WO2012057045A1

WO2012057045A1 - X-ray imaging device, x-ray imaging system

Info

Publication number: WO2012057045A1
Application number: PCT/JP2011/074364
Authority: WO
Inventors: 村越　大
Original assignee: 富士フイルム株式会社
Priority date: 2010-10-25
Filing date: 2011-10-21
Publication date: 2012-05-03

Abstract

In phase imaging using radiation such as X-rays, a scattering elimination grid is used to eliminate or reduce scattered rays, and the scattering elimination grid prevents observed X-rays from affecting phase changes, and increases the quality of the phase contrast radiological images that are obtained. A radiation imaging system (10) is provided with an imaging unit (12), a radiation source (11), and an operation unit (13) that generates a phase contrast image of an object from a plurality of radiological images acquired by the imaging unit. The imaging unit (12) is provided with a first grid (31), a second grid (32) that forms a second periodic pattern image by partially masking a first periodic pattern image formed by radiation that has passed through the first grid, a radiation image detector (30) that detects the second periodic pattern image, and a scattering elimination grid (34) that eliminates scattered rays. The scattering elimination grid (34) is positioned between the second grid and the radiation image detector.

Description

Radiography apparatus, radiation imaging system

The present invention relates to a radiation imaging apparatus and 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 X-rays, and a transmission image of the subject is captured. In this case, each X-ray emitted from the X-ray source toward the X-ray image detector is caused by a difference in characteristics (atomic number, density, thickness) of the substance existing on the path to the X-ray image detector. After receiving a corresponding amount of attenuation (absorption), it enters each pixel of the X-ray image detector. As a result, the X-ray absorption image of the subject is detected and imaged by the X-ray image detector. As an X-ray image detector, a flat panel detector (FPD: Flat Panel Detector) using a semiconductor circuit is widely used in addition to a combination of an X-ray intensifying screen and a film and a stimulable phosphor. A part of the X-rays is scattered by the subject, and the scattered X-rays (scattered rays) reduce the density (contrast) of the image. For this reason, a scattering removal grating is used for the purpose of removing or reducing scattered radiation. The scattering removal grating is typically formed by slicing a laminate in which metal foils such as lead and copper that absorb X-rays and appropriate gap materials such as paper that transmits X-rays are alternately stacked. (For example, refer to Patent Document 1).

However, since the X-ray absorption ability is lower as a substance composed of an element having a smaller atomic number, there is a problem that a sufficient soft image contrast as an X-ray absorption image cannot be obtained in a soft body tissue or a soft material. 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, an X-ray phase for obtaining an image (hereinafter referred to as a phase contrast image) based on an X-ray phase change (angle change) by an object instead of an X-ray intensity change by an object. Imaging research is actively conducted. In general, it is known that when X-rays are incident on an object, the interaction is higher in phase than in X-ray intensity. For this reason, in the X-ray phase imaging using the phase difference, a high-contrast image can be obtained even for a weakly absorbing object having a low X-ray absorption capability. As a kind of such X-ray phase imaging, in recent years, an X-ray imaging system using an X-ray Talbot interferometer comprising two transmission diffraction gratings (phase grating and absorption grating) and an X-ray image detector has been proposed. It has been devised (for example, see Patent Document 2).

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

The X-ray Talbot interferometer detects moiré fringes generated by superimposing the first image of the first diffraction grating and the second diffraction grating, and obtains subject phase information by analyzing changes in the moiré fringes caused by the subject. To do. As a method for analyzing moire fringes, for example, a fringe scanning method is known. 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. The angle of X-rays refracted by the subject from a change in the signal value of each pixel obtained by the X-ray image detector, which is taken multiple times while being translated in the vertical direction at a scanning pitch obtained by equally dividing the lattice pitch. A distribution (differential image of phase shift) is obtained, and a phase contrast image of the subject can be obtained based on this angular distribution.

As described above, the X-ray phase imaging is to observe the phase change of the X-ray caused by the subject, and observing the phase change is the change of the optical path of the X-ray caused by the subject, that is, the refraction of the X-ray. Is equivalent to observing However, in the interaction between the X-ray and the subject, there is a physical phenomenon that changes the optical path of the X-ray in addition to refraction, and one example is scattering by the subject. And these phenomena become a factor which degrades the signal of each pixel brought about by the phase change based on the refraction of X-rays. In the X-ray imaging system described in Patent Document 2, an attempt is made to remove scattered radiation by using a second diffraction grating formed with a high aspect ratio.

Japanese National Table 2003-529087 Japan Special Table 2008-545981

A metal foil having a thickness of about several millimeters is generally used for the X-ray shielding portion of the normal scattering removal grating, and gold having a thickness of about 100 μm is generally used for the X-ray shielding portion of the second diffraction grating. Is used, and there is a possibility that the scattering removal performance is insufficient. On the other hand, the second diffraction grating typically needs to be configured with a pitch on the order of μm, so that a high aspect ratio grating with a sufficiently thick X-ray shielding portion that can remove or reduce scattered radiation is used. It is very difficult to manufacture.

