WO2012057046A1

WO2012057046A1 - Radiography device and radiography system

Info

Publication number: WO2012057046A1
Application number: PCT/JP2011/074365
Authority: WO
Inventors: 拓司多田; 岩切　直人
Original assignee: 富士フイルム株式会社
Priority date: 2010-10-28
Filing date: 2011-10-21
Publication date: 2012-05-03
Also published as: JP2014014380A

Abstract

The present invention generates a high-precision radiation phase contrast image by appropriate exposure control. An X-ray imaging system (10) comprises: a first grid (31); a second grating (32) having a period that substantially matches the pattern period of a radiation picture formed by radiation that passes through the first grating; a radiation image detector (30) for detecting the radiation picture masked by the second grating; a dose detector (35) for detecting the incident radiation dose, the dose detector being positioned upstream from the second grating; and a controller (22) for controlling exposure on the basis of the dose detected by the dose detector.

Description

Radiography apparatus and 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 an X-ray image, and a transmission image of the subject is captured. In this case, each X-ray radiated from the X-ray source toward the X-ray image detector has characteristics (atomic number, density, thickness) of the substance constituting the subject existing on the path to the X-ray image detector. ), The light is incident on the X-ray image detector. As a result, an X-ray transmission image of the subject is detected and imaged by the X-ray image detector. As an X-ray image detector, there is a flat panel detector (FPD: Flat Panel Detector) using a semiconductor circuit in addition to a combination of an X-ray intensifying screen and a film, a stimulable phosphor (accumulating phosphor), and so on. Widely used. In addition, automatic exposure is used to stabilize the density of the image obtained by the X-ray image detector with respect to the required exposure amount that varies depending on the subject, or to prevent excessive exposure of the subject due to excessive exposure. Control is taking place. In the automatic exposure control, generally, the dose of X-rays transmitted through the subject is detected by a dose detector, and the X-ray irradiation is stopped when the dose detected by the dose detector reaches a preset threshold dose. .

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

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

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

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.

Patent Document 1 describes that the above automatic exposure control is performed in the X-ray phase imaging by the fringe scanning method using the first and second diffraction gratings.

International Publication No. 08/102598

In X-ray phase imaging by the fringe scanning method using the first and second diffraction gratings, the radiation image formed by the X-rays that have passed through the first diffraction grating is superimposed on the second diffraction grating as described above. As a result, moire fringes are formed on the image receiving surface of the X-ray image detector downstream of the second diffraction grating. Therefore, although the image detected by the X-ray image detector includes moire fringes, an X-ray phase contrast image from which the influence of moire fringes has been finally removed is presented in medical diagnosis and non-destructive inspection. It is preferable. However, depending on the degree of the grating pitch error of the first and second diffraction gratings, the scanning error of the second diffraction grating, etc., the moire fringes from the X-ray phase contrast image become shorter as the period of the moire fringes becomes shorter. It may be difficult to remove the influence of the image quality, and the image quality may deteriorate. Therefore, it is preferable that the period of the moire fringes is long.

Patent Document 1 does not particularly describe the position of the dose detector, but generally the dose detector is disposed behind the X-ray image detector. In that case, in the X-ray phase imaging based on the fringe scanning method using the first and second diffraction gratings, the dose detector is positioned downstream of the second diffraction grating, and thus on the light receiving portion of the dose detector. In this case, moire fringes are formed. As shown in FIG. 18, this moire fringe moves with the scanning of the second diffraction grating (FIG. 18A), and when the period of the moire fringes becomes longer than the dimension of the light receiving part of the dose detector in the period direction, The dose of X-rays incident on the dose detector per unit time varies greatly depending on whether the dark part of the moire fringes overlaps the light receiving part or when the dark part does not overlap the light receiving part.

