WO2012070661A1

WO2012070661A1 - Radiographic image detection apparatus, radiography apparatus, and radiography system

Info

Publication number: WO2012070661A1
Application number: PCT/JP2011/077260
Authority: WO
Inventors: 拓司多田; 裕康石井
Original assignee: 富士フイルム株式会社
Priority date: 2010-11-26
Filing date: 2011-11-25
Publication date: 2012-05-31
Also published as: JP2014030438A

Abstract

The invention accurately detects radiation doses and generates more precise phase contrast radiographs. The X-ray imaging system (10) is provided with a first grating (31), a second grating (32) having a period that substantially coincides with the period of the radiographic pattern formed by radiation passing through the first grating, a radiographic image detector (30) that detects the radiograph masked by the second grating, and a data processing unit (22) that processes the image data acquired by the radiographic image detector. The subject is disposed between the radiation source (11) and the first grating, the radiation irradiation time is made to be constant, and imaging is performed multiple times placing the second grating at positions relative to the radiographic image such that the phases differ from each other. Brightness is corrected for the image data acquired by the radiographic image detector in each imaging on the basis of the radiation dose detected by a radiation dose-detecting pixel during said imaging.

Description

Radiation image detection apparatus, radiation imaging apparatus, and radiation imaging system

The present invention relates to a radiation image detection apparatus, and a radiation imaging apparatus and a radiation imaging system including the radiation image detection apparatus.

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 the above X-ray imaging system, the exposure of the subject is prevented from being excessively exposed 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. Therefore, automatic exposure control is performed. In automatic exposure control, in general, the dose of X-rays transmitted through a subject is detected by a dose detector, and X-ray irradiation is stopped when the dose detected by the dose detector reaches a preset threshold value. . In addition, there is also known one that detects a dose using a part of pixels of an FPD and controls exposure by that. In addition, when detecting a dose using a part of the pixels of the FPD, since it is difficult to detect the dose in real time in relation to the time required to read out the signal of the pixel, in the fluoroscopy performed prior to imaging, The X-ray irradiation conditions are adjusted so that the dose detected by the dose detection pixels becomes a predetermined value set in advance. In imaging, the irradiation time is calculated based on the X-ray irradiation conditions at the time of fluoroscopy, further taking into account the tube voltage difference between fluoroscopy and imaging, and exposure control is performed based on the calculated irradiation time.

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 the X-ray refracted by the subject from the change of the pixel value of each pixel obtained by the X-ray image detector while performing a plurality of times of imaging while translating in the vertical direction at a scanning pitch obtained by equally dividing the lattice pitch. A distribution (differential image of phase shift) is calculated, and a phase contrast image of the subject can be obtained based on this angular distribution.

International Publication No. 08/102598

As described above, the X-ray phase imaging by the fringe scanning method detects the phase information of the subject from the change in the pixel value of each pixel accompanying the scanning of the second diffraction grating. A change in the pixel value of each pixel due to another factor reduces the detection accuracy of the phase information of the subject. Factors that change the pixel value of each pixel include, for example, variations in irradiation dose between photographings. Therefore, the irradiation dose between photographings is kept constant, or variations in the irradiation dose between photographings are measured in advance. When calculating the angle distribution of the X-rays refracted by the subject, it is necessary to correct the change in the pixel value of each pixel due to the variation in irradiation dose.

Here, in the X-ray phase imaging apparatus described in Patent Document 1, a dose detector is used, and X-ray irradiation is stopped when the dose detected by the dose detector reaches a preset threshold value. Automatic exposure control is performed. Although the patent document 1 does not particularly describe the position of the dose detector, the dose detector is generally arranged behind the FPD. In this case, in the X-ray phase imaging by the fringe scanning method, the dose detector is positioned downstream of the second diffraction grating, and moire fringes are formed on the dose detector. This moire fringe moves with the scanning of the second diffraction grating, and the X-rays incident on the dose detector per unit time when the dark part of the moire fringe overlaps the dose detector and when the dark part does not overlap. Dose varies greatly. The automatic exposure control described above extends or shortens the X-ray irradiation time so as to cancel the fluctuation of the X-ray dose incident on the dose detector per unit time. As a result, there is a variation in the irradiation dose during imaging.

Therefore, it is conceivable to perform exposure control with a constant irradiation time. However, even if the irradiation time is fixed, variations in irradiation dose may occur between imaging due to, for example, the rising characteristics and falling characteristics of the X-ray source. Therefore, it is necessary to measure the variation of the irradiation dose between the imagings, but as described above, the overlap between the dose detector and the dark part of the moire fringe changes as the second diffraction grating is scanned. Since the dose of X-rays incident on the dose detector within the irradiation time changes, it is difficult to measure variations in the irradiation dose between radiographs under moire fringes.

The above is also valid when a dose is detected by using some pixels of the FPD instead of the dose detector. The FPD is located downstream of the second diffraction grating, and moire fringes are formed on the detection surface. Therefore, as the second diffraction grating is scanned, the overlap between the dose detector and the dark portion of the moire fringe changes, and the X-ray dose incident on the dose detector per unit time changes accordingly. Therefore, even when a dose is detected using a part of the pixels of the FPD, it is still difficult to measure the variation of the irradiation dose between the images under the moire fringes.

The present invention has been made in view of the above-described circumstances, and an object thereof is to accurately detect a dose and generate a more accurate radiation phase contrast image.

