging while translating in a vertical direction with a scanning pitch obtained by equally dividing the lattice pitch. Angle distribution (differential image of phase shift) can be obtained, and a phase contrast image of the subject can be obtained based on this angle distribution.

However, according to the above-described fringe scanning method, it is necessary to perform photographing a plurality of times, and there is a concern about the movement of the subject during photographing and the resulting deterioration in image quality. In view of this, a method has been proposed in which the phase information of the subject is acquired by one shooting by using Fourier transform and inverse Fourier transform (see, for example, Patent Document 2). This is done by separating the frequency domain containing the fundamental frequency component of the periodic pattern from the spatial frequency spectrum obtained by Fourier transforming the image containing the periodic pattern, and performing the inverse Fourier transform on the separated frequency domain. A differential image is acquired. According to this, the movement of the lattice between a plurality of times of photographing and the moving mechanism that requires high accuracy are unnecessary, so that the photographing workflow can be improved and the apparatus can be simplified. In addition, it is possible to eliminate the deterioration in image quality caused by the movement of the subject between each photographing.

In the X-ray imaging systems described in Patent Documents 1 and 2, the first and second diffraction gratings need to be correspondingly large in order to expand the field of view. However, the first and second diffraction gratings typically need to be configured with a high aspect ratio with a grating pitch on the order of μm, so that it is very difficult to manufacture a large size and high accuracy. is there. Therefore, an X-ray imaging system in which each of the first and second diffraction gratings and the X-ray image detector is divided into a plurality of grating modules or a plurality of detector modules has been proposed ( (See Patent Document 3).

In addition, the self-image of the first diffraction grating is detected using a detector having a pixel pitch smaller than the period of the periodic intensity distribution of the self-image of the first diffraction grating without using the second diffraction grating, An X-ray imaging system has also been proposed in which phase information of a subject is acquired by analyzing the modulation of the periodic intensity distribution of the self-image (see Patent Document 4).

International Publication No. 04/058070 International Publication No. 10/0504843 Japanese Unexamined Patent Publication No. 2007-203061 Japanese Laid-Open Patent Publication No. 2007-203063

The X-ray imaging system described in Patent Document 4 detects a self-image of the first diffraction grating using a detector having a pixel pitch smaller than the period of the periodic intensity distribution of the self-image of the first diffraction grating. The phase information is obtained by analyzing this, and the spatial resolution is excellent because the pixel pitch is small. Further, since the second diffraction grating is not interposed, the accuracy of the phase information can be improved. However, a detector with a small pixel pitch is limited to a relatively small size, and the field of view is limited. Further, when the size of the detector is increased, typically, the S / N tends to decrease, and there is a concern that the accuracy of the phase information may decrease due to the decrease in S / N.

Therefore, if the detector is divided into a plurality of detector modules as in the X-ray imaging system described in Patent Document 3, it is possible to achieve a large pixel pitch and excellent S / N. A size detector can be obtained. However, it is very difficult to arrange all the detector modules in exactly the same relative positional relationship with respect to the first diffraction grating, and such a mismatch in relative positional relationship is caused by the period and direction of the periodic pattern in the image. Appears as a misalignment.

Further, when analysis is performed using Fourier transform and inverse Fourier transform on an image in which the period and direction of the periodic pattern in the image are not uniform, there is a possibility that the phase information of the subject cannot be obtained accurately. This is not limited to a case where a plurality of detector modules are connected and configured, and is also applicable to a case where a plurality of grating modules are connected to form the first diffraction grating and the second diffraction grating.

The present invention has been made in view of the above-described circumstances, and an object thereof is to improve the accuracy of phase information in radiation phase imaging that acquires phase information of a subject using Fourier transform and inverse Fourier transform.

An imaging unit that acquires a radiation image including a periodic pattern modulated by a subject arranged in a radiation field, and an arithmetic processing unit that generates a phase contrast image of the subject based on the periodic pattern included in the radiation image The imaging unit includes, by radiation passing therethrough, one or more gratings that form a radiation image including a periodic intensity distribution that is a basis of the periodic pattern included in the radiation image, and the radiation image A radiation image detector for detecting, and at least one of the lattice and the elements of the radiation image detector is divided into a plurality of modules, and the arithmetic processing unit includes: The radiation image corresponding to at least a part of a boundary between modules in at least one of the elements divided into modules. Is converted into a plurality of partial radiographic images, and for each partial radiographic image, a Fourier transform is performed on the partial radiographic image to obtain a spatial frequency spectrum of the partial radiographic image; Partial phase contrast image generation that separates a spatial frequency region including the fundamental frequency component of the included periodic pattern from the spatial frequency spectrum and generates a partial phase contrast image by performing an inverse Fourier transform on the separated spatial frequency region And a combination processing for combining the plurality of partial phase contrast images generated by the partial phase contrast image generation processing to generate a phase contrast image of the subject.

According to the present invention, by dividing the grating and the radiographic image detector into a plurality of modules, it is possible to obtain a large-size grating without reducing the accuracy thereof. Can be obtained in a large size without reducing its S / N, and the field of view can be easily enlarged.

Then, the radiographic image acquired by the radiographic image detector is divided into partial X-ray images corresponding to the boundaries between the modules in the module-divided elements, and the periodic pattern is analyzed for each partial X-ray image to analyze the subject. By acquiring the phase shift distribution, non-uniformity in the period and orientation of the periodic pattern between the partial X-ray images caused by the relative positional deviation of each of the modules of this element relative to the other elements is obtained. The influence on the phase shift distribution can be eliminated or reduced, and the accuracy of the obtained phase shift distribution of the subject can be increased.

It is a schematic diagram which shows the structure of an example of the radiography system for describing embodiment of this invention. It is a control block diagram of the radiography system of FIG. It is a perspective view which shows the structure of the imaging | photography part of the radiography system of FIG. It is a side view which shows the structure of the imaging | photography part of the radiography system of FIG. It is a schematic diagram which shows the structure of the radiographic image detector contained in the imaging | photography part of FIG. It is a schematic diagram which shows the mechanism for changing the periodic pattern of the image acquired by the radiographic image detector of FIG. It is a schematic diagram for demonstrating the refraction | bending of the radiation by a to-be-photographed object. FIG. 6 is a schematic diagram illustrating an example of dividing an image acquired by the radiological image detector of FIG. 5 into partial images and processing for generating a phase contrast image from the divided partial images. It is a schematic diagram which shows the other example which divides the image acquired by the radiographic image detector of FIG. 5 into a partial image. It is a schematic diagram which shows the other example which divides the image acquired by the radiographic image detector of FIG. 5 into a partial image. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows an example which classifies the image acquired by the radiographic image detector of the radiography system of FIG. 12 into a partial image. It is a schematic diagram which shows the structure of the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention.