The above automatic exposure control extends or shortens the X-ray irradiation time so as to cancel the fluctuation of the X-ray dose incident on the light receiving unit of the dose detector per unit time. As a result, there is a variation in the irradiation dose between photographings, and the variation in the irradiation dose between photographings causes a change in the signal value of each pixel separately from the scanning of the second diffraction grating. For example, X-ray image detection in which the dark part of moire fringes overlaps with the light receiving part of the dose detector for the change in signal value that should be originally obtained by scanning the second diffraction grating when the irradiation dose is constant In the pixel A of the detector, the change in the signal value is attenuated (FIG. 18B), and in the pixel B of the X-ray image detector in which the dark part of the moire fringes overlaps with the light receiving part of the dose detector, the signal The change in value is emphasized (FIG. 18C).

As described above, the X-ray phase imaging by the fringe scanning method using the first and second diffraction gratings detects the phase information of the subject from the change in the signal value of each pixel accompanying the scanning of the second diffraction grating. Thus, a change in the signal value of each pixel due to a factor different from the scanning of the second diffraction grating reduces the detection accuracy of the phase information of the subject.

The present invention has been made in view of the above-described circumstances, and an object thereof is to generate a highly accurate radiation phase contrast image by appropriate exposure control.

Detecting a first grating, a grating pattern having a period substantially matching a pattern period of a radiation image formed by radiation that has passed through the first grating, and the radiation image masked by the grating pattern A radiation imaging apparatus comprising: a radiation image detector; and a dose detector that is located upstream of the grating pattern in a traveling direction of radiation passing through the grating pattern and detects an incident radiation dose.

According to the present invention, by arranging the dose detector upstream of the lattice pattern, the dose detector detects the dose without being affected by the moire fringes formed by superimposing the radiation image and the lattice pattern. be able to. Thereby, appropriate exposure control can be performed, and a highly accurate radiation phase contrast image can be generated.

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 which shows an example of the dose detection area 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 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 radiographic image detector regarding the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the rotation mechanism which rotates the 1st grating | lattice and grating | lattice pattern centering on the axis | shaft of a radiation irradiation field regarding 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 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. It is a schematic diagram for demonstrating the conventional automatic exposure control.

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 the operation of the input device 21, an X-ray tube voltage or an X-ray dose detected by a dose detector described later can be used. X-ray imaging conditions such as a threshold dose, imaging timing, and the like are input. 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. An absorption grating 32 and a dose detector 35 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. Although described in detail later, the first and

second absorption gratings

31 and 32 are disposed between the FPD 30 and the X-ray source 11.

The dose detector 35 is arranged between the first absorption type grating 31 and the second absorption type grating 32 and is located downstream of the subject H, as will be described in detail later. The dose detector 35 detects the X-ray dose transmitted through the subject H, and sends a signal indicating that the threshold dose has been reached to the control device 20 when the detected dose reaches the threshold dose. Receiving this signal, the control device 20 sends a control signal instructing to stop the irradiation of X-rays to the X-ray source control unit 17, and the X-ray control unit 17 receiving this control signal sends the control signal to the X-ray tube 18. The high voltage generator 16 is controlled to stop the supply of power. Thereby, irradiation of the subject H with X-rays is stopped. As the X-ray light receiving element of the dose detector 35, for example, a combination of a phosphor and a photomultiplier tube, an ion chamber, an X-ray detector using a semiconductor circuit, or the like is used. In addition, while the supply of power to the X-ray tube 18 is continued, X-ray irradiation and stop may be switched by opening and closing the collimator 19, or a disk shape (or slit) in which openings and shields are alternately formed. X-ray irradiation and stop may be switched by providing a shutter plate of the shape) at the emission port of the X-ray source 11 and rotating (or translating) it in synchronization with the X-ray irradiation timing. According to this, it is possible to keep the X-ray tube 18 in a stable state and more reliably prevent variations in irradiation dose.

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. Each pixel 40 is connected to a thin film transistor (TFT) switch (not shown), and the gate electrode of the TFT switch is connected to the scanning line 45, the source electrode is connected to the capacitor, and the drain electrode is connected to the signal line 46. When the TFT switch is turned on by the drive pulse from the scanning circuit 42, the charge accumulated in the capacitor is read out to the signal line 46.