(1) masked by the first grating, the second grating having a period substantially matching the pattern period of the radiation image formed by the radiation that has passed through the first grating, and the second grating. A radiation image detector for detecting the radiation image, wherein the radiation image detector is incident with radiation propagating off at least one of the first grating region and the second grating region; A radiological image detection apparatus including at least one dose detection pixel used for detecting a radiation dose incident thereon.
(2) A radiographic apparatus comprising: the radiological image detection apparatus according to (1) above; and a radiation source that emits radiation toward the first grating.
(3) The radiation image detection device according to (1), a radiation source that emits radiation toward the first grating, and an arithmetic processing unit that processes image data acquired by the radiation image detector. An object is disposed between the radiation source and the first grating or between the first grating and the second grating, and the second grating is in phase with the radiation image. There is an imaging mode in which imaging is performed a plurality of times at different relative positions, and the arithmetic processing unit detects the dose of the image data acquired by the radiation image detector in each imaging in the imaging mode. A radiography system that corrects brightness based on the dose detected by a pixel.

According to the present invention, radiation that propagates out of at least one of the first grating and the second grating is detected by the dose detection pixel, and the radiation image of the first grating is detected on the dose detection pixel. The moire fringes are not formed by superimposing the second grating and the second grating, so that the dose can be accurately detected by the dose detection pixels without being affected by the moire fringes. Thereby, a plurality of times of imaging performed by placing the subject between the radiation source and the first grating or between the first grating and the second grating and placing the second grating at each relative position. Thus, it is possible to accurately measure the variation of the irradiation dose between the photographings. Then, by correcting the brightness of the image data acquired in each shooting based on the dose detected in the shooting, the change in the pixel value of each pixel due to the variation in the irradiation dose can be removed or reduced. . Thereby, a more 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. 3 is a flowchart for explaining a phase contrast image generation process by the radiation imaging system of FIG. 1. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. 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 example of arrangement | positioning of the dose detection pixel in a radiographic image detector. It is a schematic diagram which shows the other example of arrangement | positioning of the dose detection pixel in a radiographic image detector. It is a schematic diagram which shows the structure of the imaging | photography part regarding an example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the imaging | photography part regarding the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the imaging | photography part regarding the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the radiation attenuation body regarding the other modification of the radiography system of FIG. It is a graph which shows an example of the output characteristic of the pixel of a radiographic image detector. 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 imaging | photography part of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a graph which shows the signal of the pixel of the radiographic image detector for demonstrating the process in the arithmetic processing 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 (radiation image detection device) 12 that detects X-rays transmitted through the subject H from the X-ray source 11 and generates image data, and an exposure operation and an imaging unit of the X-ray source 11 based on the operation of the operator The console 13 is roughly divided into a console 13 that controls the image capturing operation 12 and calculates the image data acquired by the image capturing unit 12 to generate a phase contrast image.

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 of the X-ray that does not contribute to imaging of 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 is 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 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.

X-ray shielding portion 31b is in a plane perpendicular to the optical axis A of the X-ray, with grating pitch p ₁ in the constant direction (x-direction) orthogonal to the one direction, arranged at predetermined intervals d ₁ from each other Has been. Similarly, X-ray shielding portion 32b, in the plane orthogonal to the optical axis A of the X-ray, with grating pitch p ₂ of the constant in the direction (x-direction) orthogonal to the one direction, the predetermined distance d ₂ from each other It is arranged in a space. 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 type gratings

31 and 32 are configured to project the X-rays that have passed through the slit portion almost geometrically regardless of the presence or absence of the Talbot interference effect. More specifically, by setting the distances d ₁ and d ₂ to a value sufficiently larger than the effective wavelength of X-rays emitted from the X-ray source 11, most of the irradiated X-rays are not diffracted at the slit portion. In addition, a self-image of the first absorption type grating 31 can be formed behind the first absorption type grating 31. For example, when tungsten is used as the target of the radiation source and the tube voltage is 50 kV, the effective wavelength of X-ray is about 0.4 mm. In this case, if the distances d ₁ and d ₂ are set to about 1 to 10 μm, the radiation image formed by the radiation that has passed through the slit portion becomes such that the effect of diffraction can be ignored, and the first absorption grating 31 can be ignored. The self-image of the first absorption-type grating 31 is projected almost geometrically.

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 (effective wavelength) lambda, and using the positive integer m, is expressed by the following equation (2).

Equation (2) is an equation representing the Talbot interference distance when the X-ray irradiated from the X-ray source 11 is a cone beam, “Timm Weitkamp, et al., Proc. Of SPIE, Vol. 6318, 2006. It can be obtained from the formula described in “Year Salary 63180S”.

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 100 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. .

Here, the FPD 30 is a pixel on which radiation propagating off the lattice area of the first and second absorption type gratings 31 and 32 (area where the

X-ray shielding portions

31b and 32b are periodically arranged) is incident. 40. That is, in the projection onto the detection surface of the FPD 30 with the X-ray focal point 18b as the viewpoint, the projections of the first and

second absorption gratings

31 and 32 are substantially coincident, and the detection surface of the FPD 30 is It is larger than the projection of the second

absorption type gratings

31 and 32. Then, on the detection surface of the FPD 30, the group of pixels 40 belonging to the region 30A where the projections of the first and

second absorption gratings

31 and 32 overlap is the G1 image of the first absorption grating 31 and the second absorption type. An image whose intensity is modulated by superimposing with the grating 32 is detected (hereinafter, pixels belonging to the region 30A are referred to as image detection pixels). In addition, in the pixels 40 belonging to the region 30B outside the projection of the first and second

absorption type gratings

31 and 32, radiation propagating out of the grating regions of the first and second

absorption type gratings

31 and 32 is transmitted. Incident. In the illustrated example, the region 30 </ b> B is provided along one side of the detection surface of the FPD 30. Each of the plurality of pixels 40 belonging to the region 30B is used for detecting the dose of radiation incident thereon (hereinafter, the pixel belonging to the region 30B is referred to as a dose detection pixel).

