WO2013047011A1

WO2013047011A1 - Radiographic image detector, method of manufacturing same, and radiography system employing radiographic image detector

Info

Publication number: WO2013047011A1
Application number: PCT/JP2012/071074
Authority: WO
Inventors: 岩切　直人
Original assignee: 富士フイルム株式会社
Priority date: 2011-09-30
Filing date: 2012-08-21
Publication date: 2013-04-04
Also published as: JP2014238265A

Abstract

A radiographic image detector comprises: a lattice part which forms a radiographic image by radiation passing therethrough, having a periodic intensity distribution; and a detector part in which multiple pixels which detect the radiation and accumulate an electric charge are arrayed in a matrix, said detector part acquiring the radiographic image including a periodic pattern based on the periodic intensity distribution. The lattice part is anchored to the detector part.

Description

Radiation image detector, method of manufacturing the same, and radiation imaging system using the radiation image detector

The present invention relates to a radiation image detector, a manufacturing method thereof, and a radiation imaging system using the radiation image detector.

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

However, X-ray absorptivity becomes lower as a substance composed of an element having a smaller atomic number, and a difference in X-ray absorptivity is small in living soft tissue or soft material. Therefore, there is a problem that a sufficient contrast (contrast) of the X-ray transmission image 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 Talbot interferometer comprising an X-ray image detector and two separate transmission diffraction gratings (phase grating and absorption grating) has been developed. An X-ray imaging system used has been devised (see, for example, Patent Document 1).

In the X-ray Talbot interferometer, a first diffraction grating (phase type grating or absorption type grating) is arranged behind the subject, and a specific distance (Talbot interference 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 the X-rays that have passed through the first diffraction grating form a self-image (hereinafter referred to as a G1 image) that is a periodic pattern due to the Talbot interference effect. It is modulated by the interaction (phase change) between the subject and the X-rays arranged between the source and the first diffraction grating.

The X-ray Talbot interferometer detects the moiré fringes generated by superimposing the G1 image and the second diffraction grating, and obtains the phase information of the subject by analyzing the modulation of the moire fringes by the subject. 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. X-rays refracted by the subject from a change in signal value for each corresponding pixel between a plurality of image data obtained by performing a plurality of times of imaging while translating in a vertical direction with a scanning pitch obtained by equally dividing the lattice pitch. Angle distribution (differential image of phase shift) can be obtained, and a phase contrast image of the subject can be obtained based on this angle distribution.

However, according to the above-described fringe scanning method, it is necessary to perform photographing a plurality of times, and there is a concern about the movement of the subject during photographing and the resulting deterioration in image quality. In view of this, a method has been proposed in which the phase information of the subject is acquired by one shooting by using Fourier transform and inverse Fourier transform (see, for example, Patent Document 2). In this method, the frequency domain including the fundamental frequency component of the moire fringe is separated from the spatial frequency spectrum obtained by Fourier transforming the moire fringe, and the phase shift is differentiated by performing the inverse Fourier transform on the separated frequency domain. Get a statue. This eliminates the need for a moving mechanism that moves the lattice with high accuracy during a plurality of shootings, thereby improving the shooting workflow and simplifying the apparatus. In addition, it is possible to eliminate the deterioration in image quality caused by the movement of the subject between each photographing. However, when the fundamental frequency component of the moire fringe is separated, a high frequency component equal to or higher than the fundamental frequency band of the moire fringe or the fundamental frequency component is cut by a filter or the like. For this reason, the differential image of the phase shift reconstructed by the inverse Fourier transform has no cut frequency component, so that the resolution is inferior to the fringe scanning method.

Further, by detecting the periodic pattern of the G1 image using a detector having a pixel pitch smaller than the pitch of the periodic pattern of the G1 image without using the second diffraction grating, and analyzing the modulation of the periodic pattern, An X-ray imaging system that acquires phase information of an object has also been proposed (see Patent Document 3).

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

In the X-ray imaging systems described in Patent Document 1 to Patent Document 3, a first diffraction grating is provided separately from the X-ray image detector. For this reason, for example, in order to actually use the X-ray imaging system described in Patent Document 3 for imaging a subject, first, a first diffraction grating for generating a G1 image and an X pattern for detecting a periodic pattern of the G1 image are detected. The relative alignment of the line image detector must be performed with high accuracy. However, the periodic pattern of G1 is as small as about several μm, and highly accurate position adjustment is complicated.

Even if the subject can actually be photographed, the vibrations transmitted to the X-ray imaging system and the difference in the thermal expansion coefficient between the first diffraction grating and the detector as the number of times of photographing is increased. In some cases, the positional relationship between the first diffraction grating and the X-ray image detector once combined may be changed. For this reason, it is necessary to periodically align the first diffraction grating and the X-ray image detector.

In X-ray phase imaging, phase information of a subject is acquired by analyzing a minute change of modulation of the periodic pattern of the G1 image that occurs when there is no subject and when there is a subject. For this reason, if the relative positional relationship between the first diffraction grating and the X-ray image detector deviates between the shooting when there is no subject and the shooting when there is a subject, an appropriate periodic pattern modulation is performed. Cannot be obtained, greatly affecting the accuracy of the phase information.

The present invention has been made in view of the above-described circumstances, and reduces the positional deviation of the radiation image detector exemplified by the first diffraction grating and the X-ray image detector, and improves the accuracy of the obtained phase information of the subject. The purpose is to increase.

A lattice unit that forms a radiation image including a periodic intensity distribution by passing radiation and a plurality of pixels that detect the radiation and accumulate charges are arranged in a matrix and include a periodic pattern based on the periodic intensity distribution A detection unit that acquires the radiation image, and the lattice unit is fixed to the detection unit, and the lattice unit is formed of a support that transmits the radiation and a shielding member that absorbs the radiation. The support is fixed to the detection unit, and the detection unit is supported from the radiation incident side, and the shielding member is a radiation image detector formed in the support.

According to the present invention, since the shielding member functioning as a grating is integrally incorporated in the radiation image detector itself, it is not necessary to align it many times. Further, since the positional deviation between the shielding member and the detection unit is reduced, the accuracy of the obtained subject phase information can be increased. Further, since a mechanism for adjusting the position of the shielding member is not required, the configuration of the radiation imaging system is simplified.

It is a schematic diagram which shows the structure of an example of the radiography system for describing embodiment of this invention. It is a control block diagram of the radiography system of FIG. It is a perspective view 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 structure of the radiographic image detector of the radiography system of FIG. It is a schematic diagram for demonstrating the refraction | bending of the radiation by a to-be-photographed object. It is a schematic diagram which shows an example of the analysis method of periodic intensity distribution in the radiography system of FIG. It is a schematic diagram which shows an example of the manufacturing method of the radiographic image detector of FIG. It is a schematic diagram which shows an example of the manufacturing method of the radiographic image detector of FIG. It is a schematic diagram which shows the structure of the other example of the radiographic image detector of the radiography system of FIG. It is a schematic diagram which shows an example of the analysis method of a moire using the radiographic image detector of FIG. It is a schematic diagram which shows the other example of the analysis method of a moire using the radiographic image detector of FIG. It is a schematic diagram which shows the manufacturing method of the radiographic image detector of FIG. It is a schematic diagram which shows the further another example of the radiographic image detector of the radiography system of FIG. It is a schematic diagram which shows an example of the manufacturing method of the radiographic image detector of FIG. It is a schematic diagram which shows an example of the manufacturing method of the radiographic image detector 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 schematic diagram which shows the further another example of the radiographic image detector 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 generates an image data by detecting an X-ray source 11 that emits X-rays to the subject H and an X-ray source 11 that is disposed opposite to the X-ray source 11 and transmits the subject H from the X-ray source 11. The imaging unit 12 that controls the exposure operation of the X-ray source 11 and the imaging operation of the imaging unit 12 based on the operation of the operator, and the image data acquired by the imaging unit 12 is arithmetically processed to obtain a phase contrast image And the console 13 for generating

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.