Therefore, it is conceivable to use a scattering removal grating different from the second diffraction grating. However, X-ray phase imaging forms an image based on the phase change of the X-ray generated when passing through the subject, and conventional X-rays that form an image based on the intensity of the X-ray passing through the subject. It is very sensitive compared to absorption imaging. For this reason, slight irregularities such as processing traces on the surface of the scattering removal grating that did not become a shadow on the image in the X-ray absorption imaging can also become a shadow in the X-ray phase imaging. This shadow is difficult to distinguish and separate from the subject, and may cause an obstacle in image diagnosis.

The present invention has been made in view of the above-described circumstances. In phase imaging using radiation such as X-rays, the scattered radiation is removed or reduced using a scattering removal grating, and the scattering removal grating is observed. It is intended to prevent the influence of the phase change on the image and to improve the quality of the obtained radiation phase contrast image.

A radiation source, a first grating that forms a first periodic pattern image by passing radiation emitted from the radiation source, and disposed on a side opposite to the radiation source with respect to the first grating, A second grating that forms a second periodic pattern image generated by partially shielding the first periodic pattern image, a radiation image detector that detects the second periodic pattern image, and the second A radiation imaging apparatus comprising: a scatter removal grating that is disposed between a grating and the radiation image detector and removes scattered radiation.

According to the present invention, the scattered radiation can be sufficiently removed or reduced by using the scattering removal grating. Therefore, the scattering removal grating is disposed between the second grating and the radiation image detector, and the scattering removal grating can be prevented from affecting the observed X-ray phase change. . Thereby, the image quality of the obtained radiation phase contrast image can be improved.

It is a schematic diagram which shows the structure of an example of the radiography system for describing embodiment of this invention. It is a control block diagram of the radiography system of FIG. It is a schematic diagram which shows the structure of the radiographic image detector of the radiography system of FIG. It is a perspective view of the imaging part of the radiography system of FIG. It is a side view of the imaging part of the radiography system of FIG. It is a schematic diagram which shows the mechanism for changing the period of the moire fringe by superimposition of the 1st and 2nd grating | lattice. It is a schematic diagram for demonstrating the refraction | bending of the radiation by a to-be-photographed object. It is a schematic diagram for demonstrating the fringe scanning method. It is a graph which shows the signal of the pixel of the radiographic image detector accompanying a fringe scanning. It is a schematic diagram for demonstrating the other example of the production | generation method of the phase contrast image in the radiography system of FIG. It is a schematic diagram for demonstrating the other example of the production | generation method of the phase contrast image in 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 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 block diagram which shows the structure of the calculating part which produces | generates a radiographic image regarding the other example of the radiography system for describing embodiment of this invention. It is a graph which shows the signal of the pixel of the radiographic image detector for demonstrating the process in the calculating part of the radiography system of FIG.

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

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

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

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

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

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

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

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

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

The imaging unit 12 includes a flat panel detector (FPD) 30 made of a semiconductor circuit, a first absorption type grating 31 and a second absorption type for detecting phase change (angle change) of X-rays by the subject H and performing phase imaging. The absorption type grating 32 and the scatter removal grating 34 for removing scattered radiation are provided.

The FPD 30 is arranged so that the detection surface is orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11. As will be described in detail later, the first and

second absorption gratings

31 and 32 and the scattering removal grating 34 are disposed between the FPD 30 and the X-ray source 11.

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

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

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

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

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

The readout circuit 43 includes an integration amplifier circuit, an A / D converter, a correction circuit, and an image memory (all not shown). The integrating amplifier circuit integrates the charges output from each pixel 40 via the signal line 46, converts them into a voltage signal (image signal), and inputs it to the A / D converter. The A / D converter converts the input image signal into digital image data and inputs the digital image data to the correction circuit. The correction circuit performs offset correction, gain correction, and linearity correction on the image data, and stores the corrected image data in the image memory. As correction processing by the correction circuit, correction of X-ray exposure amount and exposure distribution (so-called shading) and pattern noise depending on FPD 30 control conditions (drive frequency and readout period) (for example, leak signal of TFT switch) May be included.

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

The first absorption-type grating 31 includes a substrate 31a and a plurality of X-ray shielding portions 31b arranged on the substrate 31a. Similarly, the second absorption type grating 32 includes a substrate 32a and a plurality of X-ray shielding portions 32b arranged on the substrate 32a. The

substrates

31a and 32a are both made of an X-ray transparent member such as glass that transmits X-rays.

Each of the

X-ray shielding portions

31b and 32b is in one direction in a plane orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11 (in the illustrated example, the y direction orthogonal to the x direction and the z direction). It is comprised by the linear member extended | stretched. As a material of each

X-ray shielding part

31b, 32b, a material excellent in X-ray absorption is preferable, and for example, a heavy metal such as gold or platinum is preferable. These

X-ray shielding portions

31b and 32b can be formed by a metal plating method or a vapor deposition method.

The X-ray shielding part 31b has a predetermined interval d ₁ with a constant period (lattice pitch) p ₁ in a direction (x direction) orthogonal to the one direction in a plane orthogonal to the optical axis A of the X-ray. It is arranged in a space. Similarly, X-ray shielding portion 32b, in the plane orthogonal to the optical axis A of the X-ray, in the direction predetermined period (x-direction) (grating pitch) p ₂ perpendicular to the one direction, from each other a predetermined distance _They are arranged with d _{2 in} between. Since the first and

second absorption gratings

31 and 32 do not give a phase difference to incident X-rays but give an intensity difference, they are also called amplitude gratings. Note that the slit portions (regions having the distances d ₁ and d ₂ ) may not be voids, and the voids may be filled with an X-ray low-absorbing material such as a polymer or a light metal.