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 on the detector surface 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).

In addition, it is possible to detect moire fringes even if the arrangement pitch P is larger than the moire period T within a 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.

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 present X-ray imaging system 10, the second absorption grating 32 is moved by the scanning mechanism 33 described above, but the first absorption 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. Such a change in moire fringes is obtained by photographing the moire fringes with the FPD 30 while moving the second absorption grating 32 by an integer of the grating pitch p _2, and from each of the photographed plural fringe images, 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 photographing is performed by the FPD 30 at each position of k = 0, 1, 2,..., M−1, M pixel values 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 pixel values will be described. When the pixel 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 pixel 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. Therefore, the refraction angle φ (x) is obtained by calculating the phase shift amount ψ of the signal of each pixel 40 from the M pixel values obtained in 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 pixel 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 type grating 32. A broken line in FIG. 9 indicates a change in the pixel value when the subject H does not exist, and a solid line in FIG. 9 indicates a change in the pixel 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 equation (13), the phase shift distribution is obtained by integrating the refraction angle φ (x) along the x-axis. 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).

FIG. 10 shows a flow of a phase contrast image generation process in the radiation imaging system of FIG.

In the present X-ray imaging system 10, the X-ray irradiation time is constant, and the second absorption grating 32 is placed at each position of K = 0, 1,..., M−1 (see FIG. 8). Times are taken.

In each imaging, the control device 20 sends a control signal instructing the start of X-ray irradiation to the X-ray source control unit 17. The X-ray control unit 17 that has received this control signal controls the high voltage generator 16 so as to start supplying power to the X-ray tube 18. Thereby, irradiation of the subject H with X-rays is started (step S1).

An image formed by superimposing the G1 image of the first absorption type grating 31 modulated by the subject H and the second absorption type grating 32 is picked up by the group of image detection pixels 40 of the FPD 30. At that time, X-rays propagating off the first and

second absorption gratings

31 and 32 enter the group of dose detection pixels 40 of the FPD 30, and charges corresponding to the dose of the incident X-rays are those doses. Accumulated in the detection pixel 40.

The control device 20 measures an elapsed time T after sending a control signal instructing the X-ray source control unit 17 to start X-ray irradiation, and the elapsed time T is set to a preset irradiation time T ₀ . When it reaches, a control signal for instructing to stop X-ray irradiation is sent to the X-ray source control unit 17 (step S2).

The X-ray control unit 17 receives the control signal sent from the control device 20 and controls the high voltage generator 16 to stop the supply of power to the X-ray tube 18. Thereby, irradiation of the subject H with X-rays is stopped (step S3).

After the X-ray irradiation is stopped, image data is output from the FPD 30 (step S4), and the arithmetic processing unit 22 performs luminance correction described later on the image data output from the FPD 30 (step S5).

Through the above process, M times of imaging are performed while the second absorption type grating 32 is placed at each position of K = 0, 1,... The arithmetic processing unit 22 calculates the phase shift distribution Φ according to the above-described procedure using the image data acquired by M times of photographing and subjected to luminance correction, and stores this in the storage unit 23 as a phase contrast image (step S6).

Next, luminance correction for each image data will be described.

In the above-described phase contrast image generation processing, the X-ray irradiation time in each imaging is constant, but the irradiation dose varies between imaging due to the rise characteristics and fall characteristics of the X-ray tube 18. Due to the variation in irradiation dose between photographings, an overall luminance change occurs between image data. Therefore, the arithmetic processing unit 22 performs luminance correction on each image data.

The pixel value (luminance) of each pixel in the image data corresponds to the X-ray dose incident on that pixel. X-rays propagating out of the grating regions of the first and

second absorption gratings

31 and 32 are incident on the dose detection pixel 40, and thus the G1 image and the second absorption of the first absorption grating 31 are incident. Moire fringes due to superimposition with the mold grating 32 are not formed on the dose detection pixels 40. Therefore, in the dose detection pixel 40, the dose of X-rays incident within the irradiation time does not change due to the change of the overlap with the dark part of the moire fringe, and the pixel value of the dose detection pixel 40 varies between the image data. Is due to variations in irradiation dose between imaging.

The arithmetic processing unit 22 performs luminance correction on each image data so that the pixel value of the dose detection pixel 40 in each image data is matched between the image data. In addition, when there are a plurality of dose detection pixels 40, luminance correction is performed on each image data so that the sum or average of the pixel values is matched. For example, the pixel value of the dose detection pixel 40 in the image data acquired by the first imaging is used as the reference value. Then, in each image data acquired in the second and subsequent imaging, all the pixels of the image data are multiplied by the reciprocal of the ratio of the pixel value of the dose detection pixel 40 included in the image data to the reference value. Thereby, the pixel value of the dose detection pixel 40 of each image data acquired in the second and subsequent imaging is adjusted to the reference value, and the pixel value of each pixel 40 excluding the dose detection pixel 40 is the pixel of the other pixel 40. Correction is performed while maintaining the ratio to the value. Thereby, the change of the pixel value of each pixel 40 resulting from the variation of the irradiation dose between imaging | photography is removed or reduced. It should be noted that the X-ray irradiation time does not need to be constant because the change in the pixel value of each pixel 40 due to the variation in the irradiation dose during imaging is eliminated or reduced by this luminance correction.