The X-ray source 11 includes an X-ray source control unit 17, an X-ray tube 18 that generates X-rays according to a high voltage applied from the high voltage generator 16 based on the control of the X-ray source control unit 17, A collimator unit 19 having a movable collimator 19a that limits the irradiation field so as to shield a portion of the X-rays emitted from the X-ray tube 18 that does not contribute to the inspection region of the subject H. Yes. The X-ray tube 18 is an anode rotating type, and emits an electron beam from a filament (not shown) as an electron emission source (cathode) and collides it with a rotating anode 18a rotating at a predetermined speed. X-rays are 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. And have. 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.

In the standing stand 15, a holding part 15b for holding the photographing part 12 is attached to a main body 15a installed on the floor so as 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.

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 amount of rotation of the pulley 15c or the amount of movement of the endless belt 15d. It has been. 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, and the like. 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 an X-ray image detector 30 that is integrally provided with an absorption grating unit 31 for detecting phase change (angle change) of X-rays by the subject H and performing phase imaging. .

3 and 4 schematically show the configuration of the X-ray image detector 30. FIG.

The X-ray image detector 30 includes an absorption lattice portion 31 that forms an X-ray image including a periodic intensity distribution by passing X-rays, and a plurality of pixels 40 that detect X-rays and accumulate charges in a matrix. And a sensor unit 41 that acquires an X-ray image including a periodic pattern based on the periodic intensity distribution. In the X-ray image detector 30, the detection surface of the sensor unit 41 is disposed so as to be orthogonal to the optical axis A of X-rays emitted from the X-ray source 11.

The absorption type lattice part 31 is fixed to the sensor part 41. Further, the absorption type grating unit 31 is provided on the X-ray incident side with respect to the sensor unit 41, and is provided between the sensor unit 41 and the X-ray source 11.

The absorptive lattice portion 31 is formed of a substrate portion 31a that transmits X-rays and an X-ray shielding member 31b that absorbs X-rays.

The substrate unit 31a is fixed to the sensor unit 41 and is formed of an X-ray transmissive material (low radiation absorption material) such as a silicon substrate that transmits X-rays. The substrate unit 31a supports the sensor unit 41 from the X-ray incident side.

The X-ray shielding member 31b is formed in the substrate portion 31a, and is in one direction in the plane orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11 (in the illustrated example, the x direction and the z direction). It is comprised by the linear member extended | stretched in the y direction orthogonal to. As a material of the X-ray shielding member 31b, a material excellent in X-ray absorption (radiation high absorption material) is preferable, and for example, a heavy metal such as gold or platinum is preferable. Note that the X-ray shielding member 31b can be formed by the metal plating method or the vapor deposition method using the above-described material. A method for manufacturing the X-ray shielding member 31b will be described later.

X-ray shielding member 31b is in a plane perpendicular to the optical axis A of the X-ray, at a pitch p ₁ constant in the direction (x-direction) orthogonal to the one direction, parallel at a predetermined distance d ₁ from each other It is arranged. Since the X-ray shielding member 31b does not give a phase difference to incident X-rays but gives an intensity difference to incident X-rays, it is also called an amplitude type grating.

The X-ray shielding member 31b is configured to geometrically project X-rays that have passed through the slit portion regardless of the presence or absence of the Talbot interference effect. Specifically, by setting the interval d ₁ to a value sufficiently larger than the effective wavelength of the X-rays emitted from the X-ray source 11, most of the X-rays included in the irradiated X-rays are not diffracted by the slit portion. In addition, it is configured to pass while maintaining straightness. For example, when tungsten is used as the rotary anode 18a described above and the tube voltage is 50 kV, the peak wavelength of the X-ray is about 0.4 mm. In this case, if the distance d ₁ is about 1 to 10 μm, most of the X-rays are geometrically projected without being diffracted by the slit portion.

The X-ray emitted from the X-ray source 11 is not a parallel beam but a cone beam with the X-ray focal point 18b as a light emitting point, and therefore a projection image projected through the X-ray shielding member 31b (hereinafter referred to as this 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.

Distance L ₂ from X-ray shield member 31b to the sensor unit 41, the Talbot interferometer, but is constrained to Talbot distance determined by the grating pitch and the X-ray wavelength of the diffraction grating, in the X-ray imaging system 10 Since the X-ray shielding member 31b projects incident X-rays without diffracting, the G1 image of the X-ray shielding member 31b is obtained similarly at all positions behind the X-ray shielding member 31b. , the distance L _2, can be set independently of the Talbot distance. Accordingly, the distance L _2, is set to be smaller than the Talbot interference distance, the X-ray image detector 30 can be made thinner.

The X-ray shielding member 31b preferably completely shields (absorbs) X-rays in order to generate a periodic pattern image with high contrast. However, the X-ray shielding member 31b is excellent in X-ray absorption (such as gold and platinum). Even if is used, 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 thickness h ₁ of the X-ray shielding member 31b, it is preferable to be thick 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 thickness h ₁ is 30 μm or more in terms of gold (Au). Is preferred.

On the other hand, if too thick a thickness h ₁ of the X-ray shielding member 31b, X-rays incident obliquely is less likely to pass through the slit portion, so-called vignetting occurs, the extending direction (strip direction of the X-ray shielding member 31b There is a problem that the effective field of view in the direction (x direction) orthogonal to () becomes narrow. Therefore, in view of the field of view secured to define the upper limit of the thickness h _1. In order to secure the effective field length V in the x direction on the detection surface of the sensor unit 41 of the X-ray image detector 30, let L be the distance from the X-ray focal point 18 b to the detection surface of the X-ray image detector 30. The thickness h ₁ needs to be set so as to satisfy the following expression (1) from the geometrical relationship shown in FIG.

For example, when d ₁ = 2.5 μm, assuming imaging in a normal hospital, and L = 2 m, in order to ensure a length of 10 cm as the effective field length V in the x direction, the thickness _{h 1} may be set to 100μm or less.

The aperture ratio of the X-ray shielding member 31b is preferably 0.2 to 0.6 (= 20 to 60%), particularly 0.3 to 0.5 (= 30 to 50%). Further, if the pitch p ₁ of the X-ray shielding member 31b is the phase angle resolution θ [rad] to be detected and the distance L _H (mm) from the X-ray shielding member 31b to the subject, the following equation (2) is satisfied. Is set as follows.

For example, if the phase angle resolution θ to be detected is 0.1 μrad, and the distance between the pixel 40 and the subject is L _H = 10 mm, the pitch p ₁ of the X-ray shielding member 31b may be 3.28 μm or less.

In the above configuration, the G1 image of the X-ray shielding member 31b is captured by the sensor unit 41. Next, the configuration of the X-ray image detector 30 will be described.

FIG. 5 schematically shows the configuration of the X-ray image detector 30. FIG. 5A is a schematic diagram of a front view of the X-ray image detector 30, and FIG. 5B is shown in FIG. 5A is a schematic diagram of a cross-sectional view taken along line vv of 5A. FIG.