The first and

second absorption gratings

31 and 32 are configured to geometrically project the X-rays that have passed through the slit portion regardless of the presence or absence of the Talbot interference effect. Specifically, by setting the distances d ₁ and d ₂ to a value sufficiently larger than the peak wavelength of X-rays emitted from the X-ray source 11, most of the X-rays included in the irradiated X-rays are slit at the slit portion. It is configured to pass through without being diffracted while maintaining straightness. For example, when tungsten is used as the rotary anode 18a described above and the tube voltage is 50 kV, the peak wavelength of the X-ray is about 0.4 mm. In this case, if the distances d ₁ and d ₂ are about 1 to 10 μm, most of the X-rays are geometrically projected without being diffracted at the slit portion.

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

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

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

Expression (2) is an expression that represents the Talbot interference distance when the X-ray irradiated from the X-ray source 11 is a cone beam. “Atsushi Momose, et al., Japan Journal of Applied Physics, Vol. 47, No. 10, October 2008, page 8077 ”.

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

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

The

X-ray shielding portions

31b and 32b 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. Therefore, in order to enhance the shielding of the X-rays, the X-ray shielding portion 31b, the respective thicknesses _h 1, _{h 2} of 32b, it is preferable to increase the thickness much as possible. For example, when the tube voltage of the X-ray tube 18 is 50 kV, it is preferable to shield 90% or more of the irradiated X-rays. In this case, the thicknesses h ₁ and h ₂ are 30 μm or more in terms of gold (Au). It is preferable that

However, when the X-rays irradiated from the X-ray source 11 are cone beams, if the thicknesses h ₁ and h ₂ of the

X-ray shielding portions

31b and 32b are too thick, the X-rays incident obliquely enter the slit portion. There is a problem that it becomes difficult to pass, so-called vignetting occurs, and the effective visual field in the direction (x direction) perpendicular to the extending direction (strand direction) of the

X-ray shielding portions

31b and 32b becomes narrow. Therefore, in view of the field of view secured to define the upper limit of the thickness _h 1, _{h 2.} In order to secure the effective field length V in the x direction on the detection surface of the FPD 30, assuming that the distance from the X-ray focal point 18 b to the detection surface of the FPD 30 is L, the thicknesses h ₁ and h ₂ are shown in FIG. From the scientific relationship, it is necessary to set so as to satisfy the following expressions (6) and (7).

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

The scattering removal grating 34 includes a plurality of X-ray shielding portions 34a and a plurality of X-ray transmission portions 34b.

The X-ray shielding part 34a is composed of a strip-shaped member 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. As a material of the X-ray shielding part 34a, a material excellent in X-ray absorption is preferable. For example, a metal foil such as lead, copper, or tungsten is used. The X-ray shielding portions 34a are arranged at intervals in a direction (x direction) orthogonal to the one direction in a plane orthogonal to the optical axis A of X-rays. The X-ray transmission part 34b is provided so as to fill a space between adjacent X-ray shielding parts 34a. As a material of the X-ray transmission part 34b, an X-ray low absorption material is preferable, for example, a polymer, a light metal, or the like is used.

The scattering removal grating 34 is arranged between the second absorption grating 32 downstream of the subject H and the FPD 30, and in the traveling direction among X-rays scattered by the subject H (hereinafter referred to as scattered rays). Scattered rays having a component in the arrangement direction (x direction) of the X-ray shielding part 34a are removed or reduced. Further, the scattering removal grating 34 disposed between the second absorption type grating 32 and the FPD 30 (downstream of the second absorption type grating 32) is a first absorption type grating 31 or a second absorption type grating 32. Scattered rays generated in can also be removed or reduced.

The scatter removal grating 34 is parallel to the optical axis A of the X-rays irradiated from the X-ray source 11 in the cross section along the direction (x direction) in which the X-ray shields 34 a are arranged. Although a so-called parallel grid may be used, it is preferable to use a so-called focusing grid in which the extension of each of the X-ray shielding portions 34a is focused on the radiation focus as in the illustrated example. According to this, it is possible to make it difficult to generate so-called vignetting for X-rays that travel substantially along the radiation direction from the X-ray source 11 without being scattered by the subject H. Note that the phase change (angle change) of the X-rays refracted by the subject H is typically several μrad, and the refracted X-rays travel substantially along the radiation direction from the X-ray source 11.

In the imaging unit 12 configured as described above, an intensity-modulated image is formed by superimposing the G1 image of the first absorption-type grating 31 and the second absorption-type grating 32 and is captured by the FPD 30. . Here, the scattered light due to the subject or the like is removed or reduced by the scattering removal grating 34, thereby preventing the contrast of the intensity-modulated image to be picked up from being lowered. The pattern period p ₁ ′ of the G1 image at the position of the second absorption grating 32 and the substantial grating pitch p ₂ ′ (substantial pitch after production) of the second absorption grating 32 are manufacturing errors. Some differences occur due to or placement errors. Among these, the arrangement error means that the substantial pitch in the x direction changes due to the relative inclination and rotation of the first and

second absorption gratings

31 and 32 and the distance between the two changes. I mean.