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 X-ray propagating out of the lattice area of the first and

second absorption gratings

31 and 32 is detected by the dose detection pixel 40, Moire fringes due to the superposition of the G1 image of the first absorption-type grating 31 and the second absorption-type grating 32 are not formed on the dose detection pixels, and therefore the dose can be accurately measured without being affected by the moire fringes. Can be detected. Thereby, the variation of the irradiation dose between imaging | photography can be measured correctly. Then, by correcting the brightness of the image data acquired in each shooting based on the dose detected in the shooting, the change in the pixel value of each pixel 40 due to the variation in the irradiation dose between the shootings is removed or reduced. can do. Thereby, an X-ray phase contrast image with higher accuracy can be generated.

Although only one dose detection pixel 40 is required, it is preferable to provide a plurality of dose detection pixels 40 as in the present X-ray imaging system 10 and calculate the average or sum of the pixel values. According to this, the accuracy of dose detection can be increased, and the accuracy of luminance correction of each image data can be increased.

Further, according to the present X-ray imaging system 10, since most X-rays are not diffracted by the first absorption type grating 31 and projected almost geometrically onto the second absorption type grating 32, the irradiated X-rays are obtained. Therefore, a high spatial coherence is not required, and a general X-ray source used in the medical field as the X-ray source 11 can be used. 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).

Note that the X-ray imaging system 10 performs a fringe scan on the projection image of the first grating to calculate the refraction angle φ. Therefore, both the first and second gratings are absorption type. Although described as being a 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” Various methods using Moire fringes, such as a method using Fourier transform / inverse Fourier transform known as “.72, No. 1 1982 (1982) P.156”, can also be applied.

In addition, the phase shift distribution Φ as an image has been described as being stored or displayed as a phase contrast image. As described above, the phase shift distribution Φ integrates the differential amount of the phase shift distribution Φ obtained from the refraction angle φ. Thus, the differential amounts of the refraction angle φ and the phase shift distribution Φ are also related to the X-ray phase change 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.

An X-ray imaging system 70 shown in FIG. 11 is an X-ray diagnostic apparatus that images a subject H in a supine or sitting position. FIG. 11 captures an X-ray image (phase contrast image) of the knee of the subject H. An example is shown.

The X-ray imaging system 70 is held by a bed 71 on which a subject or an imaging region of the subject is placed, and an X-ray source holding device 14 suspended from the ceiling vertically above the bed 71 and directed toward the subject placed on the bed 71. An X-ray source 11 that emits X-rays, and an imaging unit 12 that is disposed opposite to the X-ray source 11 and attached to the lower surface side of the top plate 72 of the bed 71. 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, description thereof will be omitted.

In addition, although the example which image | photographs the X-ray image of the knee of the to-be-photographed subject H was shown, this X-ray imaging system 70 is used also when image | photographing the X-ray image of the hand and elbow of the to-be-photographed subject H, for example. The hand and arm of the subject H are placed on the bed 71.

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

The radiography system shown in FIG. 12 is a mammography apparatus that captures an X-ray image (phase contrast image) of the breast B. 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, description thereof will be omitted.

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

13 differs 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 mammography apparatus 80A illustrated in FIG. The first absorption type lattice 31 is accommodated in a lattice accommodation portion 85 connected to the arm member 81. The FPD 30, the second absorption grating 32, and the scanning mechanism 33 that constitute the imaging unit 12 together with the first absorption grating 31 are housed in the imaging table 83.

Thus, even when the breast B is located between the first absorption type grating 31 and the second absorption type grating 32, the first absorption current formed at the position of the second absorption type grating 32 is the first. The projected image (G1 image) of the absorption grating 31 is deformed by the breast B. Accordingly, even in this case, the moiré fringes modulated due to the breast B can be detected by the FPD 30. That is, the mammography apparatus 80A can obtain a phase contrast image of the breast B based on the principle described above.

In this mammography apparatus 80A, since the X-ray whose dose is almost halved is irradiated to the breast B due to the shielding by the first absorption grating 31, the exposure amount of the breast B is set to the above-described mammography apparatus 80. It can be reduced to about half of the case. 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 80A can be applied to any of the X-ray imaging systems described above. Is possible.

14 and 15 show examples of the arrangement of the dose detection pixels 40.

In the X-ray imaging system described above, the plurality of dose detection pixels 40 have been described as being formed as one group along one side of the detection surface of the FPD 30. However, the arrangement of the dose detection pixels 40 is, for example, As shown in FIG. 14, it may be frame-shaped (FIG. 14A) or distributed in four corners (FIG. 14B). It can be set as appropriate according to the procedure.

Preferably, the dose detection pixel 40 is provided at a position that does not overlap the subject on the detection surface of the FPD 30, and detects X-rays that propagate outside the subject. X-rays incident on the dose detection pixel 40 where the subject overlaps are attenuated by the subject. Therefore, by providing the dose detection pixel 40 at a position where the subject does not overlap on the detection surface of the FPD 30, it is possible to detect the dose more accurately. The change of the pixel value of each pixel 40 resulting from it can be removed or reduced.

As shown in FIG. 15, for example, when imaging the knee of the subject H in the X-ray imaging system 70 described above, a plurality of dose detection pixels 40 along two sides along the longitudinal direction of the leg on the detection surface of the FPD 30. (FIG. 15A), when photographing the hand of the subject H, a plurality of dose detection pixels 40 are formed in a substantially U shape along three sides excluding the side intersecting the arm on the detection surface of the FPD 30. (FIG. 15B). Further, in the mammography apparatus 80 that captures the breast B described above, a plurality of dose detection pixels 40 may be provided in a substantially U shape along three sides excluding the side along the chest wall of the subject H on the detection surface of the FPD 30. (FIG. 15C).