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

The readout circuit 43 includes an integration amplifier circuit, an A / D converter, a correction circuit, and an image memory (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 X-rays. 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 completely shield (absorb) X-rays in order to generate a periodic pattern image with high contrast, but the above-described materials (gold, platinum) having excellent X-ray absorption properties Etc.), there are not a few X-rays that are transmitted without being absorbed. Therefore, in order to enhance the shielding of the X-rays, the X-ray shielding portion 31b, the respective thicknesses _h 1, _{h 2} of 32b, it is preferable to increase the thickness much as possible. For example, when the tube voltage of the X-ray tube 18 is 50 kV, it is preferable to shield 90% or more of the irradiated X-rays. In this case, the thicknesses h ₁ and h ₂ are 30 μm or more in terms of gold (Au). It is preferable that

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

X-ray shielding portions

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

X-ray shielding portions

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

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

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

second absorption gratings

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

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

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

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.

FIG. 7 shows one X-ray refracted according to the phase shift distribution Φ (x) of the subject H in the x direction. The dose detector 35 is not shown.

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). 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 above calculation, the change of the M signal values of each pixel 40 for calculating the phase shift amount ψ needs to be brought about by the scanning of the second absorption type grating 32. For this purpose, the X-ray irradiation dose irradiated from the X-ray source 11 to the imaging unit 12 is required to be substantially constant during imaging.

In the present X-ray imaging system 10, during each imaging, automatic exposure control is performed to stop X-ray irradiation when the X-ray dose detected by the dose detector 35 reaches a preset threshold dose. The irradiation dose is kept almost constant. The dose detector 35 is disposed between the first absorption type grating 31 and the second absorption type grating 32 and is located upstream of the second absorption type grating 32. Therefore, the G1 image on the light receiving unit of the dose detector 35 does not move even by scanning the second absorption grating 32, and the dose of X-rays incident on the dose detector 35 per unit time is constant. Therefore, the time until the dose detected by the dose detector 35 reaches the above threshold dose is constant between imagings, and variations in irradiation dose between imagings are prevented.

In addition, as described above, the G1 image formed by the X-rays that have passed through the first absorption type grating 31 is superimposed on the second absorption type grating 32 so that it is downstream of the second absorption type grating 32. Moire fringes are formed on the image receiving surface of the FPD 30. On the other hand, the dose detector 35 is located upstream of the second absorption-type grating 32, and the G1 image on the light receiving portion of the dose detector 35 does not have moire fringes, and the pattern period thereof is the first absorption. Corresponding to the grating pitch p ₁ of the mold grating 31, the order is μm. The size of the light receiving part of the dose detector 35 is generally on the order of cm. Therefore, the light receiving part of the dose detector 35 is overlapped with countless bright parts and dark parts in the G1 image, and these are averaged and detected. Is done. Accordingly, even when the first absorption-type grating 31 is scanned instead of the second absorption-type grating 32 and the G1 image moves as the first absorption-type grating 31 is scanned, the dose detector 35 has a unit. The dose of X-rays incident per hour is almost constant. Therefore, the irradiation time of X-rays until the dose detected by the dose detector 35 reaches the above threshold dose is substantially constant between imaging, and variation in irradiation dose between imaging is prevented.

In addition, the moire fringes formed downstream of the second absorption grating 32 do not affect the dose detection in the dose detector 35 disposed upstream of the second absorption grating 32. Therefore, the period T of the moire fringes can be made longer than, for example, the size of the light receiving part of the dose detector 35 by the relative rotation mechanism 50, the relative tilt mechanism 51, and the relative movement mechanism 52 described above. As a result, the influence of moire fringes can be sufficiently removed from the X-ray phase contrast image, and the image quality of the X-ray phase contrast image can be improved.

FIG. 10 shows an example of the arrangement of the light receiving parts of the dose detector 35.