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

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

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

Since the arrangement pitch P of the pixels 40 of the FPD 30 is a value determined by design (generally about 100 μm) and is difficult to change, the magnitude relationship between the arrangement pitch P and the moire period T is adjusted. Adjusts the positions of the first and

second absorption gratings

31 and 32 and changes the moire period T by changing at least one of the pattern period p ₁ ′ and the grating pitch p ₂ ′ of the G1 image. It is preferable to do.

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

The moire period T can be changed by relatively rotating one of the first and

second absorption gratings

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

As another example, the change of the moire period T is such that either one of the first and second

absorption type gratings

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

As another example, the moire period T can be changed by relatively moving one of the first and

second absorption gratings

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

In the X-ray imaging system 10, imaging unit 12 is not the Talbot interferometer as described above, since the distance L ₂ can be freely set, moire by changing the distance L ₂ as relative movement mechanism 52 A mechanism for changing the period T can be suitably employed. The change mechanism (relative rotation mechanism 50, relative tilt mechanism 51, and relative movement mechanism 52) of the first and

second absorption gratings

31 and 32 for changing the moiré period T is constituted by an actuator such as a piezoelectric element. Is possible.

When the subject H is disposed between the X-ray source 11 and the first absorption type grating 31, the moire fringes detected by the FPD 30 are modulated by the subject H. This modulation amount is proportional to the angle of the X-ray deflected by the refraction effect by the subject H. Therefore, the phase contrast image of the subject H can be generated by analyzing the moire fringes detected by the FPD 30.

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

[Analysis method 1]
FIG. 7 shows one X-ray refracted according to the phase shift distribution Φ (x) of the subject H in the x direction. The illustration of the scattering removal grating is omitted.

Reference numeral 55 indicates an X-ray path that travels straight when the subject H is not present. The X-ray that travels along the path 55 passes through the first and

second absorption gratings

31 and 32 and enters the FPD 30. To do. Reference numeral 56 indicates an X-ray path refracted and deflected by the subject H when the subject H exists. X-rays traveling along this path 56 are shielded by the second absorption type grating 32 after passing through the first absorption type grating 31.

The phase shift distribution Φ (x) of the subject H is expressed by the following equation (11), where n (x, z) is the refractive index distribution of the subject H, and z is the direction in which the X-ray travels.

The G1 image projected from the first absorptive grating 31 to the position of the second absorptive grating 32 is displaced in the x direction by an amount corresponding to the refraction angle φ due to refraction of X-rays at the subject H. become. This amount of displacement Δx is approximately expressed by the following equation (12) based on the small X-ray refraction angle φ.

Here, the refraction angle φ is expressed by Expression (13) using the X-ray wavelength λ and the phase shift distribution Φ (x) of the subject H.

Thus, the displacement amount Δx of the G1 image due to the refraction of X-rays at the subject H is related to the phase shift distribution Φ (x) of the subject H. The amount of displacement Δx is expressed by the following equation with the phase shift amount ψ of the signal output from each pixel 40 of the FPD 30 (the phase shift amount of the signal of each pixel 40 with and without the subject H): It is related as shown in (14).

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

In the fringe scanning method, imaging is performed while one of the first and second

absorption type gratings

31 and 32 is translated in a stepwise manner relative to the other in the x direction (that is, the phase of both grating periods is changed). Shoot while changing). In the X-ray imaging system 10, the second absorption type grating 32 is moved by the scanning mechanism 33 described above, but the first absorption type grating 31 may be moved. As the second absorption type grating 32 moves, the moire fringes move, and the translation distance (the amount of movement in the x direction) is one period of the grating period of the second absorption type grating 32 (grating pitch p ₂ ). (Ie, when the phase change reaches 2π), the moire fringes return to their original positions. With such a change in moire fringes, a fringe image is photographed with the FPD 30 while moving the second absorption grating 32 by an integer of the grating pitch p ₂ , and each pixel 40 is captured from the plural fringe images photographed. The signal is acquired and processed by the processing unit 22 to obtain the phase shift amount ψ of the signal of each pixel 40.

FIG. 8 schematically shows how the second absorption grating 32 is moved by the scanning pitch (p ₂ / M) obtained by dividing the grating pitch p ₂ into M (an integer of 2 or more).

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

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

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

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

Next, using the relational expression of the following expression (16), the refraction angle φ (x) is expressed as the following expression (17).

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

FIG. 9 shows a signal of one pixel of the radiation image detector that changes with the fringe scanning.

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

Since the refraction angle φ (x) is a value corresponding to the differential phase value as shown in the above equation (13), the phase shift is obtained by integrating the refraction angle φ (x) along the x-axis. 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).

[Analysis method 2]
FIG. 10 shows another example of a method for generating a phase contrast image in the X-ray imaging apparatus 1.

In the method described below, a plurality of images equivalent to the plurality of images acquired by the above-described fringe scanning method are acquired by one shooting.

By arranging the first and second

absorption type gratings

31 and 32 to rotate relative to each other by an angle θ about the optical axis C, the G1 image and the second absorption type grating 32 are relatively set to an angle θ. Rotate. The overlap between the periodic intensity distribution of the G1 image at the position of the image receiving surface of the FPD 30 and the projection of the second absorption type grating 32 (projection of the periodic arrangement of the X-ray shielding part 32b) is the same as that of the first absorption type grating 31. It changes periodically in a direction (y direction) orthogonal to the lattice pitch direction (x direction).