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

The X-ray imaging system 10 is an X-ray diagnostic apparatus that images a subject H in a standing position, and is disposed opposite to the X-ray source 11 that radiates X-rays to the subject H, and the X-ray source 11. 11 controls the exposure operation of the X-ray source 11 and the imaging operation of the imaging unit 12 based on the operation of the operator, and detects the X-ray transmitted through the subject H from 11 and generates image data. It is broadly divided into a console 13 that generates a phase contrast image by calculating the image data acquired by the imaging unit 12.

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

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

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

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

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

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

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

The imaging unit 12 is provided with an X-ray image detector 30 and a first absorption-type grating 31 and a second absorption-type grating 32 for detecting phase change of the X-ray caused by the subject H and performing phase imaging. .

3 and 4 schematically show the configuration of the photographing unit 12.

The X-ray image detector 30 is arranged so that the detection surface is orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11. The first and second absorption type gratings 31 and 32 are disposed between the X-ray image detector 30 and the X-ray source 11.

The first absorption type grating 31 is configured by connecting a plurality of first grating modules 33, and the second absorption type grating 32 is also configured by connecting a plurality of second grating modules 36. ing. In the illustrated example, the first absorption type grating 31 includes four first grating modules 33 arranged in the x and y directions in a plane orthogonal to the optical axis A, and adjacent to each other. 33 are connected to each other. In the second absorption type grating 32, five second grating modules 36 are arranged in the x direction and the y direction in a plane orthogonal to the optical axis A, and adjacent second grating modules 36 are connected to each other. Configured. The arrangement of the first grating modules 33 in the first absorption type grating 31 and the arrangement of the second grating modules 36 in the second absorption type grating 32 are not limited to the above example, and the x direction or the y direction. It may be a one-dimensional array.

Each of the first lattice modules 33 includes a substrate 34 and a plurality of X-ray shielding portions 35 disposed on the substrate 34. Each of the second grating modules 36 includes a substrate 37 and a plurality of X-ray shielding portions 38 disposed on the substrate 37. The substrates 34 and 37 are each formed of an X-ray transmissive member such as silicon, glass, or resin that transmits X-rays.

The X-ray shielding portions 35 and 38 are linear members extending in one direction (y direction in the illustrated example) in a plane perpendicular to the optical axis A of the X-rays emitted from the X-ray source 11. Composed. As a material of each X-ray shielding part 35 and 38, what is excellent in X-ray absorptivity is preferable, for example, it is preferable that they are heavy metals, such as gold | metal | money and platinum. These X-ray shielding portions 35 and 38 can be formed by a metal plating method or a vapor deposition method.

The X-ray shielding portions 35 are arranged in a plane orthogonal to the optical axis A of the X-rays at a predetermined interval p ₁ in the direction orthogonal to the one direction (x direction) with a predetermined interval d _1. ing. Similarly, the X-ray shields 38 are spaced from each other at a predetermined interval d ₂ in a direction orthogonal to the one direction (x direction) in a plane orthogonal to the optical axis A of X-rays with a constant period p _2. Are arranged. Such first and second absorption gratings 31 and 32 do not mainly give a phase difference to incident X-rays but give an intensity difference, and are also called amplitude gratings. Note that the slit portions (regions having the distances d ₁ and d ₂ ) may not be voids, and the voids may be filled with an X-ray low-absorbing material such as a polymer or a light metal.

The first absorption type grating 31 is configured to geometrically project X-rays that have passed through the slit portion regardless of the presence or absence of the Talbot interference effect. Specifically, by setting the interval d ₁ to a value sufficiently larger than the effective wavelength of the X-rays emitted from the X-ray source 11, most of the X-rays irradiated do not undergo diffraction at the slit portion. In addition, a self-projected image (hereinafter, this projected image is referred to as a self-image G1) can be formed behind the first absorption type grating 31. For example, when tungsten is used as the target material of the rotary anode 18a and the tube voltage is 50 kV, the effective wavelength of X-ray is about 0.4 mm. In this case, if the distance d ₁ is set to about 1 to 10 μm, the X-ray image formed by the X-rays that have passed through the slit portion is negligible for the diffraction effect. A self-image G1 is formed behind.

In general, the X-ray radiated from the X-ray source 11 is not a parallel beam but a cone beam with the X-ray focal point 18b as a light emission point, and the self-image G1 is at a distance from the X-ray focal point 18b. Scaled proportionally. On the other hand, the grating pitch p ₂ of the second absorption type grating 32 is determined so that the slit portion substantially coincides with the periodic intensity distribution of the self-image G 1 at the position of the second absorption type grating 32. That is, when the distance from the X-ray focal point 18b to the first absorption-type grating 31 is L ₁ and the distance from the first absorption-type grating 31 to the second absorption-type grating 32 is L ₂ , the first The lattice pitch p ₁ of the absorption-type grating 31, the pitch p ₁ ′ of the self-image G 1 at the position of the second absorption-type grating 32, and the lattice pitch p ₂ of the second absorption-type grating 32 are expressed by the following equation (1). It is determined to satisfy.

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

As described above, the imaging unit 12 does not constitute a Talbot interferometer, but the Talbot interference distance Z when it is assumed that X-rays are diffracted by the first absorption type grating 31 is the first absorption type grating. 31 grating pitch p ₁ , second absorbing grating 32 grating pitch p ₂ , X-ray wavelength (usually the effective wavelength of X-rays incident on first absorbing grating 31) λ, and positive integer m. And is represented by the following formula (2).

Formula (2) is a formula that represents the Talbot interference distance when the X-rays emitted from the X-ray source 11 are cone beams. “Atsushi Momose, et al., Japan Journal of Applied Physics, Vol. 47, No. 10, October 2008, page 8077 ”and“ Timm Weitkamp, et al., Proc. Of SPIE, Vol. 6318, 2006, 63180S-1 ”can be easily derived.

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

In the case where X-rays emitted from the X-ray source 11 can be regarded as substantially parallel beams, Talbot interference distance Z by the following equation (4), and the distance L _2, the range satisfying the following equation (5) Can be set to a value.

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

The X-ray shields 35 and 38 preferably shield (absorb) X-rays completely in order to generate an X-ray image having a periodic intensity distribution with high contrast, but are excellent in the X-ray absorbability described above. Even if materials (gold, platinum, etc.) are used, 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 make the thicknesses h1 and h2 of the X-ray shielding portions 35 and 38 as thick as possible. The X-ray shields 35 and 38 preferably shield 90% or more of the irradiated X-rays, and the thickness thereof is set according to the energy of the irradiated X-rays. For example, when tungsten is used as the target material of the X-ray tube 18 and the tube voltage is 50 kV, the thicknesses h ₁ and h ₂ are preferably 30 μm or more in terms of gold (Au).