The light receiving portions of the dose detector 35 are preferably distributed in a plane perpendicular to the optical axis A of X-rays, and particularly preferably substantially symmetric with respect to the optical axis A. For example, a plurality of light receiving portions 35a may be provided, and these light receiving portions 35a may be arranged in a distributed manner in a plane perpendicular to the optical axis A of the X-ray, and may be arranged at positions that are substantially symmetrical with respect to the optical axis A. Preferred (FIG. 10A to D). When an area dosimeter is used as a dose detector, even one light receiving unit 35a can be distributed in a plane perpendicular to the optical axis A of the X-ray and is substantially symmetrical with respect to the optical axis A. For example, it is preferable that the light receiving portion 35a has a frame shape surrounding the optical axis A (FIG. 10E). Even when an area dosimeter is used, a plurality of light receiving portions 35a can be provided, and these light receiving portions 35a may be arranged in a distributed manner in a plane orthogonal to the optical axis A of the X-ray. It is preferable to arrange at a position that is substantially symmetrical with respect to A (FIG. 10F). As described above, by distributing the light receiving portions of the dose detector 35 in a plane orthogonal to the optical axis A of the X-ray, stable dose detection can be performed for each imaging.

In the X-ray phase imaging based on the fringe scanning method using the first and

second absorption gratings

31 and 32, X-ray refraction at an object upstream of the second absorption grating 32 is detected. The refraction of X-rays at the dose detector 35 disposed upstream of the second absorption grating 32 is detected without distinction from the refraction of X-rays at the subject H. As the distance between the dose detector 35 and the second absorption grating 32 increases, the displacement amount of the G1 image due to the refraction of X-rays at the dose detector 35 increases. As shown in the above equation (14), this is related to the phase shift amount ψ of the signal of each pixel 40. Therefore, the dose detector 35 is preferably disposed between the first absorption type grating 31 and the second absorption type grating 32, and more preferably, is disposed adjacent to the second absorption type grating 32. The

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

As described above, according to the present X-ray imaging system 10, the dose detector 35 is disposed upstream of the second grating 32, and is formed by superimposing the G1 image and the second grating 32. The dose detector 35 can detect the dose without being affected by the moire fringes. Thereby, appropriate exposure control can be performed, and a highly accurate radiation phase contrast image can be generated.

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.

Note that the above-described X-ray imaging system 10 calculates the refraction angle φ by performing fringe scanning on the projection image of the first grating, and therefore the first and second gratings absorb both. Although described as a mold lattice, the present invention is not limited to this. As described above, the present invention is also useful when the refraction angle φ is calculated by performing fringe scanning on the Talbot interference image. Therefore, the first grating is not limited to the absorption type grating but may be a phase type grating. Further, the method of analyzing the moire fringes formed by superimposing the X-ray image of the first grating and the second grating is not limited to the above-described fringe scanning method. For example, “J. Opt. Soc. Am. Vol” .72, No. 1 (1982) p.156 ”, various methods using Moire fringes, such as a method using Fourier transform / inverse Fourier transform known in the art.

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.

Further, a phase differential image (a differential amount of the phase shift distribution Φ) may be created from an image group acquired by photographing (pre-photographing) in the absence of a subject. This phase differential image reflects the phase unevenness of the detection system (including phase shift due to moire, grid nonuniformity, refraction of the dose detector, etc.). Then, a phase differential image is created from a group of images acquired by shooting (main shooting) in the presence of a subject, and the phase differential image obtained by pre-shooting is subtracted from this to correct phase irregularity in the measurement system. A phase differential image can be obtained.

FIG. 11 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

A mammography apparatus 80 shown in FIG. 11 is an apparatus that captures an X-ray image (phase contrast image) of a 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 first embodiment, the same reference numerals as those of the first embodiment are given to the respective components. Since other configurations and operations are the same as those in the first embodiment, description thereof will be omitted.

FIG. 12 shows a modification of the radiation imaging system of FIG.

The mammography apparatus 90 shown in FIG. 12 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. The first absorption type lattice 31 is accommodated in a lattice accommodation portion 91 connected to the arm member 81. The imaging unit 92 includes an FPD 30, a second absorption grating 32, a scanning mechanism 33, and a dose detector 35. In the illustrated example, the dose detector 35 is disposed between the subject B and the second absorption grating 32.