The X-ray image detector 30 includes the above-described absorption-type grating unit 31, a sensor unit 41 in which a plurality of pixels 40 that detect X-rays and accumulate electric charges are two-dimensionally arranged in the xy direction, and accumulate in each pixel 40. A scanning circuit 42 for controlling the read timing of the read charges, a signal processing circuit 43 for converting and storing signals sequentially read from the respective pixels 40 into image data, and the image data via the I / F 25 of the console 13. And a data transmission circuit 44 for transmitting to the arithmetic processing unit 22. (FIG. 5A).

The X-ray image detector 30 can be configured based on a solid-state imaging device such as a CMOS (Complementary Metal Oxide Semiconductor) sensor. The sensor unit 41 includes a semiconductor substrate 48 such as a silicon substrate, a plurality of pixels 40 such as photodiodes formed on the semiconductor substrate 48, and a plurality of readout circuits that read out charges accumulated in each pixel 40 (FIG. (Not shown), a wiring unit 47 for connecting the sensor unit 41 to the scanning circuit 42 and the signal processing circuit 43, and a scintillator 49 that emits fluorescence having a wavelength suitable for the spectral sensitivity of the pixel 40 by X-ray exposure. (FIG. 5B).

The wiring portion 47 is formed on a semiconductor substrate made of single crystal silicon or the like, and includes a plurality of scanning lines 45 and a plurality of signal lines 46 provided in a matrix for reading out the electric charges accumulated in the pixels 40. The wiring part 47 is provided between the substrate part 31a of the absorption lattice part 31 and the pixel 40 in the z direction (FIG. 5B).

The scanning line 45 is formed above the region between the pixel rows of the pixels 40 that are two-dimensionally arranged. That is, the scanning line 45 does not overlap the pixel 40 in the thickness direction (z direction) of the sensor unit 41. In other words, the scanning line 45 is formed between the adjacent pixels 40 in a plan view from the X-ray incident side. Further, in the thickness direction of the sensor unit 41 (thickness direction of the X-ray shielding member 31b) (z direction), the X-ray shielding member 31b extends over the entire row direction (y direction) of some scanning lines 45 and pixels 40. overlapping. Although the X-ray shielding member 31b does not overlap all the scanning lines 45, the X-ray shielding member 31b periodically overlaps the scanning lines 45 in a plan view from the X-ray incident side. Yes. In the illustrated example, in the arrangement direction (x direction) of the X-ray shielding members 31b, the distance d between the X-ray shielding members 31b so that at least two pixels 40 are included between the adjacent X-ray shielding members 31b. ₁ is set.

Note that the scanning line 45 overlapping the X-ray shielding member 31b has an X-ray shielding member 31b whose width in the direction orthogonal to the thickness direction of the sensor unit 41 (the arrangement direction of the X-ray shielding members 31b) (x direction) corresponds. It is preferable that the width is smaller.

The signal line 46 is provided above the region between the pixel columns of the pixels 40 arranged in a two-dimensional manner. That is, the signal line 46 is formed between the adjacent pixels 40 in a plan view from the X-ray incident side.

The scintillator 49 is provided on the side opposite to the wiring portion 47 with respect to the pixel 40 (FIG. 5B). As the scintillator 49, for example, a granular scintillator such as terbium activated gadolinium oxide (Gd ₂ O ₂ S: Tb) or a columnar scintillator such as thallium activated cesium iodide (CsI: Tl) is used. When the interval between the pixels 40 is several μm or less, the scintillator is Gd ₂ SiO ₅ : Ce, Bi ₄ Ge ₃ O ₁₂ , Gd ₂ O ₂ S: Pr, Lu ₂ SiO ₅ : Ce, Lu _0.4 Gd. _1.6 A single crystal scintillator such as SiO ₅ : Ce may be used. This is because there is no reflection or scattering of light at the crystal interface unlike granular scintillators and columnar scintillators having a crystal size of about a dozen μm to 2 μm. Although the single crystal scintillator may not be configured as a single scintillator having a large area, a plurality of single crystal scintillators may be arranged in a tile shape to increase the area. Further, the light emission area of the single crystal scintillators arranged in an integer number and the total pixel area of the pixels used as image pixels among the pixels 40 may be substantially matched. It is preferable that the gap between the scintillators is located at a position different from the effective pixel. For example, a dark correction pixel area used as a dark correction pixel or the like in the wiring portion 47 or the pixel 40, an area for taking out the wiring from the IC, etc. are suitable. In addition, the sensor unit 41 may be composed of a single group of single crystal scintillators and pixels 40, or a plurality of groups of single crystal scintillators and pixels 40 are prepared and arranged. It is good.

In the X-ray image detector 30, the X-rays are incident from the absorption type grating unit 31 side, pass through the wiring unit 47 and the like, and then enter the scintillator 49. X-rays incident on the scintillator 49 are generated as fluorescence by the scintillator 49. The generated fluorescence is accumulated as charges in the pixel 40. The accumulated charges are read out based on the timing set by the scanning circuit 42 and converted into image data by the signal processing circuit 43 (FIG. 5B).

Next, a method for analyzing the G1 image will be described.

FIG. 6 shows one X-ray refracted according to the phase shift distribution Φ (x) of the subject H in the x direction.

Reference numeral 55 denotes an X-ray path that goes straight when the subject H does not exist, and the X-ray that travels along this path 55 passes through the substrate part 31 a of the absorption grating part 31 and enters the pixel 40. Reference numeral 56 indicates an X-ray path refracted and deflected by the subject H when the subject H exists. X-rays traveling along the path 56 are shielded by the X-ray shielding member 31 b of the absorption type lattice unit 31.

The phase shift distribution Φ (x) of the subject H is expressed by the following equation (3), 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 through the X-ray shielding member 31b and projected to the position of the pixel 40 is displaced in the x direction by an amount corresponding to the refraction angle φ due to refraction of the X-ray at the subject H. That is, the pattern period p ₁ ′ in the x direction of the G1 image also changes in the x direction in accordance with the change in the G1 image.

The displacement amount Δx of the G1 image is approximately expressed by the following equation (4) based on the small X-ray refraction angle φ (x).

Here, the refraction angle φ (x) is expressed by Expression (5) 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 displacement amount Δx of the G1 image is expressed by the following equation (6) based on the phase shift amount ψ of the signal output from each pixel 40 (the phase shift amount of the signal with and without the subject H). Related.

Accordingly, by obtaining the phase shift amount ψ of the signal output from each pixel 40, the refraction angle φ is obtained from the equations (4) and (6), and the phase shift distribution Φ (x) is obtained using the equation (5). Therefore, the phase shift distribution Φ (x) of the subject H, that is, the phase contrast image of the subject H can be generated by integrating this with respect to x. Hereinafter, a method of calculating the phase shift amount ψ will be described.

FIG. 7 schematically shows a signal output from each pixel 40 of the sensor unit 41.

A plurality of pixels 40 adjacent in the x direction are used as a unit, and the pixel value I of the plurality of pixels 40 constituting one unit is interpolated for each unit. In the illustrated example, the pixel values of a plurality of pixels 40 are interpolated by a sine curve, and three points need only be interpolated by the sine curve.

The phase difference between the waveforms of the signal curve (FIG. 7A) when the subject H does not exist and the signal curve (FIG. 7B) when the subject H exists corresponds to a unit pixel (one unit pixel group). This corresponds to the phase shift amount ψ of the pixel to be processed.

Since the refraction angle φ (x) is a value corresponding to the differential value of the phase shift distribution Φ (x) as shown in the equation (5), the refraction angle φ (x) is integrated along the x-axis. Thus, the phase shift distribution Φ (x) is obtained. In the above description, the y coordinate in the y direction of the pixel 40 is not taken into consideration. However, by performing the same calculation for each y coordinate, a two-dimensional phase shift distribution Φ (x , Y).