However, when the X-rays irradiated from the X-ray source 11 are cone beams, if the thicknesses h ₁ and h ₂ of the X-ray shielding portions 35 and 38 are excessively increased, the X-rays incident obliquely enter the slit portion. There is a problem that it becomes difficult to pass, so-called vignetting occurs, and the effective visual field in the direction (x direction) perpendicular to the extending direction (strand direction) of the X-ray shielding portions 35 and 38 becomes narrow. For this reason, it is preferable to limit the thicknesses h ₁ and h ₂ from the viewpoint of securing a visual field. Specifically, when the length of the effective visual field in the x direction on the detection surface of the X-ray image detector 30 is V, and the distance from the X-ray focal point 18b to the detection surface of the X-ray image detector 30 is L, the thickness h ₁ and h ₂ are preferably set so as to satisfy the following expressions (6) and (7) from the geometrical relationship shown in FIG.

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

In the above configuration, the second absorption grating 32 is superimposed on the self-image G 1 of the first absorption grating 31, and the X-ray image detector 30 is arranged immediately after the second absorption grating 32. An X-ray image is formed. The period p ₁ ′ of the periodic intensity distribution of the self-image G1 at the position of the second absorption type grating 32 and the substantial grating pitch p ₂ ′ of the second absorption type grating 32 are due to manufacturing errors and arrangement errors. Some differences occur. Among these, the arrangement error means that the substantial pitch in the x direction changes due to the relative inclination and rotation of the first and second absorption gratings 31 and 32 and the distance between the two changes. I mean.

Moire occurs due to a slight difference between the period p ₁ ′ of the periodic intensity distribution of the self-image G1 and the grating pitch p ₂ ′ of the second absorption grating 32. The period T of the moire in the x direction is expressed by the following equation (8).

Since the moire period T is larger than the period p1 ′ of the periodic intensity distribution of the self-image G1, the pixel arrangement pitch required for resolving the moire period T with respect to the X-ray image detector 30 is increased. Restrictions are relaxed. Therefore, in this example, the X-ray image detector 30 has a relatively coarse pixel arrangement pitch, but a TFT (Thin) that can relatively easily constitute a large detector with a single module. An FPD based on a film transistor panel is used.

FIG. 5 schematically shows the configuration of the X-ray image detector 30.

The X-ray image detector 30 includes a plurality of pixels 40 that convert X-rays into electric charges and accumulate them on an insulating substrate such as a glass substrate, and a plurality of readout circuits that read out the electric charges accumulated in each pixel 40. TFT switches (not shown) are two-dimensionally arranged in the xy direction, a scanning circuit 42 that controls the timing of reading charges from the image receiving section 41, and the charges read from each pixel 40. The signal processing circuit 43 converts and stores the image data, and the data transmission circuit 44 transmits the image data to the arithmetic processing unit 22 via the I / F 25 of the console 13.

Each pixel 40 is a direct conversion type in which X-rays are directly converted into electric charges by a conversion layer (not shown) such as amorphous selenium and the converted electric charges are stored in a capacitor (not shown) connected to the lower electrode. It can comprise as an element of this. A TFT switch is connected to each pixel 40, a gate electrode of the TFT switch is connected to the scanning line 45, a source electrode is connected to the capacitor, and a drain electrode is connected to the signal line 46. When the TFT switch is turned on by the drive pulse from the scanning circuit 42, the charge accumulated in the capacitor is read out to the signal line 46.

Each pixel 40 once converts X-rays into visible light by a scintillator (not shown) made of terbium activated gadolinium oxide (Gd ₂ O ₂ S: Tb), thallium activated cesium iodide (CsI: Tl), or the like. It is also possible to configure as an indirect conversion type X-ray detection element that converts the converted visible light into a charge by a photodiode (not shown) and accumulates it.

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

The X-ray image detector 30 is not limited to an FPD based on a TFT panel, and various X-ray image detectors based on a solid-state imaging device such as a CCD sensor or a CMOS sensor can also be used. .

The moiré period T in equation (8) is actually further enlarged by the distance from the second absorption grating 32 to the detection surface of the X-ray image detector 30, and therefore the detection surface of the X-ray image detector 30. The upper moire period is T ′, and in order to detect this moire by the X-ray image detector 30, the arrangement pitch P of the pixels 40 in the x direction is at least on the detection surface of the X-ray image detector 30. Is not an integral multiple of the moire period T ′.

Further, it is possible to detect moire even if the arrangement pitch P is larger than the moire period T ′, but the arrangement pitch P is preferably smaller than the moire period T ′, and preferably satisfies the following equation (9). This is because, in order to obtain a high-quality phase contrast image, it is preferable that moire is detected with high contrast in the phase contrast image generation process described later.

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

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

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

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

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

In the X-ray imaging system 10, imaging unit 12 is not the Talbot interferometer as described above, since the distance L ₂ can be freely set, moire by changing the distance L ₂ as relative movement mechanism 52 A mechanism for changing the period T can be suitably employed. The change mechanism (relative rotation mechanism 50, relative tilt mechanism 51, and relative movement mechanism 52) of the first and second absorption type gratings 31 and 32 for changing the moiré period T ′ is an actuator such as a piezoelectric element. Can be configured. In the above, the arrangement pitch of the pixels 40 in the x direction and the moire period have been described. However, the same applies to the arrangement pitch of the pixels 40 and the moire period in the y direction, and the arrangement pitch of the pixels 40 in the y direction is the moire. The period of the arrangement pitch of the pixels 40 and the period of the moire in the y direction is preferably reduced by a mechanism similar to the above-described changing mechanism (the relative rotation mechanism 50, the relative tilt mechanism 51, and the relative movement mechanism 52). The relationship can also be adjusted.

When the subject H is arranged between the X-ray source 11 and the first absorption type grating 31, the moire formed on the X-ray image detector 30 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. An image acquired by detecting the moire by the X-ray image detector 30 includes a periodic pattern corresponding to the moire, and by analyzing the periodic pattern, a phase contrast image of the subject H can be generated. it can.

Next, a method for analyzing the periodic pattern of images will be described.

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

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

The phase shift distribution Φ (x) of the subject H is expressed by the following equation (10), 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 refraction angle φ is expressed by the equation (11) using the X-ray wavelength λ and the phase shift distribution Φ (x) of the subject H.

Since the refraction angle φ (x) is a value corresponding to the differential value of the phase shift distribution as shown in the equation (11), the refraction angle φ (x) is integrated along the x-axis to obtain the phase shift. A distribution Φ (x) is obtained. In the above description, the y coordinate in the y direction of the pixel 40 is not taken into consideration. However, by performing the same calculation for each y coordinate, a two-dimensional phase shift distribution Φ (x , Y).

Here, the moire formed by the first and second absorption type gratings 31 and 32, that is, the periodic pattern of the image can be expressed by the following equation (12), and the equation (12) is expressed by the following equation (13). Can be rewritten.