Also, the pixel value of the dose detection pixel 40 where the subject overlaps is smaller than the pixel value of the dose detection pixel 40 where the subject does not overlap because the X-rays incident thereon are attenuated by the subject. Therefore, based on the pixel value of the dose detection pixel 40, it may be determined whether each dose detection pixel 40 is overlapped with the subject. For example, a predetermined threshold value is provided for the pixel value, and the pixel value of the dose detection pixel 40 is compared with the threshold value so that the dose detection pixel 40 having a pixel value less than the threshold value is determined as the dose detection pixel 40 on which the subject overlaps. If the arithmetic processing unit 22 is configured so that the pixel value of the dose detection pixel 40 with which the subject overlaps is not used in the luminance correction described above, the irradiation dose during imaging can be more reliably as in the example shown in FIG. It is possible to remove or reduce the change in the pixel value of each pixel 40 due to the variation in the number of pixels. According to such a configuration, it is not necessary to change the setting of the arrangement of the dose detection pixels 40 according to the subject in order to avoid the overlap between the dose detection pixels 40 and the subject.

FIG. 16 shows a configuration of an imaging unit related to an example of a radiation imaging system for explaining an embodiment of the present invention.

In the X-ray imaging system 10 described above, the dose detection pixel 40 of the FPD 30 has been described as being incident with X-rays propagating off both the grating regions of the first and

second absorption gratings

31 and 32. In order to accurately detect the dose in the dose detection pixel 40, it is sufficient that the moire fringes are not formed on the dose detection pixel 40, so that the first absorption type lattice 31 is out of the lattice region, and the second absorption type lattice 32 is formed. X-rays propagating through the grating region can also be configured to enter the dose detection pixel 40, and the grating region of the first absorption grating 31 can be separated from the grating region of the second absorption grating 32. It is also possible to configure so that X-rays that pass through and enter the dose detection pixel 40. In the X-ray imaging system 60 shown in FIG. 16, X-rays that are out of the lattice area of the first absorption grating 31 and propagate through the grating area of the second absorption grating 32 are incident on the dose detection pixel 40. It is configured to Since other configurations and operations are the same as those of the X-ray imaging system 10 described above, description thereof will be omitted.

In the above configuration, the X-rays incident on the dose detection pixel 40 are attenuated by passing through the second absorption grating 32. This is preferable because the pixel value of the dose detection pixel 40 can be prevented from being saturated. That is, the X-rays incident on the image detection pixel 40 of the FPD 30 are attenuated by passing through the first and

second absorption gratings

31 and 32, and a sufficient dose is secured in the image detection pixel 40 as a reference. Shooting conditions are defined. Therefore, in the case where X-rays propagating out of the lattice regions of the first and second

absorption type gratings

31 and 32 are incident on the dose detection pixel 40, the pixel of the dose detection pixel 40 depends on the imaging conditions. It is also possible that the value is saturated. Therefore, X-rays attenuated by passing through one grating region of the first and second absorption gratings 31 and 32 (in this example, the grating region of the second absorption grating 32) are dose detection pixels 40. In this case, the pixel value of the dose detection pixel 40 can be prevented from being saturated. Furthermore, according to the above configuration, one of the first and second

absorption type gratings

31 and 32 located in the traveling path of the X-rays incident on the dose detection pixel 40 also serves as a scattering removal grid. I can do it. Therefore, it is possible to exclude the influence of scattering caused by a subject or the like, and to perform dose detection more accurately.

FIG. 17 shows a configuration of an imaging unit regarding a modification of the X-ray imaging system 60 described above.

The X-ray imaging system 60A shown in FIG. 17 is configured such that X-rays propagating off both the grating regions of the first and

second absorption gratings

31 and 32 are incident on the dose detection pixel 40 of the FPD 30. On the other hand, in order to prevent the pixel value of the dose detection pixel 40 from being saturated, an X-ray attenuator 61 overlapping the dose detection pixel 40 is provided.

The X-ray attenuator 61 is formed in a foil or plate shape having a uniform thickness, and the first and second absorption gratings are projected on the detection surface of the FPD 30 with the X-ray focal point 18b as a viewpoint. 31 and 32 are provided so as to cover a region 30B that is out of projection. As a material of the X-ray attenuator 61, for example, a metal material such as gold, platinum, lead, tungsten, aluminum, copper, and iron is preferably used. The X-ray attenuator 61 is an X-ray absorptivity of the material used. Therefore, it is formed in an appropriate thickness. In addition, although it becomes thick compared with said metallic material, as a material of the X-ray attenuation body 61, nonmetallic materials, such as a polymer, a silicon | silicone, glass, can also be used.

The X-ray attenuator 61 configured as described above attenuates incident X-rays uniformly, and X-rays transmitted therethrough enter the dose detection pixels 40 belonging to the region 30B. Thereby, saturation of the pixel value of the dose detection pixel 40 can be prevented, and dose detection by the dose detection pixel 40 can be performed normally.

The X-ray attenuator 61 may be configured so that X-rays that pass through one of the first and

second absorption gratings

31 and 32 and propagate through the grating area are incident on the dose detection pixel 40. Good. According to this, since the X-rays are attenuated by passing through one of the first and second

absorption type gratings

31 and 32, the pixel value of the dose detection pixel 40 is more reliably prevented from being saturated. Or the thickness of the X-ray attenuator 61 can be kept relatively small.

Further, the dose detection pixel can be obtained by transmitting the X-ray attenuator 61 or transmitting the X-ray attenuator 61 and passing through one of the first and

second absorption gratings

31 and 32. When the pixel value of 40 is saturated, if it is due to the dynamic range of the readout circuit 43 (see FIG. 3) of the FPD 30, by using an integration amplifier circuit with a smaller amplification factor in the readout circuit 43, It is possible to prevent the pixel value of the dose detection pixel 40 from being saturated. For example, several types of integration amplifier circuits having different amplification factors may be provided in the readout circuit 43, and an integration amplifier circuit having an appropriate amplification factor may be selectively used according to the photographing conditions.