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

8 and 9 are schematic views showing an example of a method for manufacturing the X-ray image detector 30 of the X-ray imaging system 10.

First, a solid-state imaging device in which pixels 40 and wiring portions 47 are substantially formed is prepared. Next, a support substrate 51 made of an X-ray low absorption material is formed on the wiring portion 47. Then, a resist film 52 is formed on the support substrate 51.

Note that the pixels 40 formed on the semiconductor substrate 48 are not exposed to the outside at this time.

In addition, as the resist film 52, either a negative type or a positive type can be used. In this example, a case where a positive type resist film is used will be described.

Next, the resist film 52 is exposed with an electromagnetic wave having a wavelength suitable for the resist to be used from the direction of the arrow through the mask 53 (FIG. 8A). Thereby, portions other than the portion shielded by the mask 53 are cured, and etching resistance is imparted. The mask pattern formed by each mask 53 corresponds to a portion where the X-ray shielding member 31b is formed. When viewed from the direction of the arrow, the scanning line 45 overlapping the X-ray shielding member 31 b is hidden by the corresponding mask 53.

Next, after removing the mask pattern, development is performed, and the portion of the resist film 52 that is shielded from light by the mask 53 is removed, followed by washing (FIG. 8B).

Next, the semiconductor substrate 48 on which the pixels 40 are formed is polished to expose the pixels 40 to the outside (FIG. 8C).

Next, a portion of the support substrate 51 corresponding to the mask pattern is removed by etching. As a result, a plurality of grooves for filling the X-ray shielding member 31b are formed in the support substrate 51, and the substrate portion 31a of the absorption lattice portion 31 is formed (FIG. 9A).

Next, after removing the resist film 52, a plurality of grooves formed in the support substrate 51 is filled with an X-ray high absorption material by metal plating or the like. Thereby, the X-ray shielding member 31b of the absorption type lattice part 31 is formed (FIG. 9B).

Then, the pixel 40 and the scintillator 49 are optically joined. The scintillator 49 may be formed by direct vapor deposition, or may be bonded to the pixel 40 via an adhesive layer or the like (FIG. 9C).

As described above, according to the method of manufacturing the X-ray image detector 30, a plurality of grooves are formed in the support substrate 51 fixed to the sensor unit 41, and the X-ray shielding member 31b is provided by filling these grooves. Yes. For this reason, it is possible to manufacture the X-ray image detector 30 with higher accuracy than separately preparing the sensor unit 41 and the absorption type grating unit 31 and aligning the pixel 40 and the X-ray shielding member 31b later.

Further, since the semiconductor substrate 48 on which the pixels 40 are formed is polished after the support substrate 51 is formed, durability for polishing the semiconductor substrate 48 with the support substrate 51 can be increased. The support substrate 51 is preferably an X-ray low absorption material. For example, resin or Si. Further, it is desirable that the thermal expansion coefficient of the support substrate 51 is a material close to that of the layer forming the pixel 40. For example, when this layer is formed using single crystal Si, it is single crystal Si, silicon nitride Si ₃ N ₄ , a-Si, p-Si, or the like. When this layer is formed using SiC of a compound semiconductor, it is AlN or the like. Furthermore, when this layer is a TFT formed on a glass substrate, it is SiC, AlN, or the like.

As described above, according to the X-ray image detector 30 of the X-ray imaging system 10, the absorption type grating unit 31 that functions as a grating is integrated into the sensor unit 41 in the X-ray image detector 30. Therefore, it is not necessary to align this many times. Furthermore, since the positional deviation between the absorption type grating unit 31 and the sensor unit 41 is also reduced, the accuracy of the phase information can be increased in the radiation phase imaging for acquiring the phase information of the subject. Further, the configuration of the X-ray image detector 30 is simplified.

In addition, since most X-rays are not diffracted by the X-ray shielding member 31b and geometrically projected onto the sensor unit 41, the irradiated X-rays are not required to have high spatial coherence, and the X-ray source 11 may be a general X-ray source used in the medical field.

Then, the distance L ₂ from X-ray shield member 31b to the sensor unit 41 can be any value, the distance L _2, it is possible to set smaller than the minimum Talbot interference distance in Talbot interferometer The X-ray image detector 30 can be downsized (thinned). Furthermore, in this X-ray imaging system, almost all wavelength components of irradiated X-rays contribute to the projected image (G1 image) from the X-ray shielding member 31b, and the contrast of the G1 image is improved. Detection sensitivity can be improved.

In addition, since the main light emitting region of the scintillator 49 is arranged so as to be close to the pixel 40, the scintillator 49 emits light in the vicinity of the pixel 40, and sensitivity is improved.

Further, since the scanning line 45 is arranged so as to overlap the X-ray shielding member 31b, X-ray absorption by the scanning line 45 can be reduced.

In the X-ray image detector 30, the X-ray shielding member 31b and a part of the scanning line 45 are overlapped in the thickness direction (z direction) of the X-ray shielding member 31b. . For example, a plurality of X-ray shielding members 31b may be arranged in the y direction, and a part of the X-ray shielding member 31b and the signal line 46 may overlap in the thickness direction of the X-ray shielding member 31b.

It is preferable that the X-ray shielding member 31b and the signal line 46 overlap each other in terms of noise reduction.

Further, since the width of the scanning line 45 (or the signal line 46) of the wiring portion 47 is made smaller than the width of the X-ray shielding member 31b in the arrangement direction of the X-ray shielding member 31b, X-rays other than the X-ray shielding member 31b are used. Vignetting can be prevented.

In addition, as a manufacturing method of the X-ray image detector 30, methods and orders other than the above methods may be adopted.

In the X-ray image detector 30 described above, the X-ray shielding member 31b is described as an absorption type grating, but the present invention is not limited to this. As described above, even if the periodic pattern of the Talbot interference image, the G1 image can be analyzed in relation to the arrangement pitch of the pixels 40 of the X-ray image detector 30, and the present invention is also useful in that case. It is. Therefore, the grating is not limited to the absorption type grating but may be a phase type grating.

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

Further, the above-described phase contrast image generation processing may be performed on an image acquired by shooting (pre-shooting) in the absence of a subject to acquire a phase contrast image. This phase contrast image reflects, for example, phase unevenness (initial phase shift) caused by non-uniformity of the X-ray shielding member 31b or the like. By subtracting the phase contrast image in the pre-photographing from the phase contrast image obtained by photographing (main photographing) in the presence of the subject, a phase contrast image in which the phase unevenness of the photographing unit 12 is corrected can be obtained. .

FIG. 10 shows another example of the X-ray image detector 30.

The X-ray image detector 50 differs from the X-ray image detector 30 in that all the scanning lines 45 overlap the X-ray shielding member 31b of the absorption type grating portion 31 in the thickness direction of the X-ray shielding member 31b. That is, each X-ray shielding member 31b and each scanning line 45 are overlapped in a one-to-one correspondence.

For this reason, since the G1 image cannot be directly analyzed, the moire generated due to the minute difference between the pattern period p ₁ ′ in the x direction of the G1 image and the arrangement pitch P in the x direction of the pixels 40 is analyzed. To obtain a phase contrast image.

Here, the arrangement pitch P of the pixels 40, less than _'is 1 / 2p 1 pitch required to detect a periodic pattern of G1 image _{(resolution)'} period p _1.

The moire period T in the x direction of moire generated in the image is expressed by the following equation (7).

When the subject H is disposed between the X-ray source 11 and the absorption type grating unit 31 (X-ray shielding member 31b), moire generated in an image detected by the pixel 40 of the X-ray image detector 50 is modulated by the subject H. Receive. This modulation amount is proportional to the angle of the X-ray deflected by the refraction effect by the subject H. Therefore, a phase contrast image of the subject H can be generated by analyzing this moire.