In Expression (12), a (x, y) represents the background, b (x, y) represents the amplitude of the spatial frequency component corresponding to the basic period of the periodic pattern, and (f _0x, f _0y ) represents the period. Represents the basic period of the pattern. In the formula (13), c (x, y) is represented by the following formula (14).

Therefore, information on the refraction angle φ (x, y) can be obtained by extracting the component of c (x, y) or c ^* (x, y). Here, equation (13) becomes the following equation (15) by Fourier transform.

In the formula _{(15), F (f x} , f y), A (f x, f y), C (f x, f y) , respectively f (x, y), a (x, y), c It is a two-dimensional Fourier transform for (x, y).

When using the one-dimensional lattice such as the first and second absorption type gratings 31 and 32, the spatial frequency spectrum of the image, at least, a peak derived from _{A (f x, f y)} , this _{C (f} x, f _y) and ^{_{C * (f x, f y}} ) 3 peaks and the peak of the spatial frequency component corresponding to the fundamental period of the periodic pattern from the results across. A _(f x, f _y) peak derived from the origin, _{also, C (f x, f y} ) and ^{_{C * (f x, f y}} ) peak derived from the _{(± f 0x, ± f 0y} ) It occurs at the position of (combined same order).

In order to obtain the refraction angle φ (x, y) from the spatial frequency spectrum of the image, a region including the peak frequency of the spatial frequency component corresponding to the basic period of the periodic pattern is cut out so that the peak frequency overlaps the origin of the frequency space. Move the clipped area and perform inverse Fourier transform. Then, the refraction angle φ (x, y) can be obtained from the complex number information obtained by the inverse Fourier transform.

Here, the first absorption type grating 31 is configured by connecting a plurality of first grating modules 33, and the second absorption type grating 32 is also configured by connecting a plurality of second grating modules 36. ing. Therefore, the relative positional relationship of each of the first grating modules 33 with respect to the second absorption grating 32 and the X-ray image detector 30 may be shifted, and similarly, the first absorption grating 31 and the X-ray image detection are performed. In some cases, the relative positional relationship of each of the second grating modules 36 with respect to the vessel 30 is shifted. Such a mismatch in the relative positional relationship appears as a shift in the period and direction of the periodic pattern in each part of the image acquired by the X-ray image detector 30. Therefore, it corresponds to at least part of the boundary between the first grating modules 33 in the first absorption type grating 31 and / or at least part of the boundary between the second grating modules 36 in the second absorption type grating 32. Thus, the X-ray image acquired by the X-ray image detector 30 is divided into a plurality of partial X-ray images, and the periodic pattern is analyzed for each of these partial X-ray images.

FIG. 8 shows an example of X-ray image segmentation and phase contrast image generation processing based on the segmented partial X-ray image.

The example shown in FIG. 8 corresponds to all the boundaries between the first grating modules 33 in the first absorption type grating 31, and the X-ray image is a total of 16 partial X-ray images Img _1,1 of 4 × 4. , Img _2,1 ,..., Img _4,4 (FIG. 8A). Specifically, in the first absorption type grating 31, the boundary between the first grating modules 33 extends in the x-direction boundary lines Lx ₁ , Lx ₂ , Lx ₃ (see FIG. 3) and the y-direction. A boundary line Lx ′ ₁ that is defined by the boundary lines Ly ₁ , Ly ₂ , Ly ₃ (see FIG. 3) and is a projection of these boundary lines Lx _{1 to 3} , Ly _{1 to 3} onto the X-ray image detector 30. _~ 3, Ly _{'1 ~} moiety X-ray image along the ₃ X-ray image _{Img 1,1, Img 1,2, ···,} are divided into _{Img 4, 4.}

Then, by analyzing the periodic pattern for each of the divided partial X-ray images Img _1,1 , Img _2,1 ,..., Img _4,4 , partial phase shift distributions Φ _1,1 , Φ _2,1 ,..., Φ _4,4 are obtained. Then, by combining the obtained partial phase shift distributions Φ _1,1 , Φ _2,1 ,..., Φ _4,4 , the phase shift distribution Φ of the subject H is obtained (FIG. 8B).

As described above, the periodic pattern is analyzed for each partial X-ray image obtained by dividing the X-ray image corresponding to the boundary between the first grating modules 33 in the first absorption type grating 31, and the phase shift distribution of the subject H is analyzed. By obtaining Φ, at least the period between the partial X-ray images caused by the displacement of the relative positions of the first grating module 33 with respect to the second absorption grating 32 and the X-ray image detector 30. The influence of the non-uniformity of the pattern period and orientation on the phase shift distribution Φ of the subject H can be eliminated or reduced, and the accuracy of the obtained phase shift distribution Φ of the subject H can be improved.

The above processing is executed by the arithmetic processing unit 22, and the arithmetic processing unit 22 stores the phase contrast image obtained by imaging the phase shift distribution Φ (x, y) in the storage unit 23. The above-described phase contrast image generation processing is automatically performed by the respective units operating in conjunction with each other under the control of the control device 20 after an imaging instruction is given from the input device 21 by the operator. A phase contrast image is displayed on the monitor 24.

FIG. 9 shows another example of X-ray image segmentation.

The example shown in FIG. 9 corresponds to all the boundaries between the first grating modules 33 in the first absorption type grating 31 and all the boundaries between the second grating modules 36 in the second absorption type grating 32. The X-ray images are divided into partial X-ray images Img _1,1 , Img _2,1 ,..., Img _8,8 .

Specifically, in the first absorption type grating 31, the boundaries between the first grating modules 33 are boundary lines Lx ₁ , Lx ₂ , Lx ₃ extending in the x direction and boundary lines Ly ₁ , extending in the y direction. It is defined by Ly ₂ and Ly ₃ . Further, in the second absorption type grating 32, the boundary between the second grating modules 36 extends in the boundary lines Lx ₄ , Lx ₅ , Lx ₆ , Lx ₇ (see FIG. 3) extending in the x direction, and the y direction. It is defined by the boundaries Ly ₄ , Ly ₅ , Ly ₆ , Ly ₇ (see FIG. 3). An X-ray image is converted into a partial X-ray image along the boundary lines Lx ′ _{1 to 7} and Ly ′ _{1 to 7} which are projections of these boundary lines Lx _{1 to 7} and Ly _{1 to 7 on} the X-ray image detector 30. Img _1,1 , Img _2,1 ,..., Img _8,8 . That is, Img _1,1 corresponds to an image of an X-ray image formed by X-rays that have passed through the first grating module 33 _1,1 and the second grating module 36 _1,1 . Img _2,1 corresponds to an image of an X-ray image formed by X-rays that have passed through the first grating module 33 _1,1 and the second grating module 36 _2,1 .