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

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

FIG. 13 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

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

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

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

absorption type gratings

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

X-ray shielding portions

31b and 32b of the first and

second absorption gratings

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

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

The above formula (18) 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).

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

As described above, in the present X-ray imaging system, the G1 images based on the plurality of 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. be able to. The multi slit 103 described above can be applied to any of the X-ray imaging systems described above.

FIG. 14 shows the configuration of the radiation image detector in relation to 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 second absorption type grating 32 is provided independently of the FPD 30, but the X-ray image detector itself has the second absorption type grating 32 or an equivalent configuration. You may do it. As a specific embodiment, the second absorption type grating can be eliminated by using an X-ray image detector having a configuration disclosed in Japanese Patent Laid-Open No. 2009-133823. This X-ray image detector is a direct conversion type X-ray image detector including a conversion layer that converts X-rays into electric charges, and a charge collection electrode that collects electric charges converted in the conversion layer, The charge collecting electrode 121 of the pixel 120 is configured by arranging a plurality of linear electrode groups 122 to 127 formed by electrically connecting linear electrodes arranged at a constant period so that their phases are different from each other. Has been.

The pixels 120 are two-dimensionally arranged at a constant pitch along the x direction and the y direction, and each pixel 120 has a charge collection for collecting the charges converted by the conversion layer that converts the X-rays into charges. An electrode 121 is formed. The charge collection electrode 121 includes first to sixth linear electrode groups 122 to 127, and the phase of the arrangement period of the linear electrodes of each linear electrode group is shifted by π / 3. Specifically, when the phase of the first linear electrode group 122 is 0, the phase of the second linear electrode group 123 is π / 3, the phase of the third linear electrode group 124 is 2π / 3, The phase of the fourth linear electrode group 125 is π, the phase of the fifth linear electrode group 126 is 4π / 3, and the phase of the sixth linear electrode group 127 is 5π / 3.

Linear electrode groups 122-127 of the first through sixth, respectively, in which the linear electrodes extend in the y-direction and periodically arranged at a predetermined pitch p ₂ in the x-direction. The pattern period p of the G1 image at the substantial pitch p ₂ ′ (substantial pitch after manufacture) of the arrangement pitch p ₂ of the linear electrodes and the position of the charge collection electrode 121 (position of the X-ray image detector). The relationship between ₁ ′ and the arrangement pitch P of the pixels 120 in the x direction is similar to the second absorption grating 32 of the X-ray imaging system 10 described above, and the period T of the moire fringes represented by the equation (8). Therefore, it is necessary to satisfy the formula (9), and it is preferable to satisfy the formula (10).

Further, each pixel 120 is provided with a switch group 128 for reading out the charges collected by the charge collecting electrode 121. The switch group 128 includes TFT switches provided in the first to sixth linear electrode groups 121 to 126, respectively. By collecting the charges collected by the first to sixth linear electrode groups 121 to 126 individually by controlling the switch group 128, six types of fringe images having different phases can be obtained by one imaging. A phase contrast image can be generated based on these six types of fringe images.

For example, when the X-ray image detector configured as described above is applied to the X-ray imaging system 10 described above, the second absorption type grating 32 is not required from the imaging unit 12, and a plurality of images can be obtained by one imaging. Since a phase component fringe image can be acquired, physical scanning for fringe scanning becomes unnecessary, and the scanning mechanism 33 can be eliminated. Thereby, it is possible to reduce the cost and further reduce the thickness of the photographing unit. In this case, the dose detector 35 can be disposed between the subject H and the first absorption type grating 31 or between the first absorption type grating 31 and the X-ray image detector. Considering the influence of X-ray refraction at the detector 35, it is preferably arranged between the first absorption grating 31 and the X-ray image detector, more preferably the X-ray image detection described above. Placed adjacent to the vessel. It should be noted that the structure of the charge collecting electrode may be replaced with another structure described in Japanese Patent Application Laid-Open No. 2009-133823.