The rotation angle θ is a distance corresponding to n cycles of the above change (where n is an integer other than 0 and a multiple of M) where Py is the arrangement pitch of the pixels 40 in the y direction and M is the number of image data to be acquired. D is set to be D = Py × M. Thereby, with the M pixels 40 adjacent in the y direction as a unit, the periodic intensity distribution of the G1 image and the X-ray shielding part 32b of the second absorption type grating 32 between the M pixels for each unit. A state is formed in which phases with the periodic array are different from each other. The illustrated example shows a case where M = 5 and n = 1.

Then, M images are acquired by forming a plurality of pixel rows every M rows as a set and forming an image for each set based on the charges read from each pixel 40 of the pixel row group of the set. The When the rotation angle θ is set with M = 5 and n = 1 as in the example shown in the drawing, the first pixel row group of (5 × k−4) rows (k = 1, 2,...) Are acquired, the second image is acquired from the pixel row group of (5 × k−3) rows, the third image is acquired from the pixel row group of (5 × k−2) rows, and (5 A fourth image is acquired from the pixel row group of (x-1) rows, and a fifth image is acquired from the pixel row group of (5 × k) rows.

The method for generating the phase contrast image based on the M images acquired as described above is the same as the above-described fringe scanning method.

According to the above-described method for generating a phase contrast image, a plurality of images required for analysis of moire fringes can be acquired by one shooting, and the first absorption type grating 31 or the second second during a plurality of shootings can be acquired. The movement of the absorption type grating 32 and the scanning mechanism 33 that requires high accuracy are not required. Therefore, it is possible to improve the shooting workflow and simplify the apparatus. In addition, it is possible to eliminate the deterioration in image quality caused by the movement of the subject between each photographing.

[Analysis method 3]
FIG. 11 shows another example of a method for generating a phase contrast image in the X-ray imaging apparatus 1.

In the method described below, the moire fringes are analyzed using Fourier transform and inverse Fourier transform instead of the above-described fringe scanning method, and a phase contrast image is generated.

The moire fringes formed by superimposing the G1 image and the second absorption type grating 32 can be expressed by the following equation (18), and equation (18) can be rewritten into the following equation (19).

In Expression (18), a (x, y) represents the background, b (x, y) represents the amplitude of the spatial frequency component corresponding to the fundamental period of the moire fringes, and (f _0x, f _0y ) represents the moire. Represents the basic period of the stripes. In the equation (19), c (x, y) is represented by the following equation (20).

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, Expression (17) becomes the following Expression (21) by Fourier transform.

In the formula _{_{_{(21), 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 one-dimensional gratings such as the first and

second absorption gratings

31 and 32 are used, the spatial frequency spectrum of the image has at least A (f _x , f _y ) as shown in FIG. a peak derived _{from, 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 this . 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 fundamental period of the moire fringes 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. The method for generating the phase contrast image from the refraction angle φ (x, y) is the same as the above-described fringe scanning method.

According to the above-described method for generating a phase contrast image, a phase contrast image can be generated from one image, and thus only one imaging is required. The movement of the second absorption type grating 32 and the scanning mechanism 33 that requires high accuracy are not required. Therefore, it is possible to improve the shooting workflow and simplify the apparatus. In addition, it is possible to eliminate the deterioration in image quality caused by the movement of the subject between each photographing.

The above calculation is performed by the calculation processing unit 22, and the calculation processing unit 22 stores the phase contrast image in the storage unit 23.

In the X-ray phase imaging using the first and

second absorption gratings

31 and 32, X-ray refraction is observed in an object including the subject H upstream from the second absorption grating 32. Therefore, scattering removal is performed upstream of the second absorption type grating 32 (for example, between the subject H and the first absorption type grating 31 or between the first absorption type grating 31 and the second absorption type grating 32). If a grating is present, X-ray refraction at the scattering removal grating is observed without distinction from X-ray refraction at the subject H. X-ray refraction occurs remarkably in regions where the optical distances through which X-rays pass are different, particularly at the edges of the transmitting objects. It will be detected.

In the present X-ray imaging system 10, the scattering removal grating 34 is disposed between the second absorption grating 32 and the FPD 30, that is, downstream of the second absorption grating 32. Therefore, refraction at the scatter removal grating 34 is not observed, and it is possible to prevent unevenness on the surface of the scatter removal grating 34 from appearing as a shadow on the X-ray phase contrast image.

Also, in obtaining the phase shift distribution Φ (x), the refraction angle φ (x) is integrated along the x direction. The phase shift distribution Φ (x) is based on the X-ray refraction component in the x direction that is the arrangement direction of the

X-ray shielding portions

31b and 32b in the first and

second absorption gratings

31 and 32. Therefore, the scattered radiation including the x-direction component in the traveling direction, particularly the scattered radiation traveling in the x-direction, has a relatively large influence on the phase shift distribution Φ (x).

From the viewpoint of removing scattered radiation, the scattering removal grating 34 is arranged so that the arrangement direction of the X-ray shielding portions 34a is in the x direction (see FIG. 4), that is, the arrangement direction of the X-ray shielding portions 34a is the second. It is preferable to dispose the scattering removal grating 34 so as to be parallel (0 °) to the arrangement direction of the X-ray shielding portions 32 b in the absorption grating 32. Thereby, the scattering removal grating 34 can effectively remove or reduce the scattered radiation including the component in the x direction in the traveling direction, particularly the scattered radiation traveling in the x direction, and obtain the phase shift distribution Φ (x). Therefore, a phase shift distribution Φ (x) can be obtained with high accuracy by extracting the X-ray refraction component in the x direction.