According to this, in addition to the displacement of the relative position of each of the first grating modules 33 with respect to the second absorption grating 32 and the X-ray image detector 30, the first absorption grating 31 and the X-ray image detection are further performed. Of non-uniformity of the period and direction of the periodic pattern between partial X-ray images caused by the relative positional deviation of each of the second grating modules 36 relative to the detector 30 on the phase shift distribution Φ of the subject H Can be eliminated or reduced, and the accuracy of the obtained phase shift distribution Φ of the subject H can be further enhanced.

FIG. 10 shows another example of the division of the X-ray image.

The example shown in FIG. 10 corresponds to a part of the boundary between the first grating modules 33 in the first absorption type grating 31, and X-ray images are divided into 2 × 2 total of four partial X-ray images Img _{1, 1} , Img ₂ , ₁ , Img ₁ , ₂ , Img ₂ , ₂ . Specifically, in the first absorption type grating 31, the boundary between the first grating modules 33 extends in the x-direction boundary lines Lx ₁ , Lx ₂ , Lx ₃ (see FIG. 3) and the y-direction. Projection of Lx ₂ and Ly _{2 on} the X-ray image detector 30 out of these boundary lines Lx _{1 to 3} and Ly _{1 to 3} defined by the boundary lines Ly ₁ , Ly ₂ , and Ly ₃ (see FIG. 3) The X-ray images are divided into partial X-ray images Img _1,1 , Img _2,1 , Img _1,2 , Img _2,2 along the boundary lines Lx ′ ₂ and Ly ′ ₂ .

According to this, as compared with the example shown in FIG. 8, it is possible to reduce the number of sections into partial X-ray images to reduce the analysis process, and to increase the processing speed in the arithmetic processing unit 22. Furthermore, since there is a relationship that the image size and frequency resolution are proportional in discrete Fourier transform, if the number of sections is excessively increased and the image size is decreased too much, the frequency resolution after Fourier transform is reduced and the resulting phase shift The accuracy of the partial distribution Φ may be reduced. By appropriately adjusting the number of sections in this way, a frequency resolution proportional to the image size is ensured, and the accuracy of the obtained phase shift distribution Φ is ensured. You can also.

As described above, according to the radiation imaging system 10, it is made to correspond to the boundary between modules in an element divided into a plurality of modules such as the first absorption type grating 31 and the second absorption type grating 32. The X-ray image is divided into partial X-ray images, and the periodic pattern is analyzed for each partial X-ray image to obtain the phase shift distribution Φ of the subject H, whereby each of the modules constituting the module-divided elements To eliminate or reduce the influence of the non-uniformity of the period and direction of the periodic pattern between the partial X-ray images caused by the relative positional relationship with respect to other elements on the phase shift distribution Φ of the subject H The accuracy of the obtained phase shift distribution Φ of the subject H can be improved.

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

In the X-ray imaging system 10 described above, the X-ray image detector 30 is configured as a single module, but a plurality of detector modules are provided as in the first and second absorption gratings 31 and 32. It can also be divided into two parts. In that case, when the image acquired by the X-ray image detector 30 is divided into a plurality of partial images, the X-ray image detector 30 is classified according to part or all of the boundary between the detector modules. You may do it. According to this, the period and direction of the moire between partial X-ray images caused by the relative positional relationship of the detector modules with respect to the first absorption type grating 31 and the second absorption type grating 32 are not uniform. The influence of the aspect on the phase shift distribution Φ of the subject H can be eliminated or reduced.

In addition, the second grating is superimposed on the projection image of the first grating to generate moire, and therefore, the first and second gratings are both absorbing gratings. However, the present invention is not limited to this. As described above, the present invention is also useful when the moire is generated by superimposing the second grating on the Talbot interference image. Therefore, the first grating is not limited to the absorption type grating but may be a phase type grating.

In addition, the image obtained by imaging the phase shift distribution Φ is described as being stored or displayed as a phase contrast image. The phase shift distribution Φ is obtained by integrating the differential amount of the phase shift distribution Φ corresponding to the refraction angle φ. Therefore, the differential amount of the refraction angle φ and the phase shift distribution Φ is also related to the phase change of the X-ray by the subject. Therefore, an image of the refraction angle φ and an image of the differential amount of the phase shift are also included in the phase contrast image.

Also, a phase differential image (differential amount of phase shift distribution) is created from moire obtained by photographing (pre-photographing) in the absence of a subject, and obtained by photographing (main photographing) in the presence of the subject. You may make it correct | amend the phase differential image produced from the moire. The phase differential image obtained by the pre-imaging reflects the device-specific phase unevenness (for example, the grating pitch and thickness non-uniformity of the absorption grating 31). On the other hand, the phase differential image acquired by the main imaging also includes the same type of device-specific phase unevenness as the pre-imaging, and acts as an offset of the phase differential signal. Therefore, by subtracting the phase differential image obtained by the pre-photographing from the phase differential image obtained by the main imaging, it is possible to obtain a phase contrast in which the phase unevenness specific to the apparatus is corrected.

FIG. 11 shows the configuration of another example of the radiation imaging system according to the present invention. In addition, description is abbreviate | omitted or simplified by attaching | subjecting a common code | symbol to the element which is common in the X-ray imaging system 10 mentioned above.

In the X-ray imaging system 10 described above, the arrangement pitch P of the pixels 40 of the X-ray image detector 30 is the period p ₁ ′ of the periodic intensity distribution of the self-image G1 (the grating pitch p _{1 of} the first absorption type grating 31). ) And is not sufficient for resolving the periodic intensity distribution of the self-image G1, so that the moire that can be resolved by the X-ray image detector 30 is formed using the second absorption grating 32, The phase contrast image is generated by analyzing the modulation of the periodic pattern of the image corresponding to the moire. On the other hand, in the X-ray imaging system 60 shown in FIG. 11, X-ray image detection capable of resolving the self-image G1 (the pixel arrangement pitch is sufficiently smaller than the period of the periodic intensity distribution of the self-image G1). A periodic intensity distribution of the self-image G1 is detected by an X-ray image detector, and a phase contrast image is obtained by analyzing a periodic pattern of an image corresponding to the periodic intensity distribution of the self-image G1. Generated.

The imaging unit 61 is provided with an X-ray image detector 62 and a first absorption type grating 31. The first absorption type grating 31 is configured by connecting a plurality of first grating modules 33, and the X-ray image detector 62 is also configured by connecting a plurality of detector modules 63. In the illustrated example, the first absorption type grating 31 includes four first grating modules 33 arranged in the x and y directions in a plane orthogonal to the optical axis A, and adjacent to each other. 33 are connected to each other. The X-ray image detector 62 includes five detector modules 63 arranged in the x direction and y direction in a plane orthogonal to the optical axis A, and the adjacent detector modules 63 connected to each other. Yes.

Each detector module 63 has an image receiving unit in which a plurality of pixels that detect X-rays and accumulate electric charges are two-dimensionally arranged in the xy direction. Note that a scanning circuit that controls the readout timing of the charges accumulated in each pixel, a signal processing circuit that converts and stores signals read sequentially from each pixel into image data, and image data stored in the console 13 A data transmission circuit that transmits to the arithmetic processing unit 22 via F25 may be provided for each detector module 63, or may be provided in the X-ray image detector 62 so as to control the plurality of detector modules 63. It may be.