Here, attention is paid to the period p ₁ ′ of the G1 image and the period T of the moire. The G1 image projected from the X-ray shielding member 31b to the position of the pixel 40 is displaced in the x direction by an amount corresponding to the refraction angle φ due to refraction of the X-ray by the subject H. The moire generated in the image due to a minute difference between the pattern period p ₁ ′ in the x direction of the G1 image and the arrangement pitch P in the x direction of the pixels 40 also changes in the x direction in accordance with the change in the G1 image.

When the displacement amount Δx of the G1 image reaches the period p ₁ ′, the moire returns to the original state. Therefore, the displacement amount ΔX of the moire is expressed by the following equation (8) using the displacement amount Δx of the G1 image. expressed.

This displacement amount ΔX is expressed by the following equation (9) based on the phase shift amount ψ of the signal output from each pixel 40 of the X-ray image detector 50 (the phase shift amount of the signal with and without the subject H). Are related.

Accordingly, by obtaining the phase shift amount ψ of the signal output from each pixel 40, the refraction angle φ is obtained from the equations (4), (8), and (9), and the above equation (5) is used. Since the differential amount of the phase shift distribution Φ (x) is obtained, by integrating this with respect to x, the phase shift distribution Φ (x) of the subject H, that is, the phase contrast image of the subject H can be generated.

Hereinafter, a method for calculating the phase shift amount ψ will be described.

FIG. 11 schematically shows a signal output from each pixel 40 of the X-ray image detector 50.

Since the period p ₁ ′ in the x direction of the periodic pattern of the G1 image and the arrangement pitch P in the x direction of the pixels 40 are in a relationship that causes moire in the image, the signal curve changes periodically with the period T of moire. To do. When the G1 image changes in the x direction, the moire also changes in the x direction, and the phase of the signal curve corresponding to the moire changes. When the displacement amount Δx of the G1 image reaches the period p ₁ ′ of the periodic pattern, the moire displacement amount ΔX becomes the moire period T, and the moire and signal curve return to the original state.

The phase difference between the waveforms of the signal curve (FIG. 7A) when the subject H is not present and the signal curve (FIG. 7B) when the subject H is present is the signal of each pixel 40 constituting the unit. This corresponds to the phase shift amount ψ.

FIG. 12 shows another example of a moire analysis method using the X-ray image detector 50.

In this example, moire is analyzed using Fourier transform and inverse Fourier transform. The moire formed by the interference between the period of the periodic pattern of the X-ray shielding member 31b and the arrangement pitch of the pixels 40 of the X-ray image detector 50 can be expressed by the following expression (10). (11) can be rewritten.

In equation (10), a (x, y) represents the background, b (x, y) represents the amplitude of the spatial frequency component corresponding to the fundamental period of moire, and (f _0x, f _0y ) represents the moire. Represents the basic period. In the formula (11), c (x, y) is represented by the following formula (12).

Therefore, information on the refraction angle φ (x, y) can be obtained by extracting the component of c (x, y) or c ^* (x, y). Here, Formula (11) becomes following Formula (13) by Fourier-transform.

In the formula _{_{_{(13), F (f x}}} , f y), A (f x, f y), C (f x, f y) , respectively f (x, y), a (x, y), c It is a two-dimensional Fourier transform for (x, y).

When using the one-dimensional lattice such as X-ray shielding member 31b, the spatial frequency spectrum of the moire, at least, a peak derived from _{A _(f} x, f _y), the sandwich so C _(f x, Three peaks are generated, including the peak of the spatial frequency component corresponding to the fundamental period of moire derived from f _y ) and C ^* (f _x , f _y ). A _{_(f} x, f _y) peak derived from the origin, _{_{also, C (f x, f y}} ) and ^{_{_{C * (f x, f y}}} ) peak derived from the _{_{(± f 0x, ± f 0y}} ) It occurs at the position of (combined same order).

In order to obtain the refraction angle φ (x, y) from the moire spatial frequency spectrum, the region R including the peak frequency of the spatial frequency component corresponding to the fundamental period of moire is cut out so that the peak frequency overlaps the origin of the frequency space. The extracted region R is moved and inverse Fourier transform is performed. Then, the refraction angle φ (x, y) can be obtained from the complex number information obtained by the inverse Fourier transform.

FIG. 13 shows a manufacturing method of the X-ray image detector 50 of FIG.

As shown in FIG. 13, the manufacturing method of the X-ray image detector 50 is different from the manufacturing method shown in FIGS. 8 and 9 in that a mask pattern is not separately prepared.

First, a solid-state imaging device is prepared in which a sensor unit 41 and a wiring unit 47 in which pixels 40 are already arranged in a two-dimensional manner are formed. At this time, the scanning line 45 of the wiring portion 47 is provided above the region between the pixel rows of the pixels 40.

Next, a support substrate 51 formed of an X-ray low absorption material is formed on the wiring portion 47. Then, a resist film 52 is formed on the wiring part 47.

Also, as the resist film 52, either a negative type or a positive type can be used. In this example, a case where a positive resist film is used will be described.

Then, from the semiconductor substrate 48 side, the resist film 52 is exposed with X-rays or the like from the direction of the arrow (from the wiring portion 47 side). At this time, since the scanning lines 45 of the wiring portion 47 are provided avoiding the pixels 40, each scanning line 45 functions as a mask constituting a mask pattern. The location where the X-ray shielding member 31b is to be formed can be determined by such a self-alignment method.

In order to form a mask, each scanning line 45 preferably has a low X-ray transmittance. For this reason, it is preferable to manufacture the scanning line 45 thickly in the thickness direction. On the other hand, the signal line 46 that does not function as a mask is preferably formed thinner than the scanning line 45. Further, since the scanning line 45 is manufactured to be thick, the wiring resistance can be lowered, and it is preferable to use it for a wiring that requires a low wiring resistance, such as a readout wiring from the pixel 40.

Hereinafter, since the resist film 52 development process is the same as the manufacturing method of the X-ray image detector 30, description thereof is omitted.

As described above, according to the manufacturing method of the X-ray image detector 50, since the scanning line 45 of the wiring part 47 is used as a mask pattern, a groove can be formed in the substrate part 31a without using the mask pattern again.

FIG. 14 shows still another example of the X-ray image detector 30.

The X-ray image detector 60 is different from the X-ray image detector 30 in that the X-ray image detector 60 is configured based on an organic CMOS sensor, and the detection is performed by two-dimensionally arranging the pixels 40 with a photoelectric conversion element using an organic photoelectric conversion material. The unit 61 is configured.

The plurality of pixels 40 include an upper electrode film 64, a lower electrode film 63, and a photoelectric conversion film 62 disposed therebetween. The photoelectric conversion film 62 is composed of an organic photoelectric conversion film.

For the photoelectric conversion film 62, for example, an organic photoelectric conversion material described in JP2009-32854A is used. The photoelectric conversion film 62 absorbs light emitted from the scintillator 49 and generates electric charges according to the absorbed light. In this way, the photoelectric conversion film 62 including the organic photoelectric conversion material has a sharp absorption spectrum in the visible region, and electromagnetic waves other than light emitted by the scintillator 49 are hardly absorbed by the photoelectric conversion film 62, and noise is generated. It can be effectively suppressed.