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

In the X-ray imaging system 10 described above, the phase contrast image obtained by the fringe scanning is an X-ray refraction component in the periodic array direction (x direction) of the X-ray shielding portions of the first and

second absorption gratings

31 and 32. And the refractive component in the extending direction (y direction) of the X-ray shielding part is not included. For this reason, there is a portion that cannot be depicted depending on the shape and orientation of the subject H. For example, when the direction of the load surface of the articular cartilage is matched with the y direction, it is considered that the peripheral tissue of the cartilage (such as tendons and ligaments) having a shape perpendicular to the load surface is insufficiently depicted. By moving the subject H, it is possible to recapture a region that is not sufficiently depicted, but in addition to increasing the burden on the subject H and the operator, the position reproducibility with the recaptured image is ensured. There is a problem that it is difficult.

Accordingly, the first and

second absorption gratings

31 and 32 are centered on a virtual line (X-ray optical axis A) orthogonal to the center of the grating surface of the first and

second absorption gratings

31 and 32. It is also possible to provide a lattice rotation mechanism 105 that rotates integrally from the first direction and sets the second direction to generate a phase contrast image in each of the first direction and the second direction. Is preferred.

In the illustrated example, the first and

second absorption gratings

31 and 32 are rotated by 90 °, and the first direction and the second direction are orthogonal to each other. As long as the direction intersects, the rotation angle of the first and

second absorption gratings

31 and 32 is not limited to 90 °. Further, the grating rotating mechanism 105 may be configured to rotate only the first and second

absorption type gratings

31 and 32 separately from the FPD 30, or the first and second absorption type gratings 31. , 32 and the FPD 30 may be rotated together. Further, when the multi-slit 103 is provided, the multi-slit 103 and the collimator 109 or the radiation source formed integrally with them is rotated so that the rotation coincides with the first and

second absorption gratings

31 and 32. . Furthermore, the generation of phase contrast images in the first and second orientations using the grating rotation mechanism 105 can be applied to any of the X-ray imaging systems described above.

FIG. 16 shows a configuration of a calculation unit of another example of the radiation imaging system for explaining the 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. 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 I _k (x, y) obtained for each pixel with respect to k and calculating an average value as shown in FIG. To do. The average value may be calculated by simply averaging the pixel data I _k (x, y) with respect to k. However, when M is small, the error increases, so that the pixel data I _k ( After fitting x, 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 I _k (x, y) with respect to k can be used as long as the amount corresponds to the average value.

Note that an absorption image may be created from an image group acquired by photographing (pre-photographing) in the absence of a subject. This absorption image reflects the transmittance unevenness of the detection system (including information such as the transmittance unevenness of the grid and the influence of the absorption of the dose detector). Therefore, a correction coefficient map for correcting the transmittance unevenness of the detection system can be created from this image. Absorption of the subject, in which an absorption image is created from a group of images obtained by shooting in the state of the subject (main shooting), and the above-described correction coefficient is applied to each pixel, thereby correcting the transmittance unevenness of the detection system. An image can be obtained.

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 I _k (x, y) obtained for each pixel. The amplitude value may be calculated by obtaining the difference between the maximum value and the minimum value of the pixel data I _k (x, y). However, when M is small, the error increases, so that the pixel data After fitting I _k (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.

It should be noted that a small angle scattered image may be created from an image group obtained by photographing (pre-photographing) in the absence of a subject. This small-angle scattered image reflects the amplitude value unevenness of the detection system (including information such as grid pitch non-uniformity, aperture ratio non-uniformity, and non-uniformity due to relative displacement between grids). . Therefore, a correction coefficient map for correcting the amplitude irregularity of the detection system can be created from this image. A small-angle scattered image is created from a group of images acquired by shooting (main shooting) in the presence of the subject, and the amplitude value unevenness of the detection system is corrected by applying the correction coefficient described above to each pixel. A small-angle scattered image can be obtained.