The plurality of pixels 40 are arranged at an arrangement pitch capable of resolving the periodic intensity distribution of the self-image G1 formed on the X-ray image detector 62. Specifically, the pixel arrangement pitch P is set to a pitch of 1/2 or less, preferably 1/5 or less of the period p ₁ ′ of the periodic intensity distribution of the self-image G1, which is typically several μm. An image receiving unit in which a plurality of pixels are arranged at such a minute arrangement pitch is a CCD (Charge Coupled Device) in which a readout circuit for reading out the electric charge accumulated in each pixel is formed on a semiconductor substrate made of single crystal silicon or the like. A solid-state imaging device such as a sensor or a complementary metal oxide semiconductor (CMOS) sensor can be used as a base. In addition, as long as the above-described pixel arrangement pitch is satisfied, an image receiving unit that is configured based on a TFT panel can also be used.

In the above configuration, the self-image G1 of the first absorption grating 31 is formed on the X-ray image detector 62. When the subject H is disposed between the X-ray source 11 and the first absorption type grating 31, the periodic intensity distribution of the self-image G 1 formed on the X-ray image detector 62 depends on the subject H. Undergo modulation. The image acquired by detecting the self-image G1 by the X-ray image detector 62 includes a periodic pattern corresponding to the periodic intensity distribution of the self-image G1, and the subject H is analyzed by analyzing the periodic pattern. Phase contrast images can be generated.

In generating the phase contrast image, at least part of the boundary between the first grating modules 33 in the first absorption grating 31 and / or at least the boundary between the detector modules 63 in the X-ray image detector 62 is used. Corresponding to a part, the X-ray image acquired by the X-ray image detector 62 is divided into a plurality of partial X-ray images, and a periodic pattern is analyzed for each of these partial X-ray images.

FIG. 12 shows an example of the division of the X-ray image.

The example shown in FIG. 12 corresponds to all of the boundaries between the first grating modules 33 in the first absorption grating 31 and all of the boundaries between the detector modules 63 in the X-ray image detector 62. The line image is divided into partial X-ray images Img _1,1 , Img _2,1 ,..., Img _8,8 .

Specifically, in the first absorption type grating 31, the boundaries between the first grating modules 33 are boundary lines Lx ₁ , Lx ₂ , Lx ₃ extending in the x direction and boundary lines Ly ₁ , extending in the y direction. It is defined by Ly ₂ and Ly ₃ . In the X-ray image detector 62, the boundaries between the detector modules 63 are boundary lines Lx ₈ , Lx ₉ , Lx ₁₀ , Lx ₁₁ (see FIG. 11) extending in the x direction, and a boundary line Ly extending in the y direction. ₈ , Ly ₉ , Ly ₁₀ , Ly ₁₁ (see FIG. 11). Border _{Lx 1 ~ 3, Ly 1 ~} 3 of the X-ray image detector 62 is a projection onto the boundary _{Lx '1 ~ 3, Ly'} 1 ~ 3, and the X-ray image detector 62 itself boundary line Lx _{8 ~ 11, Ly 8} ~ 11 along with X-ray image the partial X-ray image _{Img 1,1, Img 2,1, ···,} are divided into _{Img 8, 8.} That is, Img _1,1 corresponds to an image obtained by detecting an X-ray image formed on the detector module 63 _1,1 by X-rays passing through the first grating module 33 _1,1 . Img _2,1 corresponds to an image acquired by detecting an X-ray image formed on the detector module 63 _2,1 by X-rays passing through the first grating module 33 _1,1 .

By analyzing the periodic pattern for each of the segmented partial X-ray images Img _1,1 , Img _2,1 ,..., Img _8,8 , partial phase shift distributions Φ _{m, n} (m = 1 to 8, N = 1 to 8) are obtained, and the phase shift distribution Φ of the subject H is obtained by combining the obtained partial phase shift distributions Φ _{m, n} .

According to this X-ray imaging system 60, an X-ray image detector capable of resolving the periodic intensity distribution of the self-image G1 (the arrangement pitch of the pixels 40 is sufficiently smaller than the period of the periodic intensity distribution of the self-image G1). 62 is used, the periodic intensity distribution of the self-image G1 is detected by the X-ray image detector 62, and the phase information is analyzed by analyzing the periodic pattern of the image corresponding to the periodic intensity distribution of the self-image G1. Since it is acquired and the arrangement pitch of the pixels 40 is minute, the spatial resolution is excellent. Further, since the second grating is not used, the accuracy of the obtained phase shift distribution Φ of the subject H can be improved.

An X-ray image detector with a small pixel arrangement pitch is limited to a relatively small size, and the S / N tends to decrease as the size increases. By configuring as one X-ray image detector 62, the size can be increased to ensure a field of view, and the S / N reduction can be suppressed, and the accuracy of the obtained phase shift distribution Φ of the subject H can be improved. Can do.

Further, for each partial X-ray image obtained by dividing the X-ray image corresponding to the boundary between the first grating modules 33 in the first absorption grating 31 and the boundary between the detector modules 63 in the X-ray image detector 62. By analyzing the periodic pattern and acquiring the phase shift distribution Φ of the subject H, the relative positional relationship of each of the first grating modules 33 with respect to the X-ray image detector 62 is further shifted, and further the first absorption grating The influence of the non-uniformity of the period and direction of the periodic pattern between the partial X-ray images caused by the deviation of the relative positional relationship of each of the detector modules 63 with respect to 31 on the phase shift distribution Φ of the subject H is eliminated. The accuracy of the obtained phase shift distribution Φ of the subject H can be increased.

FIG. 13 shows a modification of the X-ray imaging system 60 described above.

In the X-ray imaging system 70 shown in FIG. 13, the imaging unit 71 includes an X-ray image detector 72 and the first absorption grating 31, and the X-ray image detector 72 connects a plurality of detector modules 73. Configured. The arrangement pitch of the pixels in each detector module 73 is an arrangement of several μm, which is about the same as the period of the periodic intensity distribution of the self-image G1, so that moire is generated in relation to the periodic intensity distribution of the self-image G1. It is said to be a pitch. The arrangement pitch P of the pixels is preferably greater than _'sequences pitch necessary to resolve periodic intensity distribution of the self image G1 exhibiting a periodic intensity distribution of 1 / 2p _1' period p _1.

Since the pixel arrangement pitch P is a value determined by design and difficult to change, in order to adjust the magnitude relationship between the pixel arrangement pitch P and the period p ₁ ′ of the self-image G1, It is preferable to adjust the position of one absorption grating 31 by changing the period p ₁ ′ of the self-image G1. For the position adjustment of the first absorption type grating 31, for example, a mechanism similar to the above-described relative rotation mechanism 50, relative tilt mechanism 51, and relative movement mechanism 52 (see FIG. 6) can be used.