The organic photoelectric conversion material of the photoelectric conversion film 62 is preferably such that its absorption peak wavelength is closer to the emission peak wavelength of the scintillator 49 in order to absorb light emitted by the scintillator 49 most efficiently. Ideally, the absorption peak wavelength of the organic photoelectric conversion material matches the emission peak wavelength of the scintillator 49, but if the difference between the two is small, the light emitted from the scintillator 49 can be sufficiently absorbed. . Specifically, the difference between the absorption peak wavelength of the organic photoelectric conversion material and the emission peak wavelength with respect to the radiation of the scintillator 49 is preferably within 10 nm, and more preferably within 5 nm.

Examples of organic photoelectric conversion materials that can satisfy such conditions include arylidene organic compounds, quinacridone organic compounds, and phthalocyanine organic compounds. For example, since the absorption peak wavelength in the visible region of quinacridone is 560 nm, if quinacridone is used as the organic photoelectric conversion material and CsI: Tl is used as the material of the scintillator 49, the difference between the peak wavelengths can be made within 5 nm. The amount of charge generated in the photoelectric conversion film 62 can be substantially maximized.

Note that the pixel 40 may be configured by an organic layer including an upper electrode film 64, a lower electrode film 63, and a photoelectric conversion film 62 disposed therebetween. More specifically, this organic layer is a part that absorbs electromagnetic waves, a photoelectric conversion part, an electron transport part, a hole transport part, an electron blocking part, a hole blocking part, a crystallization preventing part, an electrode, and an interlayer contact improvement. It can be formed by stacking or mixing parts.

The thickness of the photoelectric conversion film 62 is preferably as large as possible in terms of absorbing light from the scintillator 49, but considering the ratio that does not contribute to charge separation, it is preferably 30 nm to 300 nm, more preferably 50 nm to 250 nm. Hereinafter, it is particularly preferably 80 nm or more and 200 nm or less.

The upper electrode film 64 is preferably made of a conductive material that is transparent at least with respect to the emission wavelength of the scintillator 49 because light generated by the scintillator 49 needs to enter the photoelectric conversion film 62. Specifically, it is preferable to use a transparent conductive oxide (TCO) that has a high visible light transmittance and a low resistance value.

Although a metal thin film such as Au can be used as the upper electrode film 64, the resistance value tends to increase if an attempt is made to obtain a transmittance of 90% or more, so the TCO is preferred. For example, ITO, IZO, AZO, FTO, SnO2, TiO2, and ZnO2 can be preferably used, and ITO is most preferable from the viewpoint of process simplicity, low resistance, and transparency.

It should be noted that the upper electrode film 64 may have a single configuration common to all the pixels 40 or may be divided for each pixel 40.

Also, the thickness of the upper electrode film 64 can be, for example, 30 nm or more and 300 nm or less.

The lower electrode film 63 is a thin film divided for each pixel 40. The lower electrode film 63 can be made of a transparent or opaque conductive material, and aluminum, silver, or the like can be suitably used. The thickness of the lower electrode film 63 can be, for example, 30 nm or more and 300 nm or less.

For the configuration of the photoelectric conversion element using the organic photoelectric conversion material described above, for example, the description in JP-A-2009-32854 is helpful.

As described above, according to the X-ray image detector 60, since the photoelectric conversion film 62 using the organic photoelectric conversion material is provided in the pixel 40, noise in the phase contrast image can be reduced.

15 and 16 show a method for manufacturing the X-ray image detector 60 of FIG.

First, after a readout circuit (not shown) for each pixel 40 is formed on a support substrate 51 such as a silicon substrate, a wiring portion 47 having a plurality of scanning lines 45 and a plurality of signal lines 46 is formed on the support substrate 51. To do. Further, a lower electrode film 63 is formed on the surface of the wiring portion 47 opposite to the support substrate 51 (FIG. 15A). The lower electrode film 63 is connected to the readout circuit.

Next, similarly to the manufacturing method of the X-ray image detector 30, the absorption type lattice portion 31 having the X-ray shielding member 31b is formed on the side where the lower electrode film 63 is not provided (FIG. 15B). At this time, the support formed of the low X-ray absorption material may be formed on the silicon substrate. However, the groove is formed by etching or the like using the silicon substrate itself as the support and then X-ray shielding is performed by metal plating or the like. The member 31b is formed.

Next, a photoelectric conversion film 62 is formed on the lower electrode film 63 (FIG. 15C).

Next, the upper electrode film 64 is formed on the photoelectric conversion film 62, and the protective film 65 is formed on the upper electrode film 64 (FIG. 16A).

Then, the pixel 40 and a separately manufactured scintillator 49 are optically bonded. The scintillator 49 may be directly deposited on the protective film 65, or an adhesive layer may be provided separately from the protective film 65, and the scintillator 49 may be bonded to the protective film 65 (FIG. 16B).

FIG. 17 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

A mammography system 170 shown in FIG. 17 is an apparatus that captures an X-ray image (phase contrast image) of a breast B as a subject. The mammography system 170 is disposed at one end of an arm member 81 that is pivotably connected to a base (not shown), and disposed at the other end of the arm member 81. An imaging table 83 and a compression plate 84 configured to be movable in the vertical direction with respect to the imaging table 83 are provided.

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

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

FIG. 18 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

The X-ray imaging system 180 is different from the X-ray imaging system 10 of the first embodiment in that the multi-slit 103 is disposed in the collimator unit 102 of the X-ray source 101. Since other configurations are the same as those of the X-ray imaging system 10 described above, description thereof will be omitted.

In the X-ray imaging system 10 described above, when the distance from the X-ray source 11 to the X-ray image detector 30 is set to a distance (1 m to 2 m) set in a general hospital imaging room, X The blur of the G1 image due to the focal size of the line focal point 18b (generally about 0.1 mm to 1 mm) is affected, and there is a possibility that the image quality of the phase contrast image is 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 absorptive grating like the X-ray shielding member 31b, and a plurality of X-ray shielding members extending in one direction (y direction) are in the same direction as the X-ray shielding member 31b of the absorptive grating part 31. They are periodically arranged in the (x direction). The multi-slit 103 partially shields the radiation emitted from the X-ray focal point 18b, thereby reducing the effective focal size in the x direction and forming a large number of point light sources (dispersed light sources) in the x direction. The purpose is to do.

Grating pitch _{p 3} of the multi-slit 103, the distance from the multi-slit 103 to the X-ray shielding member 31b as _{L 3,} the position of the substantially multi-slit 103 is X-ray focal position, the following equation (14) It is necessary to set to satisfy.

Expression (14) is for the projection image (G1 image) of the X-rays emitted from the respective point light sources dispersedly formed by the multi-slit 103 to coincide with each other at the position of the pixel 40 (overlapping). It is a geometric condition.

As described above, in the present X-ray imaging system 180, G1 images based on a plurality of point light sources formed by the multi-slit 103 are superimposed, thereby improving the image quality of the phase contrast image without reducing the X-ray intensity. Can be made. The multi slit 103 described above can be applied to any of the X-ray imaging systems described above.

FIG. 19 is a schematic diagram showing still another example of the radiation image detector of the radiation imaging system of FIG.

The X-ray shielding member 31b of the above-described X-ray image detector 30 is configured such that the periodic arrangement direction of the X-ray shielding member 31b is linear (that is, the lattice plane is planar). Instead, as shown in FIG. 13, the X-ray image detector 70 uses an absorption type grating unit 110 in which the grating surface of the X-ray shielding member 31 b is concaved into a curved surface. Accordingly, in the X-ray image detector 70, the G1 image detection surface is cylindrical. That is, the detection surface of the G1 image by the X-ray image detector 70 is a cylindrical surface having a straight line extending in the y direction passing through the X-ray focal point 18b as a central axis.