In this example, the X-ray attenuator 61 is placed on the detection surface of the FPD 30. As described above, the X-ray attenuator 61 is preferably arranged so that the transmission surface of the transmitted X-ray is in close contact with or very close to the detection surface of the FPD 30. Although scattering occurs due to transmission through the X-ray attenuator 61, the smaller the distance between the exit surface of the X-ray attenuator 61 and the detection surface of the FPD 30, the more the X-ray dissipation can be reduced. More accurate dose detection is possible. When the X-ray propagating through one of the first and second absorption-

type gratings

31 and 32 is incident on the dose detection pixel 40, the X-ray attenuator 61 is connected to one of the X-ray attenuators 61. You may arrange | position to the upstream (X-ray source 11 side) rather than a grating | lattice. One of the first and second absorption-

type gratings

31 and 32 positioned in the traveling path of the X-rays incident on the dose detection pixel 40 can also serve as a scatter removal grid. Is disposed upstream of one of the gratings, the influence of scattering by the X-ray attenuator 61 can be eliminated, and accurate dose detection can be performed.

18 and 19 show the configuration of the imaging unit in relation to the modification of the radiation imaging system of FIG.

In the example shown in FIG. 18, the thickness of the X-ray attenuator 61 is different in each part, and in the example shown in the figure, the thickness increases or decreases stepwise in the width direction. According to such a configuration, for example, even if the pixel value is saturated in the dose detection pixel 40 where the portion 61a having the smallest thickness overlaps in the X-ray attenuator 61, the pixel in the dose detection pixel 40 where the

portions

61b and 61c having larger thickness overlap. The value is not saturated and dose detection can be performed normally. Thereby, it is possible to cope with more imaging conditions using the single X-ray attenuator 61.

An X-ray attenuator can also be configured by arranging a plurality of attenuation materials (for example, platinum, gold, lead, silver, tungsten, molybdenum, etc.) having the same thickness but different attenuation coefficients in a direction perpendicular to the thickness direction. . In this case, the thickness of the X-ray attenuator can be made uniform, and the attenuation can be changed in each part.

Further, as shown in FIG. 19, first and second damping

materials

62 and 63 having different damping coefficients are laminated in the thickness direction, and the ratio of the thicknesses of the first and second damping

materials

62 and 63 is set for each part. The X-ray attenuator 61 can be configured differently. Also in this case, the thickness of the X-ray attenuator can be made uniform and the attenuation can be changed in each part.

In the brightness correction of image data, only the dose detection pixel 40 whose pixel value is unsaturated is extracted from the dose detection pixels 40, and the brightness correction of each image data is performed based on the extracted pixel value of the dose detection pixel 40. Made.

Here, depending on the characteristics of the FPD, the output of each pixel with respect to the incident dose may not be linear. For example, as shown in FIG. 20, the slope of output characteristics may differ between a high dose and a low dose. In such a case, the average pixel value of the image detection pixels 40 that detect moire fringes or the pixel value closest to the pixel value of the specific image detection pixel 40 among the dose detection pixels 40 whose pixel values are unsaturated. It is preferable to perform luminance correction of each image data based on the pixel value of the dose detection pixel 40.

As described above, the configuration for preventing saturation of the pixel value of the dose detection pixel 40 described with reference to FIGS. 16 to 20 can be applied to any of the X-ray imaging systems described above.

FIG. 21 shows an example of a radiation imaging system for explaining the embodiment of the present invention, and FIG. 22 shows a configuration of an imaging unit of the radiation imaging system of FIG.

According to each of the X-ray imaging systems 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 referenced 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.

An X-ray imaging system 90 shown in FIG. 21 includes a phase imaging mode in which the first and

second absorption gratings

31 and 32 are arranged in the X-ray irradiation field, and a phase contrast image of the subject H is generated by the above-described fringe scanning. The first and second

absorption type gratings

31 and 32 are retracted from the X-ray irradiation field, and a normal imaging mode for generating an image (absorption image) based on an X-ray intensity change by the subject H is provided. Further, a moving mechanism 91 for retracting the first and second

absorption type gratings

31 and 32 from the X-ray irradiation field is further provided. Switching between the phase imaging mode and the normal imaging mode, that is, insertion of the first and

second absorption gratings

31 and 32 into the X-ray irradiation field and withdrawal from the irradiation field, for example, according to an input operation on the console 13 This is done by the control device 22 driving the moving mechanism 91. As the moving mechanism 91, for example, a linear motion mechanism such as a ball screw or a linear motor can be used.

Further, the X-ray imaging system 90 is provided with an X-ray attenuator 61 for preventing saturation of the pixel value of the dose detection pixel 40 of the FPD 30 as in the X-ray imaging system 60A shown in FIG. . In the X-ray imaging system 60A described above, the X-ray attenuator 61 is placed on the detection surface of the FPD 30, but in the X-ray imaging system 90, the X-ray attenuator 61 is the second absorption type grating. 32 is integrated.

The generation process of the absorption image in the normal imaging mode is different from the above-described generation process of the phase contrast image by the fringe scanning, and only needs to be performed once. For this reason, it is not necessary to measure the variation in the irradiation dose between radiographs, and the X-ray attenuator 61 is also unnecessary. Therefore, in the present X-ray imaging system 90, the X-ray attenuator 61 and the first

absorption type gratings

31 and 32 are also retracted from the X-ray irradiation field. Thereby, the entire detection surface of the FPD 30 can be effectively utilized.