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. No deviation occurs, and the phase contrast image can be satisfactorily superimposed with the absorption image or the small-angle scattered image.

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

As described above, the present specification includes a first grating and a grating pattern having a period substantially matching the pattern period of a radiation image formed by radiation that has passed through the first grating; A radiation image detector for detecting the radiation image masked by the grating pattern; a dose detector for detecting an incident radiation dose positioned upstream of the grating pattern in a traveling direction of radiation passing through the grating pattern; A radiation imaging apparatus comprising:

Also, in the radiographic apparatus disclosed in this specification, the dose detector is disposed between the first grating and the grating pattern.

Also, in the radiation imaging apparatus disclosed in this specification, the pattern period of the radiation image masked by the lattice pattern is longer than the dimension of the light receiving unit of the dose detector with respect to the period direction.

Further, the radiographic apparatus disclosed in the specification changes at least one of a relative posture and a relative position of the lattice pattern with respect to the first lattice, and the radiation image masked by the second lattice. And a change mechanism for changing the pattern period.

Further, in the radiation imaging apparatus disclosed in this specification, the changing mechanism causes at least one of the first grating and the grating pattern to be arranged around an optical axis of radiation irradiated to the first grating. Rotate.

Further, in the radiographic apparatus disclosed in this specification, the change mechanism has at least one of the first grating and the grating pattern with respect to an optical axis of radiation applied to the first grating. Tilt.

Further, in the radiation imaging apparatus disclosed in this specification, the changing mechanism may change at least one of the first grating and the grating pattern along an optical axis of radiation irradiated on the first grating. Move.

Further, in the radiographic apparatus disclosed in this specification, the light receiving units of the dose detector are provided in a distributed manner in a plane orthogonal to the optical axis of the radiation irradiated on the first grating.

Also, in the radiation imaging apparatus disclosed in this specification, the grating pattern is placed at a plurality of relative positions having different phases from each other with respect to the radiation image.

Further, in the radiation imaging apparatus disclosed in this specification, the lattice pattern is a second lattice, and the second lattice is moved, and the second lattice is moved. A scanning mechanism is further provided that places a grating at the plurality of relative positions with respect to the radiation image.

Also, in the radiographic apparatus disclosed in this specification, the dose detector is disposed adjacent to the second grating.

Further, in the radiation imaging apparatus disclosed in this specification, the lattice pattern is provided in the radiation image detector.

Further, in the radiation imaging apparatus disclosed in this specification, the radiological image detector includes a conversion layer that converts radiation into electric charge, and a charge collection electrode that collects electric charge converted in the conversion layer, for each pixel. The charge collection electrode includes a plurality of linear electrode groups having a period substantially matching the pattern period of the radiation image, and the plurality of linear electrode groups are arranged so that their phases are different from each other. The lattice pattern is constituted by each of the plurality of linear electrode groups.

Further, in the radiation imaging apparatus disclosed in this specification, the dose detector is provided adjacent to the radiation image detector.

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

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

The radiation imaging apparatus disclosed in this specification further includes a radiation source that irradiates radiation toward the first grating.

Further, the present specification discloses a radiation imaging system including any one of the above-described radiation imaging apparatuses and a control unit that performs exposure control based on the dose detected by the dose detector.

Further, in the radiation imaging system disclosed in this specification, the lattice pattern is placed at a plurality of relative positions having different phases with respect to the radiation image, and the radiation image detector has the lattice pattern different from each other. In each imaging step placed at the relative position, the radiation image masked by the lattice pattern is detected, and the dose detected by the dose detector is constant between the imaging steps. The exposure is controlled so that

Further, the radiation imaging system disclosed in this specification calculates a distribution of refraction angles of radiation incident on the radiation image detector from a radiation image acquired by the radiation image detector of the radiation detector, A calculation unit is further provided for generating a phase contrast image of the subject based on the distribution of the refraction angles.