When the pixel arrangement pitch P satisfies the above condition, the period T in the x direction of moire generated in the image is expressed by the following equation (16).

In the above configuration, the self-image G1 of the first absorption grating 31 is formed on the X-ray image detector 72. When the subject H is arranged between the X-ray source 11 and the first absorption type grating 31, the periodic intensity distribution of the self-image G 1 formed on the X-ray image detector 72 depends on the subject H. Undergo modulation. As described above, the image acquired by detecting the self-image G1 by the X-ray image detector 72 has a moire (periodic pattern) according to the relationship between the period of the periodic intensity distribution of the self-image G1 and the arrangement pitch of the pixels. This moire is based on the periodic intensity distribution of the self-image G1. Therefore, a phase contrast image of the subject H can be generated by analyzing the moire.

In generating the phase contrast image, at least part of the boundary between the first grating modules 33 in the first absorption grating 31 and / or at least the boundary between the detector modules 73 in the X-ray image detector 72. Corresponding to a part, the X-ray image acquired by the X-ray image detector 72 is divided into a plurality of partial X-ray images, and moire analysis is performed for each of these partial X-ray images.

According to the present X-ray imaging system 70, since the period of moire is larger than the period of the periodic intensity distribution of the self-image G1, in terms of the arrangement pitch of pixels necessary for resolving this, Restrictions on the X-ray image detector can be relaxed.

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

A mammography apparatus 80 shown in FIG. 14 is an apparatus that captures an X-ray image (phase contrast image) of the breast B as a subject. The mammography apparatus 80 is disposed at one end of an arm member 81 that is pivotally connected to a base (not shown), and disposed at the other end of the arm member 81. An imaging table 83 and a compression plate 84 configured to be movable in the vertical direction with respect to the imaging table 83 are provided.

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

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

FIG. 15 shows a modification of the radiation imaging system of FIG.

15 is different from the mammography apparatus 80 described above in that the first absorption type grating 31 is disposed between the X-ray source 11 and the compression plate 84.

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

In the present mammography apparatus 90, the X-ray whose dose is almost halved is irradiated to the subject B due to the shielding by the first absorption type grating 31. Therefore, the exposure amount of the subject B is determined as described above. It can be reduced to about half that of the device 80. Note that, as in the mammography apparatus 90, the subject is disposed between the first absorption type grating 31 and the second absorption type grating 32, in other words, the first absorption type grating 31 is attached to the subject. The arrangement on the front side (X-ray source side) can be applied to any of the X-ray imaging systems described above.

FIG. 16 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

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

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

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

The lattice pitch p ₃ of the multi-slit 103 needs to be set so as to satisfy the following expression (17), where L ₃ is the distance from the multi-slit 103 to the first absorption-type lattice 31.

Expression (17) indicates that the projection image (self-image G1) of the X-rays emitted from the small focus light sources dispersedly formed by the multi-slit 103 by the first absorption-type grating 31 is the second absorption-type grating 32. This is a geometric condition for matching (overlapping) in position.

In addition, since the position of the multi-slit 103 is substantially the X-ray focal position, if the pitch p1 ′ of the self-image G1 at the position of the second absorption grating 32 is assumed, the grating pitch of the first absorption grating 31 p1 and the lattice pitch p2 of the second absorption type lattice 32 are determined so as to satisfy the relationship of the following equation (18).

As described above, in the present X-ray imaging system 100, the self-images G1 formed by the plurality of small focus light sources formed by the multi-slits 103 are superimposed, so that the phase contrast image is not reduced without reducing the X-ray intensity. Image quality can be improved. The multi slit 103 can be applied to any of the X-ray imaging systems described above.

FIG. 17 shows the configuration of the first and second gratings for another example of the radiation imaging system for explaining the embodiment of the present invention.

In the X-ray imaging system 10 described above, the first and second absorption gratings 31 and 32 are arranged such that the periodic arrangement direction of the X-ray shielding portions 31b and 32b is linear (that is, the grating surface is planar). However, instead of this, as shown in FIG. 19, it is also possible to use first and second absorption type gratings 110 and 111 having a substantially concave curved surface. In this case, it is preferable to use the X-ray image detector 112 having a cylindrical detection surface, and the detection surface of the X-ray image detector 112 has a straight line passing through the X-ray focal point 18b and extending in the y direction as a central axis. Cylindrical surface.

The first absorption type grating 110 is configured by connecting a plurality of first grating modules 33, and the first grating module 33 passes through the X-ray focal point 18b and extends in the extending direction (y Are arranged along a cylindrical surface whose center axis is an imaginary line extending in the direction). Similarly, the second absorption-type grating 111 is also configured by connecting a plurality of second grating modules 36, and the second grating module 36 passes through the X-ray focal point 18 b and has the X-ray shielding portion 38. They are arranged along a cylindrical surface having a virtual axis extending in the extending direction (y direction) as a central axis.

In this way, by configuring the first and second absorption type gratings 110 and 111 by connecting a plurality of grating modules, the grating surfaces can be easily formed into a substantially concave curved surface shape. Then, by making the grating surfaces of the first and second absorption gratings 110 and 111 substantially concave curved surfaces, the X-rays irradiated from the X-ray focal point 18 b since made incident substantially perpendicularly to the respective units, the upper limit of the limitation of the thickness h ₂ of the thickness h ₁ and the X-ray shielding portion 38 of the X-ray shielding portion 35 is reduced, the above expression (6) and (7) There is no need to consider.

The X-ray image detector 112 is also divided into a plurality of detector modules in the same manner as the first and second absorption type gratings 110 and 111, and these detector modules are arranged along the cylindrical surface. If arranged, the detection surface can be easily formed into a cylindrical surface.

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

As described above, the following radiographic imaging systems (1) to (9) are disclosed in this specification.