Note that the second absorption type grating 32 also has a considerable function of removing scattered radiation. Therefore, the scattering direction of the scattering removal grating 34 is such that the arrangement direction of the X-ray shielding portions 34a is X in the second absorption type grating 32. You may arrange | position so that the arrangement | sequence direction of the line shielding part 32b may be crossed. However, if the scattering removal grating 34 is arranged so that the arrangement direction of the X-ray shielding part 34a is orthogonal to the arrangement direction of the X-ray shielding part 32b in the second absorption type grating 32, the phase shift distribution Φ (x) The removal of the scattered radiation traveling in the x-direction having the greatest influence depends only on the second absorption type grating 32 that is inferior to the scattering removal grating 34 in terms of the scattering removal ability. Therefore, the angle formed by the arrangement direction of the plurality of X-ray shielding portions 34a in the scattering removal grating 34 and the arrangement direction (x direction) of the X-ray shielding portions 32b in the second absorption type grating 32 is 0 ° or more and 90 °. Is less than 0, preferably 0 °.

The scattering removal grating 34 is arranged so that the angle formed by the arrangement direction of the X-ray shielding part 34a and the arrangement direction (x direction) of the X-ray shielding part 32b in the second absorption grating 32 is not less than 0 ° and less than 90 °. When arranged, depending on the arrangement pitch of the X-ray shields 34a of the scatter removal grating 34, a moire having periodicity in the x direction is generated in relation to the arrangement pitch P of the pixels 40 in the spatial frequency response in the x direction. There is a case. In that case, since the G1 image is also a moire fringe having periodicity in the x direction, the above moire caused by the scattering removal grating 34 and the FPD 30 is removed from the image by appropriate image processing such as applying a frequency filter. Need to be done. Accordingly, the arrangement pitch of the X-ray shielding portions 34a of the scatter removal grating 34 is a pitch that does not cause a moire of a frequency that causes an image problem in relation to the arrangement pitch P of the pixels 40 in the x direction in the spatial frequency response. It is preferable to set to.

From the viewpoint of preventing the occurrence of moire having periodicity in the x direction due to the scattering removal grating 34 and the FPD 30, the scattering removal grating 34 is arranged so that the arrangement direction of the X-ray shielding portions 34a is in the y direction. That is, the scattering removal grating 34 is arranged so that the arrangement direction of the X-ray shielding part 34 a is orthogonal (90 °) to the arrangement direction (x direction) of the X-ray shielding part 32 b in the second absorption type grating 32. It may be.

Further, in a conventional X-ray imaging system that obtains an image (X-ray absorption contrast image) based on X-ray absorption, scattering removal is performed in order to prevent the X-ray shielding portion of the scattering removal grating from appearing in the image. The grating is moved during X-ray irradiation. In the present X-ray imaging system 10 that obtains an X-ray phase contrast image, the scatter removal grating 34 may be similarly moved during X-ray irradiation. Preferably, however, the scatter removal grating 34 is stationary at least during X-ray irradiation. Is done. This is because vibration is generated by moving the scattering removal grating 34, and relative vibration between the first and

second absorption gratings

31 and 32 can occur due to the vibration. The refraction angle of X-rays generated by passing through the subject H is as small as several μrad, and the displacement amount of the G1 image generated along with this and the phase shift amount of the signal of each pixel 40 related to this displacement amount are also slight. It is. Therefore, the relative positional deviation between the first and

second absorption gratings

31 and 32 affects the phase detection capability. Therefore, in order to prevent the relative displacement between the first and second

absorption type gratings

31 and 32, it is preferable that the scattering removal grating 34 is stationary during the X-ray irradiation.

The above-described phase contrast image generation processing is automatically performed after the imaging instruction is given by the operator from the input device 21, and the units are linked and operated automatically under the control of the control device 20. A phase contrast image is displayed on the monitor 24.

As described above, according to the present X-ray imaging system 10, the scattered radiation can be sufficiently removed or reduced by using the scattering removal grating 34. In this case, the scattering removal grating 34 is disposed between the second grating 32 and the FPD 30, and the scattering removal grating 34 can be prevented from affecting the observed X-ray phase change. . Thereby, the image quality of the obtained X-ray phase contrast image 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 this X-ray imaging system, almost all wavelength components of irradiated X-rays contribute to the projection image (G1 image) from the first absorption type grating 31 and the contrast of moire fringes is improved. Contrast image detection sensitivity can be improved.

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

Further, although the X-ray imaging system 10 has been described as one that stores or displays an image of the phase shift distribution Φ as a phase contrast image, as described above, the phase shift distribution Φ is a phase determined from the refraction angle φ. The differential amount of the shift distribution Φ is integrated, and 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 having the refraction angle φ as an image and an image having the differential amount of the phase shift Φ are also included in the phase contrast image.