Here, a separate moving mechanism may be provided to insert and retract the X-ray attenuator 61 from the X-ray irradiation field. However, in the X-ray imaging system 90, the X-ray attenuator 61 is provided. Is integrated with the second absorption type grating 32, and is moved together with the second absorption type grating 32 by the moving mechanism 91. Thereby, the configuration of the apparatus can be simplified.

Further, in the present X-ray imaging system 90, the X-ray attenuator 61 formed separately from the second absorption type grating 32 is assembled to the second absorption type grating 32, and both are integrated. It is also possible to form the attenuating body 61 in the second absorption type grating 32 so that both are integrated. Furthermore, if the same material (gold, platinum, etc.) as the X-ray shielding part 32b of the second absorption type grating 32 is used as the material of the X-ray attenuator 61, the X-ray shielding part can be obtained by metal plating or vapor deposition. It can be formed simultaneously with 32b. Instead of the second absorption type grating 32, an X-ray attenuation body 61 is provided integrally with the first absorption type grating 31, and the X-ray attenuation body 61 is moved together with the first absorption type grating 31. May be.

FIG. 23 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 set to 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 deteriorated. 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.

Expression (18) indicates that the projection image (G1 image) of the X-rays emitted from the small-focus light sources dispersedly formed by the multi-slit 103 by the first absorption-type grating 31 is the 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 100, the G1 images based on the plurality of small focus light sources formed by the multi slit 103 are superimposed, thereby improving the image quality of the phase contrast image without decreasing the X-ray intensity. Can be improved. The multi slit 103 described above can be applied to any of the X-ray imaging systems described above. When applied to the X-ray imaging system 90 described above, in normal imaging performed by retracting the first and

second absorption gratings

31 and 32 from X-ray irradiation, the multi-slit 103 is also used in the X-ray irradiation field. Evacuate from.

FIG. 24 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

In the X-ray imaging system 90 described above, an absorption image is captured separately from the phase contrast image. However, it is difficult to satisfactorily superimpose due to a shift in the imaging position between the phase contrast image capture and the absorption image capture. There is a case. Moreover, it may be a burden on the subject due to an increase in the number of imaging. 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 I _k (x, y) obtained for each pixel with respect to k and calculating an average value as illustrated 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 an object in which an absorption image is created from a group of images acquired by photographing (main shooting) in the presence of the object, 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 scatter image is created from a group of images acquired by shooting (main shooting) in the presence of a subject, and the amplitude value unevenness of the detection system is corrected by applying the above correction coefficient 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. 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.

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

As described above, the present specification includes a first grating and a second grating having a period that substantially matches the pattern period of the radiation image formed by the radiation that has passed through the first grating. And a radiation image detector that detects the radiation image masked by the second grating, wherein the radiation image detector has at least one grating region of the first grating and the second grating. A radiation image detection apparatus is disclosed that includes at least one dose detection pixel that is used to detect the amount of radiation that is incident on the radiation that propagates off the beam.

The radiological image detection apparatus disclosed in the present specification further includes a radiation attenuator that overlaps each of the dose detection pixels.

Also, the radiation image detection apparatus disclosed in this specification is provided with a plurality of the dose detection pixels, and the radiation attenuator has different attenuation amounts in each part.

Further, in the radiological image detection apparatus disclosed in this specification, the thickness of the radiation attenuator is different in each part.

In the radiological image detection apparatus disclosed in this specification, the radiation attenuator includes first and second radiation attenuating materials having different attenuation coefficients, and the first and second radiation attenuating materials have a thickness. The thickness ratios of the first and second radiation attenuating materials are different in each part of the radiation attenuating body.

Further, in the radiological image detection apparatus disclosed in this specification, the radiation attenuator includes a plurality of radiation attenuating materials having different attenuation coefficients, and the plurality of radiation attenuating materials are arranged in a direction orthogonal to the thickness direction. Configured.

Also, in the radiological image detection apparatus disclosed in the present specification, the radiation attenuator is disposed in close contact with the detection surface of the radiological image detector.

The radiological image detection apparatus disclosed in the present specification further includes a moving mechanism for retracting the first grating, the second grating, and the radiation attenuator from the radiation irradiation field.

Also, in the radiological image detection apparatus disclosed in this specification, the radiation attenuator is provided integrally with the first grating or the second grating.

Further, the present specification discloses a radiation imaging apparatus including the above-described radiation image detection apparatus and a radiation source that emits radiation toward the first grating.

Further, the present specification includes the radiation image detection apparatus, a radiation source that emits radiation toward the first grating, and an arithmetic processing unit that processes image data acquired by the radiation image detector. A subject is disposed between the radiation source and the first grating, or between the first grating and the second grating, and the second grating is positioned with respect to the radiation image. There is an imaging mode in which imaging is performed a plurality of times at relative positions different from each other in phase, and the arithmetic processing unit is configured to acquire the image data acquired by the radiation image detector in each imaging in the imaging mode in the imaging. A radiation imaging system that corrects brightness based on a dose detected by a dose detection pixel is disclosed.

Further, the present specification includes the radiation image detection apparatus, a radiation source that emits radiation toward the first grating, and an arithmetic processing unit that processes image data acquired by the radiation image detector. A subject is disposed between the radiation source and the first grating, or between the first grating and the second grating, and the second grating is positioned with respect to the radiation image. A first imaging mode in which imaging is performed a plurality of times at relative positions different from each other in phase; and the first grating, the second grating, and the radiation attenuator are retracted from a radiation irradiation field; There is a second imaging mode in which a subject is placed between the radiographic image detector and imaging is performed, and the arithmetic processing unit is acquired by the radiographic image detector at each imaging in the first imaging mode. Processed image data Radiation imaging system for brightness correction on the basis of the dose detected by said dose detection pixel are disclosed in shooting.