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. 2010-242779) filed on Oct. 28, 2010, 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 35 Dose Detector 40 Pixel

Claims

A first lattice;
A grating pattern having a period substantially matching a pattern period of a radiation image formed by radiation that has passed through the first grating;
A radiation image detector for detecting the radiation image masked by the lattice pattern;
A dose detector that is located upstream of the grating pattern in the traveling direction of the radiation passing through the grating pattern and detects the amount of incident radiation;
A radiographic apparatus comprising:
The radiographic apparatus according to claim 1,
The dose detector is a radiographic apparatus arranged between the first grating and the grating pattern.
The radiographic apparatus according to claim 1 or 2, wherein
A radiation imaging apparatus, wherein a pattern period of the radiation image masked by the lattice pattern is longer than a dimension of a light receiving unit of the dose detector with respect to the period direction.
The radiographic apparatus according to any one of claims 1 to 3,
A radiation imaging apparatus further comprising: a changing mechanism that changes a pattern period of the radiation image masked by the second grating by changing at least one of a relative posture and a relative position of the grating pattern with respect to the first grating. .
The radiographic apparatus according to claim 4,
The change mechanism is a radiation imaging apparatus that rotates at least one of the first grating and the grating pattern around an optical axis of radiation applied to the first grating.
The radiographic apparatus according to claim 4,
The change mechanism is a radiation imaging apparatus that tilts at least one of the first grating and the grating pattern with respect to an optical axis of radiation applied to the first grating.
The radiographic apparatus according to claim 4,
The change mechanism is a radiation imaging apparatus that moves at least one of the first grating and the grating pattern along an optical axis of radiation applied to the first grating.
The radiographic apparatus according to any one of claims 1 to 7,
The light-receiving unit of the dose detector is a radiation imaging apparatus provided in a distributed manner in a plane orthogonal to the optical axis of radiation applied to the first grating.
The radiographic apparatus according to any one of claims 1 to 8,
The radiographic apparatus in which the lattice pattern is placed at a plurality of relative positions whose phases are different from each other with respect to the radiation image.
The radiographic apparatus according to claim 9, wherein
The lattice pattern is a second lattice,
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 10, wherein
The dose detector is a radiation imaging apparatus arranged adjacent to the second grating.
The radiographic apparatus according to claim 9, wherein
The grid pattern is a radiation imaging apparatus provided in the radiation image detector.
The radiographic apparatus according to claim 12, comprising:
The radiation image detector includes, for each pixel, a conversion layer that converts radiation into charges, and a charge collection electrode that collects charges converted in the conversion layer,
The charge collection electrode includes a plurality of linear electrode groups having a period substantially matching the pattern period of the radiation image,
The plurality of linear electrode groups are arranged so that their phases are different from each other,
The grid pattern is a radiation imaging apparatus configured by each of the plurality of linear electrode groups.
A radiographic apparatus according to claim 12 or claim 13, wherein
The dose detector is a radiation imaging apparatus provided adjacent to the radiation image detector.
The radiographic apparatus according to any one of claims 1 to 14,
The radiographic apparatus wherein the first grating is an absorption grating.
The radiographic apparatus according to any one of claims 1 to 14,
The radiographic apparatus wherein the first grating is a phase type grating.
The radiographic apparatus according to any one of claims 1 to 16, wherein
A radiation imaging apparatus further comprising a radiation source that irradiates radiation toward the first grating.
A radiation imaging apparatus according to any one of claims 1 to 17,
A control unit for controlling exposure based on a dose detected by the dose detector;
A radiography system comprising:
The radiation imaging system according to claim 18,
The grating pattern is placed at a plurality of relative positions with different phases with respect to the radiation image,
The radiological image detector detects the radiographic image masked by the grid pattern in each imaging step in which the grid pattern is placed at the relative position different from each other,
The said control part is a radiography system which controls exposure so that the dose detected by the said dose detector becomes fixed between each said imaging step.
The radiation imaging system according to claim 19,
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 image detector of the radiation detector, and the phase contrast image of the subject is calculated based on the distribution of the refraction angle. A radiation imaging system further comprising a calculation unit for generating the.