Absorption grating 110 is formed by X-ray transparent, and the curved surface of the substrate portion 110a, a plurality of X-ray shielding member 110b is periodically arranged at a predetermined pitch p _1. Each X-ray shielding member 110b is formed of a material excellent in X-ray absorption and extends linearly in the y direction, and the lattice plane of the absorption type grating portion 110 passes through the X-ray focal point 18b and is X-ray shielded. It has a shape along a cylindrical surface with a straight line extending in the extending direction of the member 110b as the central axis.

In addition, when the X-ray image detector 70 is deployed so that the X-ray image detector 70 becomes a surface perpendicular to the optical axis A, the X-ray shielding members 110b in the absorption type grating unit 110 are arranged at a predetermined pitch. are those substantially parallel, and a plurality of X-ray shielding member 110b is included that are arranged parallel to at a predetermined pitch p _1.

The pixel pitch of the X-ray image detector 70 is a pitch that causes moire in the image in relation to the pattern period of the G1 image formed in the X-ray image detector 70.

By making the lattice plane of the X-ray shielding member 110b cylindrical, the X-rays irradiated from the X-ray focal point 18b are all incident perpendicularly to the lattice plane when the subject H is not present. is relaxed the line shielding members 110b limit constraints thickness h ₁ of is not necessary to consider the above-mentioned formula (1).

As described above, each X-ray imaging system described above has the following features compared to the conventional example.

First, since the array of high-definition pixels 40 provided in the X-ray image detector also serves as the second diffraction grating, the G1 image of the first diffraction grating can be directly detected. For this reason, the second diffraction grating of the conventional example becomes unnecessary. For this reason, since the attenuation of X-rays in the second diffraction grating is eliminated, X-rays can be used effectively and the patient's X-ray exposure can be greatly reduced.

In addition, since the first diffraction grating is integrated with the sensor unit 41, there is no relative displacement between the first diffraction grating and the pixel as in the conventional example, and the position adjustment is highly accurate. There is no need to repeat.

Further, when using a high-definition X-ray image detector of about several μm having a pitch equal to or smaller than that of the first diffraction grating, the phase information of the subject can be obtained by one imaging using Fourier transform and inverse Fourier transform. Even in this method, since the fundamental frequency component of moire is higher than that of a subject structure such as a human body structure, the frequency domain including the fundamental frequency component is separated and the differential image of the phase shift is reproduced by inverse Fourier transform. Even if configured, the resolution degradation of the subject structure can be suppressed to a problem-free level. Therefore, it is possible to obtain sufficient resolution for observing the human body structure even with a single shot, and there is no image quality degradation due to subject blurring and grid movement accuracy during multiple shots, such as the fringe scanning method. An accurate phase contrast image can be obtained.

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

Although an indirect conversion type radiation detector that photoelectrically converts the light emitted from the scintillator 49 has been described as an example of a method for detecting X-rays, the present invention is not limited thereto. For example, it is possible to use a direct conversion type radiation detector that converts X-ray photons into electric charges in a semiconductor layer such as a-Se. The direct conversion type has no blur due to scattering of light emitted from the scintillator, and an image with higher resolution than the indirect conversion type can be obtained.

Furthermore, as a radiation image detector, a CMOS (Complementary Metal Metal Oxide Semiconductor) sensor has been described as an example, but the present invention is not limited to this. For example, a CCD (Charge-Coupled Device) sensor or a TFT (Thin-Film-Transistor) sensor may be employed if the resolution is high enough to detect moiré. Currently, a TFT sensor made of a-Si has a higher pixel size than a radiation image detector using a CMOS sensor or a CCD sensor, but has a higher mobility than an element structure change or a-Si. TFT sensor using p-Si, CGS (Continuous Grain Silicon, continuous grain boundary crystal silicon), polycrystalline system represented by ZnO, and amorphous oxide material represented by IGZO (In-Ga-Zn-O) By using, it is possible to achieve higher definition than the current situation. When a TFT sensor is used, a groove such as a glass substrate serving as a support for the TFT sensor is formed by etching as described above, and the groove is filled to form the X-ray shielding member 31b.

As described above, the following items (1) to (20) are disclosed in this specification.
(1) a lattice portion that forms a radiation image including a periodic intensity distribution by passing radiation;
A plurality of pixels for detecting the radiation and accumulating charges, arranged in a matrix and acquiring the radiation image including a periodic pattern based on the periodic intensity distribution, and the lattice unit includes the The lattice unit is fixed to the detection unit, and the lattice unit is formed of a support that transmits the radiation and a shielding member that absorbs the radiation. The support is fixed to the detection unit, and the detection unit is The radiation image detector is supported from the radiation incident side, and the shielding member is formed in the support body.
(2) The radiation image detector according to (1),
The shielding member is a radiation image detector that forms the radiation image by projecting the passing radiation.
(3) The radiation image detector according to (1) or (2),
The shielding member is a radiographic image detector which is an amplitude type grating which is composed of linear members arranged in parallel with each other at a constant pitch and gives an intensity difference to the incident radiation.
(4) The radiation image detector according to any one of (1) to (3),
The detection unit has a wiring unit including a plurality of scanning lines and a plurality of signal lines provided in a matrix for reading the radiation image,
The wiring portion is a radiation image detector provided between the support on which the shielding member is formed and the plurality of pixels.
(5) The radiation image detector according to (4),
The radiographic image detector, wherein the detection unit includes a phosphor that emits light having a wavelength suitable for spectral sensitivity of the plurality of pixels by exposure of the radiation on a side opposite to the wiring unit with respect to the plurality of pixels.
(6) The radiation image detector according to (4) or (5),
The support is formed of a radiation-absorbing material,
The said shielding member is a radiographic image detector formed with the radiation high absorption material with which the several linear groove | channel formed in the said support body with a fixed pitch was filled.
(7) The radiation image detector according to (5) or (6),
The radiation image detector in which the plurality of scanning lines and the plurality of signal lines constituting the wiring unit do not overlap with the plurality of pixels in the thickness direction of the detection unit.
(8) The radiation image detector according to (7),
The shielding member is a radiation image detector that periodically overlaps at least a part of one of the plurality of scanning lines or the plurality of signal lines in the thickness direction of the detection unit.
(9) The radiation image detector according to (8),
The shielding member is a radiation image detector that periodically overlaps at least a part of the plurality of signal lines.
(10) The radiation image detector according to (8) or (9),
The scanning line or signal line that overlaps the shielding member has a width that is perpendicular to the thickness direction of the detection unit and is smaller than the width of the corresponding shielding member.
(11) The radiation image detector according to any one of (8) to (10),
One of the scanning line and the signal line on which the shielding member overlaps is a radiation image detector formed thicker than the other.
(12) The radiation image detector according to any one of (1) to (11),
The plurality of pixels are arranged at a pitch capable of resolving the periodic intensity distribution, and the periodic pattern is a radiation image detector corresponding to the periodic intensity distribution.
(13) The radiological image detector according to any one of (1) to (11),
The plurality of pixels are arranged at a pitch that forms moire in relation to the period of the periodic intensity distribution, and the periodic pattern corresponds to the moire and is 1/2 times the period of the periodic intensity distribution. Smaller radiation image detector.
(14) The radiographic image detector according to any one of (1) to (13), wherein the radiographic image detector includes a CMOS (Complementary Metal Oxide Semiconductor) sensor.
(15) The radiation image detector according to (14), wherein the plurality of pixels include an organic photoelectric conversion film.
(16) A method for manufacturing a radiation image detector according to (6),
On the detection unit, a support substrate to be the support formed of a low radiation absorbing material is formed,
Forming a plurality of linear grooves on the support substrate at a constant pitch;
A method for manufacturing a radiation image detector, wherein the shielding member is formed by filling the plurality of linear grooves with a radiation-absorbing material.
(17) A method for manufacturing the radiation image detector according to (16),
Forming a resist film on the support substrate;
In the thickness direction of the detection portion, a portion of the resist film that is shielded from light by the mask pattern by exposing the resist film from the detection portion side using a scanning line or a signal line on which the shielding member overlaps as a mask pattern Radiation that forms the plurality of linear grooves on the support substrate by curing other portions, imparting etching resistance, and removing the portion of the support substrate corresponding to the mask pattern by etching Manufacturing method of image detector.
(18) The radiation image detector according to any one of (1) to (15),
Arithmetic processing for generating a phase contrast image of the subject based on the modulation generated in the radiation image acquired by the radiation image detector due to the subject arranged in the radiation field incident on the shielding member And
A radiography system comprising:
(19) The radiographic system according to (18),
The arithmetic processing unit calculates a phase of an intensity modulation signal obtained by interpolating pixel values of a plurality of pixels constituting each set, with three or more adjacent pixels among the plurality of pixels as a set. A radiation imaging system that generates a phase contrast image of the subject based on the phase shift amount of the intensity modulation signal when the subject is present and when the subject is absent.
(20) The radiographic system according to (18),
The arithmetic processing unit performs a Fourier transform on the radiation image acquired by the radiation image detector to acquire a spatial frequency spectrum of the radiation image, and obtains a fundamental frequency component of moire in the spatial frequency spectrum. A radiography system that separates a spatial frequency region including the spatial frequency spectrum from the spatial frequency spectrum and generates a partial phase contrast image by performing inverse Fourier transform on the separated spatial frequency region.