Also, in the radiation imaging system disclosed in this specification, the arithmetic processing unit corrects the luminance based on the dose detected by the dose detection pixel on which the radiation propagating off the subject is incident.

Further, the radiation imaging system disclosed in the present specification includes a plurality of the dose detection pixels, and the arithmetic processing unit is detected by a pixel whose pixel value is unsaturated among the plurality of dose detection pixels. Brightness correction based on the dose.

Further, in the radiation imaging system disclosed in this specification, the calculation processing unit is a dose detection pixel whose pixel value is closest to a pixel value of a pixel group that detects the radiation image among the plurality of dose detection pixels. The luminance is corrected based on the dose detected by.

Further, in the radiation imaging system disclosed in the present specification, the calculation processing unit calculates a distribution of refraction angles of radiation incident on the radiation image detector from a plurality of image data whose luminance has been corrected. A phase contrast image is generated based on the angular 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 November 26, 2010 (Japanese Patent Application No. 2010-264241), the contents of which are incorporated herein by reference.

10 X-ray imaging system 11 X-ray source 12 Imaging unit (radiation image detection apparatus)
13 Console 20 Control device 30 FPD
31 First absorption type grating 32 Second absorption type grating 33 Scanning mechanism 40 Pixel

Claims

A first lattice;
A second grating having a period that substantially matches the pattern period of the radiation image formed by the radiation that has passed through the first grating;
A radiation image detector for detecting the radiation image masked by the second grating;
With
The radiation image detector includes at least a dose detection pixel that is used to detect a radiation dose incident on the radiation that propagates out of at least one of the first grating and the second grating. Radiation image detection device including one.
The radiological image detection apparatus according to claim 1,
A radiation image detection apparatus further comprising a radiation attenuator that overlaps each of the dose detection pixels.
The radiological image detection apparatus according to claim 2,
A plurality of the dose detection pixels are provided,
The radiation attenuator is a radiological image detection apparatus in which attenuation is different in each part.
The radiological image detection apparatus according to claim 3,
The radiation attenuating body is a radiation image detecting device in which the thickness is different in each part.
The radiological image detection apparatus according to claim 3,
The radiation attenuating body includes first and second radiation attenuating materials having different attenuation coefficients, and the first and second radiation attenuating materials are laminated in the thickness direction.
The radiation image detection apparatus in which the ratio of the thickness of the first and second radiation attenuating materials is different in each part of the radiation attenuating body.
The radiological image detection apparatus according to claim 3,
The radiation attenuator includes a plurality of radiation attenuating materials having different attenuation coefficients, and the plurality of radiation attenuating materials are arranged in a direction perpendicular to the thickness direction.
The radiological image detection apparatus according to any one of claims 2 to 6,
The radiation image detecting device, wherein the radiation attenuator is disposed in close contact with a detection surface of the radiation image detector.
The radiological image detection apparatus according to any one of claims 2 to 7,
A radiation image detection apparatus further comprising a moving mechanism for retracting the first grating, the second grating, and the radiation attenuator from a radiation irradiation field.
The radiological image detection apparatus according to claim 8,
The radiation image detecting apparatus, wherein the radiation attenuator is provided integrally with the first grating or the second grating.
The radiological image detection apparatus according to any one of claims 1 to 9,
A radiation source that emits radiation toward the first grating;
A radiographic apparatus comprising:
The radiological image detection apparatus according to any one of claims 1 to 9,
A radiation source that emits radiation toward the first grating;
An arithmetic processing unit for processing image data acquired by the radiation image detector;
With
An object is disposed between the radiation source and the first grating, or between the first grating and the second grating, and the second grating is different in phase from the radiation image. There is a shooting mode that takes multiple shots at a relative position,
The radiation processing system, wherein the arithmetic processing unit corrects the brightness of image data acquired by the radiation image detector in each imaging in the imaging mode based on a dose detected by the dose detection pixel in the imaging.
A radiological image detection apparatus according to claim 8 or 9,
A radiation source that emits radiation toward the first grating;
An arithmetic processing unit for processing image data acquired by the radiation image detector;
With
An object is disposed between the radiation source and the first grating, or between the first grating and the second grating, and the second grating is different in phase from the radiation image. A first shooting mode in which a plurality of shootings are performed at a relative position;
A second imaging mode in which the first grating, the second grating, and the radiation attenuator are retracted from a radiation field, and an object is placed between the radiation source and the radiation image detector to perform imaging. When,
There is
The arithmetic processing unit is a radiographic system that corrects the brightness of image data acquired by the radiographic image detector in each imaging in the first imaging mode based on a dose detected by the dose detection pixels in the imaging. .
The radiographic system according to claim 11 or 12,
The radiation processing system in which the arithmetic processing unit performs brightness correction based on a dose detected by a dose detection pixel on which radiation propagating off the subject enters.
The radiation imaging system according to any one of claims 11 to 13,
A plurality of the dose detection pixels are provided,
The said arithmetic processing part is a radiography system which carries out brightness correction | amendment based on the dose detected by the pixel whose pixel value is unsaturated among these dose detection pixels.
The radiation imaging system according to claim 14,
The arithmetic processing unit is a radiography system that corrects luminance based on a dose detected by a dose detection pixel whose pixel value is closest to a pixel value of a pixel group that detects the radiation image among the plurality of dose detection pixels. .
The radiation imaging system according to any one of claims 11 to 15,
The arithmetic processing unit calculates a refraction angle distribution of radiation incident on the radiation image detector from a plurality of brightness corrected image data, and generates a phase contrast image based on the refraction angle distribution. Shooting system.