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

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

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

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

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

A mammography apparatus 90 shown in FIG. 13 is different from the mammography apparatus 80 described above in that the first absorption grating 31 is disposed between the X-ray source 11 and the compression plate 84. The first absorption type lattice 31 is accommodated in a lattice accommodation portion 91 connected to the arm member 81. The imaging unit 92 includes the FPD 30, the second absorption type grating 32, the scanning mechanism 33, and the scattering removal grating 34. In the illustrated example, the scattering removal grating 34 is disposed between the second absorption type grating 32 and the FPD 30.

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

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

Preferably, in the case of an imaging technique in which the influence of scattered radiation is strong, the image is taken using the scattering removal grating 34, and in the case where the influence of scattered radiation is weak, the image is taken without using the scattered radiation removal grid from the viewpoint of reducing exposure. The Examples of imaging techniques that are strongly influenced by scattered radiation include those that take images of thick parts such as the waist (FIG. 1) and the side of the body, or those that depict pale contrasts such as the lungs and breasts (FIG. 12). The On the other hand, as a case where the influence of scattered radiation is weak, for example, a case where a thin part such as a finger or a toe is photographed is exemplified. Therefore, it is preferable that the scattering removal grating 34 can be retracted from the X-ray irradiation field. For example, in the housing of the imaging unit 12 that houses the scatter removal grating 34, the scatter removal is movably movable in one direction (for example, the y direction) in the plane orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11. In the case of an imaging technique that supports the grating 34 and advances or retracts the scatter removal grating 34 in one direction using an appropriate driving mechanism, and the influence of the scattered radiation is weak, the scatter removal grating 34 is converted into an X-ray by the above driving mechanism. Evacuate from field. Alternatively, the scattering removal grating 34 may be removed from the X-ray irradiation field by being removed from the housing of the imaging unit 12.

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

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

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

The multi-slit 103 is an absorption type grating (third absorption type grating) having a configuration similar to that of the first and second

absorption type gratings

31 and 32 provided in the imaging unit 12, and is in one direction (y direction). The extended X-ray shielding portions are periodically arranged in the same direction (x direction) as the

X-ray shielding portions

31b and 32b of the first and

second absorption gratings

31 and 32. The multi-slit 103 is intended to form a large number of small-focus light sources (dispersed light sources) arranged at a predetermined pitch in the x direction by partially shielding the radiation emitted from the X-ray focal point 18b. .

The grating pitch p ₃ of the multi slit 103 needs to be set so as to satisfy the following expression (22), where L ₃ is the distance from the multi slit 103 to the first absorption type grating 31.

The above formula (22) indicates that the projection image (G1 image) of the X-rays emitted from the respective point light sources dispersedly formed by the multi slit 103 by the first absorption type grating 31 is the position of the second absorption type grating 32. This is a geometric condition for matching (overlapping).

Further, since the position of the substantially multi-slit 103 is X-ray focal position, the grating pitch p ₂ of the second absorption-type grating 32 is determined so as to satisfy the following relation (23).

As described above, in the present X-ray imaging system 100, the G1 images based on the plurality of point light sources formed by the multi slit 103 are superimposed, thereby improving the image quality of the phase contrast image without reducing the X-ray intensity. Can be made. The multi slit 103 described above can be applied to any of the X-ray imaging systems described above.

In the above description, the first and

second absorption gratings

31 and 32 and the multi slit 103 are described as one-dimensional gratings. However, the first and

second absorption gratings

31 and 32 are described. In addition, the multi-slit 103 can be a two-dimensional lattice, and the present invention is useful also in that case.

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

According to each X-ray imaging system described above, a high-contrast image (phase contrast image) of an X-ray weakly absorbing object that has been difficult to draw can be obtained. In addition, an absorption image is referred to corresponding to the phase contrast image. What you can do will help you interpret. For example, it is effective to supplement the portion that could not be represented by the absorption image with the information of the phase contrast image by superimposing the absorption image and the phase contrast image by appropriate processing such as weighting, gradation, and frequency processing. However, capturing an absorption image separately from the phase contrast image makes it difficult to superimpose images due to the shift in the shooting position between the phase contrast image capture and the absorption image capture. Increasing the burden on the subject. In recent years, small-angle scattered images have attracted attention in addition to phase contrast images and absorption images. The small-angle scattered image can express tissue properties resulting from the fine structure inside the subject tissue, and is expected as a new expression method for image diagnosis in the fields of cancer and cardiovascular diseases.

Therefore, this X-ray imaging system uses an arithmetic processing unit 190 that can generate an absorption image and a small-angle scattered image from a plurality of images acquired for a phase contrast image. Since other configurations are the same as those of the X-ray imaging system 10 described above, description thereof will be omitted. The arithmetic processing unit 190 includes a phase contrast image generation unit 191, an absorption image generation unit 192, and a small angle scattered image generation unit 193. These all perform arithmetic processing based on image data obtained at M scanning positions of k = 0, 1, 2,..., M−1. Among these, the phase contrast image generation unit 191 generates a phase contrast image according to the above-described procedure.

The absorption image generation unit 192 generates an absorption image by averaging the pixel data Ik (x, y) obtained for each pixel with respect to k and calculating an average value as shown in FIG. . The average value may be calculated by simply averaging the pixel data Ik (x, y) with respect to k. However, when M is small, the error increases, so the pixel data Ik (x, y After fitting y) with a sine wave, an average value of the fitted sine wave may be obtained. The generation of the absorption image is not limited to the average value, and an addition value obtained by adding the pixel data Ik (x, y) with respect to k can be used as long as the amount corresponds to the average value.