The radiation image detector and the radiation imaging system are useful when used for inspection of a subject in medical diagnosis, inspection of an object in nondestructive inspection, and the like.

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 September 30, 2011 (Japanese Patent Application No. 2011-218506), the contents of which are incorporated herein by reference.

DESCRIPTION OF SYMBOLS 10 X-ray imaging system 11 X-ray source 12 Imaging part 13 Console 30 X-ray image detector 31 Absorption-type grating 31a Substrate part 31b X-ray shielding member 40 Pixel 41 Sensor part

Claims

A grating portion that forms a radiation image including a periodic intensity distribution by passing radiation;
A plurality of pixels for detecting the radiation and accumulating charges, arranged in a matrix, and including a detection unit that acquires the radiation image including a periodic pattern based on the periodic intensity distribution, and
The lattice unit is fixed to the detection unit,
The lattice portion is formed of a support that transmits the radiation and a shielding member that absorbs the radiation,
The support is fixed to the detection unit, and the detection unit is supported from the incident side of the radiation,
The shielding member is a radiation image detector formed in the support.
The radiological image detector according to claim 1,
The shielding member is a radiation image detector that forms the radiation image by projecting the radiation passing therethrough.
The radiological image detector according to claim 1 or 2,
The shielding member is a radiographic image detector which is composed of linear members arranged in parallel to each other at a constant pitch and is an amplitude type grating which gives an intensity difference to the incident radiation.
The radiation image detector according to any one of claims 1 to 3,
The detection unit includes a wiring unit including a plurality of scanning lines and a plurality of signal lines provided in a matrix for reading the radiation image,
The wiring portion is a radiation image detector provided between the support on which the shielding member is formed and the plurality of pixels.
The radiological image detector according to claim 4,
The detection unit is a radiological image detector including a phosphor that emits light having a wavelength suitable for a spectral sensitivity of the plurality of pixels by exposure of the radiation on a side opposite to the wiring unit with respect to the plurality of pixels.
The radiological image detector according to claim 4 or 5, wherein
The support is formed of a radiation-absorbing material;
The said shielding member is a radiographic image detector formed with the radiation high absorption material with which the several linear groove | channel formed in the said support body with a fixed pitch was filled.
The radiological image detector according to claim 5 or 6,
The plurality of scanning lines and the plurality of signal lines constituting the wiring unit are radiation image detectors that do not overlap the plurality of pixels in the thickness direction of the detection unit.
The radiological image detector according to claim 7,
The shielding member is a radiological image detector that periodically overlaps at least a part of either one of the plurality of scanning lines or the plurality of signal lines in the thickness direction of the detection unit.
The radiation image detector according to claim 8, comprising:
The shielding member is a radiation image detector that periodically overlaps at least a part of the plurality of signal lines.
The radiological image detector according to claim 8 or 9, wherein
The scanning line or signal line that overlaps the shielding member is a radiation image detector having a width that is perpendicular to the thickness direction of the detection unit and that is smaller than the width of the corresponding shielding member.
The radiographic image detector according to any one of claims 8 to 10,
One of the scanning line and the signal line on which the shielding member overlaps is a radiation image detector formed thicker than the other.
The radiation image detector according to any one of claims 1 to 11,
The plurality of pixels are arranged at a pitch capable of resolving the periodic intensity distribution, and the periodic pattern is a radiation image detector corresponding to the periodic intensity distribution.
The radiation image detector according to any one of claims 1 to 11,
The plurality of pixels are arranged at a pitch that forms moire in relation to the period of the periodic intensity distribution, and the periodic pattern corresponds to the moire and is 1/2 times the period of the periodic intensity distribution. Smaller radiation image detector.
The radiological image detector according to any one of claims 1 to 13, wherein the radiological image detector includes a CMOS (Complementary Metal Oxide Semiconductor) sensor.
The radiation image detector according to claim 14, wherein the plurality of pixels have an organic photoelectric conversion film.
It is a manufacturing method of the radiographic image detector according to claim 6,
On the detection unit, a support substrate to be the support formed of a low radiation absorbing material is formed,
Forming a plurality of linear grooves on the support substrate at a constant pitch;
A method of manufacturing a radiation image detector, wherein the shielding member is formed by filling the plurality of linear grooves with a radiation-absorbing material.
It is a manufacturing method of the radiographic image detector according to claim 16,
Forming a resist film on the support substrate;
In the thickness direction of the detection unit, a portion shielded from light by the mask pattern of the resist film by exposing the resist film from the detection unit side using a scanning line or a signal line on which the shielding member overlaps as a mask pattern Curing other parts, giving etching resistance,
The manufacturing method of the radiographic image detector which forms the several linear groove | channel in the said support substrate by removing the part corresponding to the said mask pattern among the said support substrates by an etching.
The radiation image detector according to any one of claims 1 to 15,
Arithmetic processing for generating a phase contrast image of the subject based on the modulation generated in the radiation image acquired by the radiation image detector due to the subject arranged in the radiation field incident on the shielding member And
A radiography system comprising:
The radiation imaging system according to claim 18,
The arithmetic processing unit calculates a phase of an intensity modulation signal formed by interpolating pixel values of a plurality of pixels constituting each set, with three or more pixels adjacent to each other as a set among the plurality of pixels. A radiation imaging system for generating a phase contrast image of the subject based on a phase shift amount of the intensity modulation signal when the subject is present and when the subject is absent.
The radiation imaging system according to claim 18,
The arithmetic processing unit performs a Fourier transform on the radiation image acquired by the radiation image detector to acquire a spatial frequency spectrum of the radiation image, and obtains a fundamental frequency component of moire in the spatial frequency spectrum. A radiation imaging system that separates a spatial frequency region including the spatial frequency spectrum from the spatial frequency spectrum and generates a partial phase contrast image by performing an inverse Fourier transform on the separated spatial frequency region.