(1) An imaging unit that acquires a radiographic image including a periodic pattern modulated by a subject arranged in a radiation irradiation field, and generates a phase contrast image of the subject based on the periodic pattern included in the radiographic image An arithmetic processing unit, wherein the imaging unit forms one or more gratings that form a radiation image including a periodic intensity distribution that is a basis of the periodic pattern included in the radiation image by passing radiation; and A radiation image detector for detecting a radiation image, wherein at least one of the lattice and the elements of the radiation image detector is divided into a plurality of modules, and the arithmetic processing The unit is configured to correspond to at least a part of a boundary between modules in at least one of the elements divided into modules. A line image is divided into a plurality of partial radiation images, and for each partial radiation image, a Fourier transform is performed on the partial radiation image to obtain a spatial frequency spectrum of the partial radiation image, and the partial radiation A partial phase contrast that separates a spatial frequency region including a fundamental frequency component of a periodic pattern included in an image from the spatial frequency spectrum, and performs an inverse Fourier transform on the separated spatial frequency region to generate a partial phase contrast image. A radiographic imaging system that executes a combining process of combining a plurality of partial phase contrast images generated by the partial phase contrast image generating process to generate a phase contrast image of the subject. .
(2) The radiation imaging system according to (1), further including a first grating that forms a first radiation image including the periodic intensity distribution by passing radiation, wherein the radiation image detector includes A radiation imaging system for detecting a radiation image of 1.
(3) The radiation imaging system according to (2), wherein the radiation image detector is divided into a plurality of detector modules, and the arithmetic processing unit is at least detected by the radiation image detector. A radiation imaging system that divides the radiation image into the plurality of partial radiation images in correspondence with at least a part of a boundary between the instrument modules.
(4) In the radiation imaging system according to (2) or (3), the first grating is configured by being divided into a plurality of first grating modules, and the arithmetic processing unit includes at least the arithmetic processing unit. A radiation imaging system that divides the radiation image into the plurality of partial radiation images in correspondence with at least a part of a boundary between first grating modules in a first grating.
(5) The radiation imaging system according to any one of (2) to (4), wherein the radiation image detector includes an array of a plurality of pixels that detect radiation and accumulate electric charges, and The pixels are arranged at a pitch of ½ or less of the period of the periodic intensity distribution of the first radiation image, and the periodic pattern included in the radiation image is the periodic pattern of the first radiation image. Radiography system corresponding to intensity distribution.
(6) The radiation imaging system according to any one of (2) to (4), wherein the radiation image detector includes an array of a plurality of pixels that detect radiation and accumulate electric charges. The pixels are arranged in a pitch that forms moire in relation to the period of the periodic intensity distribution of the first radiation image, and the periodic pattern included in the radiation image has a radiation imaging system corresponding to the moire. .
(7) In the radiation imaging system according to (5) or (6), the radiation image detector includes a readout circuit that reads out the electric charges accumulated in the pixels, and the readout circuit is provided on a semiconductor substrate. The said radiography system currently formed.
(8) The radiation imaging system according to any one of (2) to (4), further including a second grating that masks the first radiation image to form a second radiation image including moire. And the radiographic image detector detects the second radiographic image instead of detecting the first radiographic image.
(9) In the radiographic system according to (8), the second grating is configured to be divided into a plurality of second grating modules, and the arithmetic processing unit includes at least the second grating. A radiation imaging system that divides the radiation image into the plurality of partial radiation images corresponding to at least a part of a boundary between the second grating modules in FIG.

According to the present invention, the interference between the period of the periodic pattern of the radiographic image and the pixel pitch of the radiographic image detector causes moiré to occur in the image acquired by the radiographic image detector, and is based on the modulation of moiré caused by the subject. To generate a phase contrast image. Therefore, it is not necessary to reduce the pixel pitch so that the periodic pattern of the radiation image can be detected, and S / N can be ensured to improve the accuracy of the phase information.

Although the present invention has been described in detail and with reference to specific embodiments, it will be apparent to those skilled in the art that various changes and modifications can be made without departing from the spirit and scope of the invention.
This application is based on a Japanese patent application (Japanese Patent Application No. 2011-130709) filed on June 10, 2011, the contents of which are incorporated herein by reference.

DESCRIPTION OF SYMBOLS 10 X-ray imaging system 11 X-ray source 12 Imaging part 13 Console 30 X-ray image detector 31 1st absorption grating 32 2nd absorption grating 40 Pixel

Claims (9)

1. An imaging unit for acquiring a radiographic image including a periodic pattern modulated by a subject arranged in a radiation irradiation field;
   An arithmetic processing unit for generating a phase contrast image of the subject based on the periodic pattern included in the radiation image;
   With
   The imaging unit includes one or more gratings that form a radiation image including a periodic intensity distribution that is a basis of the periodic pattern included in the radiation image, and a radiation image detection that detects the radiation image. And
   At least one element of the grating and the elements of the radiation image detector is configured by being divided into a plurality of modules.
   The arithmetic processing unit divides the radiographic image into a plurality of partial radiographic images corresponding to at least a part of a boundary between modules in at least one of the elements divided into modules, and each of the partial radiographic images A spatial frequency region including a fundamental frequency component of a periodic pattern included in the partial radiation image, and a transformation process for performing a Fourier transform on the partial radiation image to obtain a spatial frequency spectrum of the partial radiation image. Performing a partial phase contrast image generation process for generating a partial phase contrast image by performing an inverse Fourier transform on the separated spatial frequency domain to generate a partial phase contrast image, and performing the partial phase contrast image generation process Combining the plurality of partial phase contrast images generated by Executing a binding process to generate a phase contrast image of the body, a radiation imaging system.
2. The radiation imaging system according to claim 1,
   A first grating for forming a first radiation image including the periodic intensity distribution by passing radiation;
   The radiation image detector is a radiation imaging system that detects the first radiation image.
3. The radiographic system according to claim 2,
   The radiation image detector is configured by being divided into a plurality of detector modules,
   The radiation processing system in which the arithmetic processing unit classifies the radiation image into the plurality of partial radiation images in correspondence with at least a part of a boundary between detector modules in the radiation image detector.
4. The radiographic system according to claim 2 or 3,
   The first lattice is divided into a plurality of first lattice modules,
   The arithmetic processing unit divides the radiographic image into the plurality of partial radiographic images in correspondence with at least a part of a boundary between first grid modules in the first grid.
5. The radiographic system according to any one of claims 2 to 4,
   The radiation image detector has an array of a plurality of pixels that detect radiation and accumulate electric charge, and these pixels are equal to or less than ½ of the period of the periodic intensity distribution of the first radiation image. Arranged on the pitch,
   The periodic pattern included in the radiation image is a radiation imaging system corresponding to the periodic intensity distribution of the first radiation image.
6. The radiographic system according to any one of claims 2 to 4,
   The radiation image detector has an array of a plurality of pixels that detect radiation and accumulate electric charges, and these pixels are moire in relation to the period of the periodic intensity distribution of the first radiation image. Arranged in the pitch to be formed,
   The periodic pattern included in the radiation image is a radiation imaging system corresponding to the moire.
7. The radiographic system according to claim 5 or 6,
   The radiation image detector includes a readout circuit that reads out charges accumulated in the pixels, and the readout circuit is formed on a semiconductor substrate.
8. The radiographic system according to any one of claims 2 to 4,
   Further comprising a second grating for masking the first radiation image to form a second radiation image including moire,
   The radiographic image detector detects the second radiographic image instead of detecting the first radiographic image.
9. The radiation imaging system according to claim 8,
   The second grating is configured by being divided into a plurality of second grating modules,
   The radiographic system that divides the radiographic image into the plurality of partial radiographic images in correspondence with at least a part of a boundary between second grid modules in the second grid.

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