The small angle scattered image generation unit 193 generates a small angle scattered image by calculating and imaging the amplitude value of the pixel data Ik (x, y) obtained for each pixel. The amplitude value may be calculated by obtaining a difference between the maximum value and the minimum value of the pixel data Ik (x, y). However, when M is small, the error increases, and therefore the pixel data Ik. After fitting (x, y) with a sine wave, the amplitude value of the fitted sine wave may be obtained. In addition, the generation of the small-angle scattered image is not limited to the amplitude value, and a dispersion value, a standard deviation, or the like can be used as an amount corresponding to the variation centered on the average value.

According to the present X-ray imaging system, an absorption image and a small angle scattered image are generated from a plurality of images acquired for the phase contrast image of the subject. There is no deviation, and it is possible to superimpose the phase contrast image with the absorption image and the small-angle scattered image, and the burden on the subject is reduced as compared with the case of separately shooting for the absorption image and the small-angle scattered image. be able to.

As described above, the present specification includes a radiation source, a first grating that forms a first periodic pattern image by passing the radiation emitted from the radiation source, and the first grating. A second grating that is disposed on the opposite side of the radiation source and forms a second periodic pattern image generated by partially shielding the first periodic pattern image, and the second periodic pattern image There is disclosed a radiographic apparatus including a radiological image detector for detecting scatter, and a scatter removal grating disposed between the second grating and the radiographic image detector for removing scattered radiation.

Further, the radiation imaging apparatus disclosed in the present specification moves one of the first grating and the second grating, and moves the second grating to the plurality of relative positions with respect to the radiation image. And a scanning mechanism.

Also, in the radiation imaging apparatus disclosed in the present specification, the scatter removal grating is stationary at least while the radiation image is detected by the radiation image detector.

Further, in the radiation imaging apparatus disclosed in this specification, the scatter removal grating can be retracted from the radiation irradiation field.

Further, in the radiation imaging apparatus disclosed in the present specification, the scatter removal grating has a plurality of radiation shielding units arranged in parallel at intervals, and the arrangement direction of the plurality of radiation shielding units And the pitch direction of the first grating are orthogonal to each other.

Also, in the radiation imaging apparatus disclosed in this specification, the first grating is a phase-type grating.

Also, in the radiation imaging apparatus disclosed in this specification, the first grating is an absorption grating.

The radiation imaging apparatus disclosed in the present specification further includes a third grating that irradiates the first grating by selectively passing the irradiated radiation.

Further, in the present specification, the distribution of the refraction angle of the radiation incident on the radiation image detector is calculated from the radiation image acquired by the radiation imaging apparatus and the radiation image detector, and the refraction angle of the radiation image is calculated. A radiographic imaging system is disclosed that includes a calculation unit that generates a phase contrast image of a subject based on a distribution.

Although the present invention has been described in detail and with reference to specific embodiments, it will be apparent to those skilled in the art that various changes and modifications can be made without departing from the spirit and scope of the invention.
This application is based on a Japanese patent application filed on October 25, 2010 (Japanese Patent Application No. 2010-239108), the contents of which are incorporated herein by reference.

10 X-ray imaging system 11 X-ray source 12 Imaging unit 13 Console 30 FPD
31 First Absorption Type Grating 32 Second Absorption Type Grating 33 Scanning Mechanism 34 Scattering Removal Grating 40 Pixel

Claims

A radiation source;
A first grating that forms a first periodic pattern image by passing radiation emitted from the radiation source;
A second grating disposed on the opposite side of the first grating relative to the radiation source and forming a second periodic pattern image generated by partially shielding the first periodic pattern image;
A radiation image detector for detecting the second periodic pattern image;
A scatter removing grating disposed between the second grating and the radiation image detector for removing scattered radiation;
A radiographic apparatus comprising:
The radiographic apparatus according to claim 1,
A radiation imaging apparatus further comprising: a scanning mechanism that moves one of the first grating and the second grating and places the second grating at the plurality of relative positions with respect to the radiation image.
The radiographic apparatus according to claim 1 or 2,
The scatter removal grating is a radiation imaging apparatus that is stationary at least while the second periodic pattern image is detected by the radiation image detector.
The radiographic apparatus according to any one of claims 1 to 3,
The scatter removing grating is a radiation imaging apparatus that can be retracted from a radiation irradiation field.
A radiographic apparatus according to any one of claims 1 to 4, wherein
The scatter removal grating has a plurality of radiation shielding portions arranged in parallel at intervals from each other,
A radiation imaging apparatus in which an arrangement direction of the plurality of radiation shielding portions and a pitch direction of the first grating are orthogonal to each other.
The radiographic apparatus according to any one of claims 1 to 5,
The radiographic apparatus wherein the first grating is a phase type grating.
The radiographic apparatus according to any one of claims 1 to 5,
The radiographic apparatus wherein the first grating is an absorption grating.
The radiographic apparatus according to any one of claims 1 to 7,
A radiation imaging apparatus further comprising a third grating that irradiates the first grating by selectively passing the irradiated radiation.
A radiation imaging apparatus according to any one of claims 1 to 8,
A computing unit that computes a refraction angle distribution of radiation incident on the radiation image detector from a radiation image acquired by the radiation image detector and generates a phase contrast image of the subject based on the refraction angle distribution. When,
A radiography system comprising: