WO2012070580A1 - Radiograph detection device, radiography device, and radiography system - Google Patents

Abstract

The present invention increases image quality and expands the imaging range in radiation phase imaging. An X-ray image detection device (12) is provided with: a first lattice (31); a lattice pattern (32) that has a cycle that essentially matches the pattern cycle of a radiological image formed by means of radiation that has passed through the first lattice; and a radiograph detector (30) that detects the radiological image masked by the lattice pattern. The first lattice contains a plurality of lattice parts (31A) arrayed in a plane that intersects with the direction of progress of the radiation that passes through the first lattice, and the connection section that connects adjacent lattice parts (31A) to each other is formed from a radiation shielding body. Also, in the image data acquired by the radiograph detector, the data of each pixel (40) belonging to a defective region projected by the connection section are complemented by data of pixels (40) belonging to the region that excludes the defective region and that is at the periphery of the pixels (40) of the defective region.

Description

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

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

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

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

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

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

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

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

In an X-ray imaging system using an X-ray Talbot interferometer, the first and second diffraction gratings must be correspondingly large in order to expand the imaging range. However, since the first and second diffraction gratings typically need to be configured with a high aspect ratio with a grating pitch on the order of μm, it is very difficult to accurately manufacture a large-size grating. . Therefore, in the X-ray imaging system described in Patent Document 1, each of the first and second diffraction gratings is configured by connecting a plurality of grating pieces, and each grating piece is relatively small. It has been.

Japanese Unexamined Patent Publication No. 2007-203061

When each of the first and second diffraction gratings is configured by connecting a plurality of grating pieces, the above-described fringe scanning cannot be normally performed at the connecting part of adjacent grating pieces, and the X-rays transmitted through the connecting part are not generated. The area of the incident X-ray image detector is a defective area in which the X-ray phase change caused by the subject cannot be accurately detected. In Patent Document 1, the data of each pixel belonging to a defective area is complemented based on the data of surrounding pixels, or the first and second diffractions are performed so as to avoid the occurrence of such a defective area. Although adjusting the lattice is described, no specific measures are disclosed.

The present invention has been made in view of the above-described circumstances, and is intended to increase the imaging range and improve the image quality in radiation phase imaging.

(1) a first grating, a grating pattern having a period substantially matching a pattern period of a radiation image formed by radiation that has passed through the first grating, and the radiation image masked by the grating pattern The first grating includes a plurality of grating pieces arranged in a plane intersecting the traveling direction of the radiation passing through the first grating, and is adjacent to the grating. A radiographic image detection apparatus in which a connecting portion for connecting pieces is formed of a radiation shield.
(2) A radiographic apparatus comprising: the radiological image detection apparatus according to (1) above; and a radiation source that irradiates radiation toward the first grating.
(3) In the image data acquired by the radiation imaging apparatus of (2) and the radiation image detector, the data of each pixel belonging to the defect area onto which the connecting portion is projected is placed around the pixel. A radiation imaging system including an arithmetic processing unit that complements data of pixels belonging to an area excluding the defective area.

According to the present invention, a plurality of lattice pieces are connected to form a lattice, thereby obtaining a large-size lattice while maintaining the accuracy by using a relatively small individual lattice piece. it can. Then, by forming the connecting portion of the adjacent lattice pieces with a radiation shield, the pixels belonging to the defect area can be easily and reliably extracted, and the data of the pixels can be complemented with certainty. The image quality can be improved.

It is a schematic diagram which shows the structure of an example of the radiography system for describing embodiment of this invention. It is a control block diagram of the radiography system of FIG. It is a schematic diagram which shows the structure of the radiographic image detector contained in the radiography system of FIG. It is a perspective view which shows the structure of the 1st and 2nd grating | lattice. It is a side view which shows the structure of the 1st and 2nd grating | lattice. It is a schematic diagram which shows the mechanism for changing the period of the moire fringe by superimposition of the 1st and 2nd grating | lattice. It is a schematic diagram for demonstrating the refraction | bending of the radiation by a to-be-photographed object. It is a schematic diagram for demonstrating the fringe scanning method. It is a graph which shows the signal of the pixel of the radiographic image detector accompanying a fringe scanning. It is a schematic diagram which shows the structure of the 1st and 2nd grating | lattice for demonstrating distribution of the defect area | region in a radiographic image detector. It is a top view which shows distribution of the defect area | region in a radiographic image detector. It is a schematic diagram which shows the structure of the 1st and 2nd grating | lattice regarding the modification of the radiography system of FIG. It is a schematic diagram which shows distribution of the defect area | region in a radiographic image detector. It is a schematic diagram which shows the structure of the 1st and 2nd grating | lattice regarding the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the radiographic image detector regarding the other example of the radiography system for describing embodiment of this invention. It is a block diagram which shows the structure of the arithmetic processing part regarding the other example of the radiography system for describing embodiment of this invention. It is a graph which shows the signal of the pixel of the radiographic image detector for demonstrating the process in the arithmetic processing part of the radiography system of FIG. It is a schematic block diagram of the other example of the radiography system for describing embodiment of this invention. It is a figure which shows schematic structure of the radiographic image detector of an optical reading system. It is a figure which shows the arrangement | positioning relationship of the pixel of a 1st grating | lattice, a 2nd grating | lattice, and a radiographic image detector. It is a figure for demonstrating the method to set the inclination-angle of the 1st grating | lattice with respect to a 2nd grating | lattice. It is a figure for demonstrating the adjustment method of the inclination-angle of the 1st grating | lattice with respect to a 2nd grating | lattice. It is a figure for demonstrating the effect | action of recording of the radiographic image detector of an optical reading system. It is a figure for demonstrating the effect | action of the reading of the radiation image detector of an optical reading system. It is a figure for demonstrating the effect | action which acquires a some fringe image based on the image signal read from the radiographic image detector of an optical reading system. It is a figure for demonstrating the effect | action which acquires a some fringe image based on the image signal read from the radiographic image detector of an optical reading system. It is a figure which shows the arrangement | positioning relationship between the radiographic image detector using a TFT switch, and the 1st and 2nd grating | lattice. It is a figure which shows schematic structure of the radiographic image detector using CMOS. It is a figure which shows the structure of one pixel circuit of the radiographic image detector using CMOS. It is a figure which shows the arrangement | positioning relationship between the radiographic image detector using CMOS, and the 1st and 2nd grating | lattice. It is a schematic block diagram of the other example of the radiography system for describing embodiment of this invention. It is a figure which shows schematic structure of one Embodiment of a radiographic image detector. It is a figure for demonstrating the effect | action of recording of one Embodiment of a radiographic image detector. It is a figure for demonstrating the effect | action of reading of one Embodiment of a radiographic image detector. It is a figure which shows other embodiment of a radiographic image detector. It is a figure for demonstrating the effect | action of recording of other embodiment of a radiographic image detector. It is a figure for demonstrating the effect | action of the reading of other embodiment of a radiographic image detector. It is a figure which shows other embodiment of a radiographic image detector.

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 (X-ray 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 imaging of the X-ray source 11 based on the operation of the operator. The console 12 is broadly classified into a console 13 that controls the photographing operation of the unit 12 and performs arithmetic processing on image data acquired by the photographing 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 that does not contribute to the inspection area of the subject H. The X-ray tube 18 is of an anode rotating type, and emits an electron beam from a filament (not shown) as an electron emission source (cathode) and collides with a rotating anode 18a rotating at a predetermined speed, thereby causing X-rays. Is generated. The colliding portion of the rotating anode 18a with the electron beam becomes the X-ray focal point 18b.

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

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

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

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

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

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

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 the configurations of the first and second gratings.

The first absorption type lattice 31 is configured by connecting a plurality of first lattice pieces 31A. Each of the first lattice pieces 31A includes a substrate 31a and a plurality of Xs arranged on the substrate 31a. It is comprised from the line shielding part 31b. The second absorption type grating 32 is also configured by connecting a plurality of second grating pieces 32A. Each of the second grating pieces 32A includes a substrate 32a and a plurality of substrates arranged on the substrate 32a. It is comprised from the X-ray shielding part 32b. The substrates 31a and 32a are both made of an X-ray transparent member such as glass that transmits X-rays.

The X-ray shielding portions 31b and 32b are both linear members extending in one direction (y direction in the illustrated example) in a plane perpendicular to the optical axis A of the X-rays emitted from the X-ray source 11. Composed. As a material of each X-ray shielding part 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.

In the illustrated example, the plurality of first lattice pieces 31A are arranged in the same x direction as the arrangement direction of the X-ray shielding portions 31b in a plane orthogonal to the optical axis A, and the lattice pieces 31A adjacent to each other in the x direction are arranged. It is connected. Similarly, the plurality of second lattice pieces 32A are arranged in the x direction within a plane orthogonal to the optical axis A, and the lattice pieces 32A adjacent in the x direction are connected to each other. The lattice pieces 31A and 32A are not limited to the same direction (x direction) as the arrangement direction of the X-ray shielding portions 31b and 32b, and may be arranged in a direction (y direction) orthogonal thereto. They may be two-dimensionally arranged in the x direction and the y direction.

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. Specifically, by setting the distances d ₁ and d ₂ 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 can be obtained at the slit portion. It is configured to pass through without being diffracted while maintaining straightness. For example, when tungsten is used as the rotary anode 18a described above and the tube voltage is 50 kV, the 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, most of the X-rays are projected almost geometrically 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 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 is known from the annual salary 63180S-1.

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

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

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

However, when the X-rays irradiated from the X-ray source 11 are cone beams, if the thicknesses h ₁ and h ₂ of the X-ray shielding portions 31b and 32b are too thick, the X-rays incident obliquely enter the slit portion. There is a problem that it becomes difficult to pass, so-called vignetting occurs, and the effective visual field in the direction (x direction) perpendicular to the extending direction (strand direction) of the X-ray shielding portions 31b and 32b becomes narrow. Therefore, in view of the field of view secured to define the upper limit of the thickness _h 1, _{h 2.} In order to secure the effective field length V in the x direction on the detection surface of the FPD 30, assuming that the distance from the X-ray focal point 18 b to the detection surface of the FPD 30 is L, the thicknesses h ₁ and h ₂ are shown in FIG. From the scientific relationship, it is necessary to set so as to satisfy the following expressions (6) and (7).

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

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

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

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

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

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

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

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

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

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

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

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

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

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

Since the refraction angle φ (x) is a value corresponding to the differential phase value as shown in the above equation (13), the phase shift is obtained by integrating the refraction angle φ (x) along the x-axis. A distribution Φ (x) is obtained. In the above description, the y coordinate in the y direction of the pixel 40 is not taken into consideration. However, by performing the same calculation for each y coordinate, a two-dimensional phase shift distribution Φ (x , Y).

The above calculation is performed by the calculation processing unit 22. The arithmetic processing unit 22 further includes a connection portion between the first lattice pieces 31A adjacent to each other in the first absorption type lattice 31 and a second absorption type lattice in the projection onto the FPD 30 with the X-ray focal point 18b as the viewpoint. The pixels 40 belonging to the region (defect region) where the connection part of the 32 adjacent second lattice pieces 32A is projected are extracted, and the signal values of these pixels 40 are complemented with the signal values of the surrounding pixels 40. The phase shift distribution Φ is calculated based on the complemented signal value. Hereinafter, extraction of the pixel 40 and complementation of the signal value will be described.

FIG. 10 shows the configuration of the first and second gratings, and FIG. 11 shows the distribution of defect areas in the radiation image detector.

The first lattice pieces 31A adjacent to each other in the first absorption type lattice 31 are dispersed with fine particles of heavy metal having an atomic number of 40 or more as an X-ray absorber, such as gold paste, platinum paste, lead-containing solder, etc. The connecting portion 31c is constituted by an X-ray shield made of a hardened adhesive layer. The second lattice pieces 32A adjacent to each other in the second absorption type grating 32 are also connected using an adhesive in which an X-ray absorbing material is dispersed, and the connecting portion 32c is an X-ray made of a cured adhesive layer. It is constituted by a shield. The connecting portions 31c and 32c configured as described above are typically formed to a thickness of about several tens to several hundreds of μm, and the distances d ₁ and d ₁ between the slit portions of the first and second absorption gratings 31 and 32 are formed. It is larger than d ₂ (about 1 to 10 μm) and is substantially equal to or larger than the pitch of the pixels 40 in the FPD 30.

Therefore, the pixel 40 belonging to the first area A1 where the connection part 31c of the first absorption type grating 31 is projected and the second area A2 where the connection part 32c of the second absorption type grating 32 is projected. The signal value of the pixel 40 is smaller than the signal value of the pixel 40 belonging to the third area A3 excluding the first and second areas A1 and A2. Therefore, a predetermined threshold is set for the signal value of the pixel 40, and the pixel 40 having a signal value smaller than this threshold is extracted. The extracted signal value of each pixel 40 is complemented by an appropriate method such as replacement with an average value of signal values of a plurality of pixels 40 that belong to the third region A3 around the pixel.

Here, each pixel 40 belonging to the first area A1 where the connection part of the first absorption type grating 31 is projected and belongs to the second area A2 where the connection part of the second absorption type grating 32 is projected. In other words, at least one pixel 40 belonging to the third region A3 is interposed between each pixel 40, in other words, the first region A1 and the second region A2 do not overlap each other and are adjacent to each other. The first and second absorption type gratings 31 and 32 are arranged so that a gap G larger than the pitch of the pixels 40 of the FPD 30 is placed between both the areas A1 and A2. With this configuration, the pixels 40 belonging to the third region A3 in the immediate vicinity of the pixels 40 belonging to the first and second regions A1 and A2, that is, effective pixels capable of accurately detecting the phase information of the subject by fringe scanning. 40 is provided.

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

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 X-ray imaging system 10, each of the first and second absorption type gratings 31 and 32 is configured by connecting a plurality of grating pieces to each of the grating pieces. The first and second absorption gratings 31 and 32 having a large size can be obtained while maintaining the accuracy by using relatively small ones for 31A and 32A. Thereby, the imaging range can be expanded. Then, by forming a connecting portion between adjacent lattice pieces with an X-ray shield, in the projection onto the FPD 30 with the X-ray focal point 18b as a viewpoint, the pixels 40 belonging to the regions A1 and A2 in which the connecting portion is projected. Extraction can be performed easily and reliably, and the signal value of the pixel 40 can be reliably complemented to improve the image quality of the obtained phase contrast image.

In addition, each pixel 40 belonging to the first area A1 where the connection part 31c of the first absorption type grating 31 is projected and the second area A2 where the connection part 32c of the second absorption type grating 32 is projected. By interposing at least one pixel 40 belonging to the third region A3 between each pixel 40 belonging to the pixel 40, the pixels 40 belonging to the first and second regions A1 and A2 are in close proximity to each other by stripe scanning. The effective pixels 40 capable of accurately detecting the phase information of the subject can be provided, and the signal values of the pixels 40 belonging to the first and second regions A1 and A2 are obtained using the signal values of the effective pixels 40. Can be complemented with high accuracy. Thereby, the image quality of the obtained phase contrast image can be improved.

In addition, since most X-rays are not diffracted by the first absorption type grating 31 and are projected almost geometrically onto the second absorption type grating 32, the irradiated X-rays have high spatial coherence. A general X-ray source used in the medical field can be used as the X-ray source 11 without being required. The distance L ₂ from the first absorption type grating 31 to the second absorption type grating 32 can be set to an arbitrary value, and the distance L ₂ is smaller than the minimum Talbot interference distance in the Talbot interferometer. Since it can be set, the photographing unit 12 can be downsized (thinned). Furthermore, in this X-ray imaging system, almost all wavelength components of irradiated X-rays contribute to the projection image (G1 image) from the first absorption type grating 31 and the contrast of moire fringes is improved. Contrast image detection sensitivity can be improved.

Note that the above-described X-ray imaging system 10 calculates the refraction angle φ by performing fringe scanning on the projection image of the first grating, and therefore the first and second gratings absorb both. Although described as a mold lattice, the present invention is not limited to this. As described above, the present invention is also useful when the refraction angle φ is calculated by performing fringe scanning on the Talbot interference image. Therefore, the first grating is not limited to the absorption type grating but may be a phase type grating. Further, the method of analyzing the moire fringes formed by superimposing the X-ray image of the first grating and the second grating is not limited to the above-described fringe scanning method. For example, “J. Opt. Soc. Am. Vol” 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.

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

Further, a phase differential image (a differential amount of the phase shift distribution Φ) may be created from an image group acquired by photographing (pre-photographing) in the absence of a subject. This phase differential image reflects the phase unevenness of the detection system (including phase shift due to moire, grid nonuniformity, 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. 12 shows the configuration of the first and second gratings for a modification of the radiation imaging system of FIG. 1, and FIG. 13 shows the distribution of defect areas in the radiation image detector.

In this modification, the first absorption grating 31 includes a plurality of first grating pieces 31A arranged in the x direction and the y direction, and the first grating pieces 31A adjacent in the x direction are arranged in the y direction. Adjacent first lattice pieces 31A are connected to each other. Also in the second absorption type grating 32, a plurality of second grating pieces 32A are arranged in the x direction and the y direction, and the second grating pieces 32A adjacent in the x direction are adjacent to each other in the y direction. The lattice pieces 32A are connected to each other. The connecting portions between adjacent lattice pieces are all formed by an X-ray shield.

And each pixel 40 which belongs to 1st area | region A1 in which the connection part 31cx of two 1st grating | lattice pieces 31A adjacent to x direction is projected, and two 2nd grating | lattice pieces 32A adjacent to x direction similarly. At least one pixel 40 belonging to the third area excluding the first and second areas A1 and A2 is interposed between each pixel 40 belonging to the second area A2 onto which the connecting portion 32cx is projected. . In other words, the first area A1 and the second area A2 do not overlap each other, and a gap larger than the pitch of the pixels 40 of the FPD 30 is placed between the adjacent areas A1 and A2. Thereby, the effective pixel 40 can be provided in the immediate vicinity of each pixel 40 belonging to the first and second regions A1 and A2, and the first and second regions are used by using the signal values of these effective pixels 40. The signal values of the pixels 40 belonging to A1 and A2 can be complemented with high accuracy.

Furthermore, each pixel 40 belonging to the fourth region A4 on which the connecting portion 31cy of the two first grid pieces 31A adjacent in the y direction is projected, and two second grid pieces 32A that are also adjacent in the y direction At least one pixel 40 belonging to the sixth area A6 excluding the fourth and fifth areas A4 and A5 is interposed between each pixel 40 belonging to the fifth area A5 on which the connection 32cy portion is projected. Yes. In other words, the fourth area A4 and the fifth area A5 do not overlap each other, and a gap larger than the pitch of the pixels 40 of the FPD 30 is placed between the adjacent areas A4 and A5. Thereby, the effective pixel 40 can be provided in the immediate vicinity of each pixel 40 belonging to the fourth and fifth regions A4, A5, and the fourth and fifth regions can be obtained using the signal values of the effective pixels 40. The signal values of the pixels 40 belonging to A4 and A5 can be complemented with high accuracy.

In addition, it is preferable that the first region A1 and the second region A2 do not overlap with each other in the x direction, and the fourth region A4 and the fifth region A5 do not overlap with each other in the y direction. Only one of the x direction and the y direction may be such that the two regions do not overlap. In this direction, the effective pixel 40 may be provided in the immediate vicinity of each pixel 40 belonging to both regions. it can.

FIG. 14 shows the configuration of the first and second gratings for another modification of the radiation imaging system of FIG.

In the present modification, the first grating piece 31A and the second grating piece 32A have X-rays of the first absorption type grating 31 and the second absorption type grating 32 in a geometric shape excluding their thickness. It is similar according to the ratio (L ₁ / (L ₁ + L ₂ )) of the distance from the focal point 18b. Then, the arrangement direction of the first grating pieces 31A in the first absorption type grating 31 and the number of arrangements in each direction, and the arrangement direction of the second grating pieces 32A in the second absorption type grating 32 and the number of arrangements in each direction. The centers of the first and second absorption gratings 31 and 32 are both located on the optical axis A of X-rays.

According to the above configuration, the region where the connection part 32c of the second absorption type grating 32 is projected is included in the region where the connection part 31c of the first absorption type grating 31 is projected. Thereby, the number of pixels 40 whose signal values are complemented can be reduced, and the phase information of the subject can be acquired with higher accuracy.

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

A mammography apparatus 80 shown in FIG. 15 is an apparatus that captures an X-ray image (phase contrast image) of a breast B as a subject. The mammography apparatus 80 is disposed at one end of an arm member 81 that is pivotally connected to a base (not shown), and disposed at the other end of the arm member 81. An imaging table 83 and a compression plate 84 configured to be movable in the vertical direction with respect to the imaging table 83 are provided.

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

Since the X-ray source 11 and the imaging unit 12 have the same configuration as that of the X-ray imaging system 10 described above, the same reference numerals as those of the X-ray imaging system 10 described above are attached to the respective components. Yes. 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. 16 shows a modification of the radiation imaging system of FIG.

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

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

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

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

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

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

The multi-slit 103 is an absorption type grating (third absorption type grating) having a configuration similar to that of the first and second absorption type gratings 31 and 32 provided in the imaging unit 12, and is in one direction (y direction). The extended X-ray shielding portions are periodically arranged in the same direction (x direction) as the X-ray shielding portions 31b and 32b of the first and second absorption gratings 31 and 32. The multi-slit 103 is intended to form a large number of small-focus light sources (dispersed light sources) arranged at a predetermined pitch in the x direction by partially shielding the radiation emitted from the X-ray focal point 18b. .

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

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

In the X-ray imaging system 10 described above, the phase contrast image obtained by the fringe scanning is an X-ray refraction component in the periodic array direction (x direction) of the X-ray shielding portions of the first and second absorption gratings 31 and 32. And the refractive component in the extending direction (y direction) of the X-ray shielding part is not included. For this reason, there is a portion that cannot be depicted depending on the shape and orientation of the subject H. For example, when the direction of the load surface of the articular cartilage is matched with the y direction, it is considered that the peripheral tissue of the cartilage (such as tendons and ligaments) having a shape perpendicular to the load surface is insufficiently depicted. By moving the subject H, it is possible to recapture a region that is not sufficiently depicted, but in addition to increasing the burden on the subject H and the operator, the position reproducibility with the recaptured image is ensured. There is a problem that it is difficult.

Accordingly, the first and second absorption gratings 31 and 32 are centered on a virtual line (X-ray optical axis A) orthogonal to the center of the grating surface of the first and second absorption gratings 31 and 32. It is also possible to provide a lattice rotation mechanism 105 that rotates integrally from the first direction and sets the second direction to generate a phase contrast image in each of the first direction and the second direction. it can.

In the illustrated example, the first and second absorption gratings 31 and 32 are rotated by 90 °, and the first direction and the second direction are orthogonal to each other. As long as the direction intersects, the rotation angle of the first and second absorption gratings 31 and 32 is not limited to 90 °. Further, the grating rotating mechanism 105 may be configured to rotate only the first and second absorption type gratings 31 and 32 separately from the FPD 30, or the first and second absorption type gratings 31. , 32 and the FPD 30 may be rotated together. Further, when the multi-slit 103 is provided, an X-ray source in which the multi-slit 103 and the collimator unit 102 are integrally formed so that the rotation coincides with the first and second absorption gratings 31 and 32 is used. Rotate. Furthermore, the generation of phase contrast images in the first and second orientations using the grating rotation mechanism 105 can be applied to any of the X-ray imaging systems described above.

FIG. 19 shows the configuration of the first and second gratings for another example of the radiation imaging system for explaining the embodiment of the present invention.

In the X-ray imaging system 10 described above, the first and second absorption gratings 31 and 32 are arranged such that the periodic arrangement direction of the X-ray shielding portions 31b and 32b is linear (that is, the grating surface is planar). However, instead of this, as shown in FIG. 19, it is also possible to use first and second absorption type gratings 110 and 111 having a substantially concave curved surface.

The first absorption type grating 110 is configured by connecting a plurality of first grating pieces 110A, and each of the first grating pieces 110A is formed on the surface of a planar substrate 110a that is X-ray transparent. , a plurality of X-ray shielding section 110b that extends linearly in the y direction are periodically arranged at a predetermined pitch p _1. The first absorption type grating 110 includes a plurality of first gratings in the circumferential direction of the cylindrical surface with the imaginary line passing through the X-ray focal point 18b and extending in the extending direction (y direction) of the X-ray shielding part 110b as a central axis. When the pieces 110A are arranged and the first lattice pieces 110A adjacent in the circumferential direction are connected to each other, the lattice surface is formed in a substantially concave curved surface shape. A connecting portion between adjacent first lattice pieces 110A is formed of an X-ray shield.

Similarly, the second absorption type grating 111 is configured by connecting a plurality of second grating pieces 111A, and each of the second grating pieces 111A is an X-ray transparent and planar substrate 111a. on the surface of a plurality of X-ray shielding section 111b that extends linearly in the y direction are periodically arranged at a predetermined pitch p _2. The second absorption type grating 111 includes a plurality of second gratings in the circumferential direction of the cylindrical surface with a virtual line passing through the X-ray focal point 18b and extending in the extending direction (y direction) of the X-ray shielding part 111b as a central axis. When the pieces 111A are arranged and the first lattice pieces 111A adjacent to each other in the circumferential direction are connected to each other, the lattice surface is formed in a substantially concave curved surface shape. A connecting portion between the adjacent second lattice pieces 111A is formed of an X-ray shield.

_{L 1} the distance from the X-ray focal point 18b to the first absorption grating 110, when the distance from the first absorption grating 110 to the second absorption grating 111 was _{L 2,} the grating pitch _{p 2} are Are determined so as to satisfy the relationship of the above formula (1).

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

Further, when the grating surfaces of the first and second absorption type gratings 110 and 111 are formed in a concave curved surface shape, one of the first and second absorption type gratings 110 and 111 is centered on the X-ray focal point 18b. As described above, the above-described fringe scanning is performed by moving in a direction along the lattice plane. Further, when the grating surfaces of the first and second absorption gratings 110 and 111 are formed in a concave curved surface, the detection surface of the FPD 112 is also a cylinder whose central axis is a straight line that extends in the y direction through the X-ray focal point 18b. It is preferable to form a concave curved surface along the surface.

The first and second absorption gratings 110 and 111 and the FPD 112 can be applied to any of the X-ray imaging systems described above. Furthermore, it is also preferable that the multi slit 103 has the same shape as the first and second absorption gratings 110 and 111.

FIG. 20 shows the configuration of the radiation image detector in relation to another example of the radiation imaging system for explaining the embodiment of the present invention.

In the X-ray imaging system 10 described above, the second absorption type grating 32 is provided independently of the FPD 30, but the X-ray image detector itself has the second absorption type grating 32 or an equivalent configuration. You may do it. As a specific embodiment, the second absorption type grating can be eliminated by using an X-ray image detector having a configuration disclosed in Japanese Patent Laid-Open No. 2009-133823. This X-ray image detector is a direct conversion type X-ray image detector including a conversion layer that converts X-rays into electric charges, and a charge collection electrode that collects electric charges converted in the conversion layer, The charge collecting electrode 121 of the pixel 120 is configured by arranging a plurality of linear electrode groups 122 to 127 formed by electrically connecting linear electrodes arranged at a constant period so that their phases are different from each other. Has been.

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

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

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

For example, when the X-ray image detector configured as described above is applied to the X-ray imaging system 10 described above, the second absorption type grating 32 is not required from the imaging unit 12, and a plurality of images can be obtained by one imaging. Since a phase component fringe image can be acquired, physical scanning for fringe scanning becomes unnecessary, and the scanning mechanism 33 can be eliminated. Thereby, it is possible to reduce the cost and further reduce the thickness of the photographing unit. It should be noted that the structure of the charge collecting electrode may be replaced with another structure described in Japanese Patent Application Laid-Open No. 2009-133823.

FIG. 21 shows the configuration of the calculation unit of another example of the radiation imaging system for explaining the embodiment of the present invention.

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

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

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

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

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

According to the present X-ray imaging system, an absorption image and a small angle scattered image are generated from a plurality of images acquired for the phase contrast image of the subject. No deviation occurs, and the phase contrast image can be satisfactorily superimposed with the absorption image or the small-angle scattered image.

FIG. 23 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

The X-ray imaging system of the present example has a first grating 131 that forms the first periodic pattern image by passing the X-rays emitted from the X-ray source 11 and the first grating 131. A second grating 132 that forms a second periodic pattern image by modulating the intensity of the periodic pattern image, and an X-ray image detector (radiation) that detects the second periodic pattern image formed by the second grating 132. Image detector) 240 and a second periodic pattern image detected by X-ray image detector 240 acquire a fringe image, and generate a phase contrast image based on the acquired fringe image Part 260. The phase contrast image generation unit 260 constitutes a part of the processing of the control device 20 in the console 13 (FIG. 2) and the processing of the arithmetic processing unit 22.

The X-ray source 11 emits X-rays toward the subject H, and has spatial coherence that can generate a Talbot interference effect when the first grating 131 is irradiated with X-rays. is there. For example, a microfocus X-ray tube or a plasma X-ray source having a small X-ray emission point size can be used. Further, when an X-ray source having a relatively large X-ray emission point (so-called focal spot size) as used in a normal medical field is used, a multi-slit (for example, the multi-slit 103 described above) having a predetermined pitch is used. It can be used by being installed between the X-ray source 11 and the first grating 131.

The first grating 131 is desirably a so-called phase modulation type grating that gives a phase modulation of about 90 ° or about 180 ° with respect to the irradiated X-ray. For example, when the X-ray shielding portion is gold In the normal X-ray energy region for medical diagnosis, the necessary thickness h ₁ is about 1 μm to several μm. An amplitude modulation type grating can also be used as the first grating 131. On the other hand, the second grating 132 is preferably an amplitude modulation type grating.

When the X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam, the self-image of the first grating 131 formed through the first grating 131 is an X-ray source. It is enlarged in proportion to the distance from 11. In this example, the grating pitch P ₂ of the second grating 132 substantially matches the periodic pattern of the bright part of the self-image of the first grating 131 at the position of the second grating 132. To be determined. That is, when the distance from the focal point of the X-ray source 11 to the first grating 131 is L ₁ and the distance from the first grating 131 to the second grating 132 is L ₂ , the second grating pitch p ₂ is Are determined so as to satisfy the relationship of the above formula (1).

When the X-rays emitted from the X-ray source 11 are parallel beams, it is determined so as to satisfy p ₂ = p ₁ .

As in the first absorption type grating 31 of the X-ray imaging system 10 described above, the first grating 131 has a plurality of first grating pieces arranged one-dimensionally or two-dimensionally, and these first grating pieces. It is configured by connecting each other, and the connecting portion between the adjacent first lattice pieces is configured by an X-ray shield. Similarly to the second absorption grating 32 of the X-ray imaging system 10 described above, the second grating 132 also has a plurality of second grating pieces arranged one-dimensionally or two-dimensionally. The pieces are connected to each other, and the connecting portion between the adjacent second lattice pieces is formed of an X-ray shield.

The X-ray image detector 240 detects an image in which the self-image of the first grating 131 formed by the X-rays incident on the first grating 131 is intensity-modulated by the second grating 132 as an image signal. . In this example, the X-ray image detector 240 is a direct-conversion X-ray image detector that reads an image signal by scanning with linear reading light. X-ray image detector.

FIG. 24 shows the appearance (FIG. 24A), xz plane cross section (FIG. 24B), and yz plane cross section (FIG. 24C) of the X-ray image detector 240.

The X-ray image detector 240 of this example includes a first electrode layer 241 that transmits X-rays, and a recording photoconductive layer 242 that generates charges when irradiated with X-rays transmitted through the first electrode layer 241. The charge transport layer 244, which acts as an insulator for charges of one polarity among the charges generated in the recording photoconductive layer 242, and acts as a conductor for charges of the other polarity, reading light The photoconductive layer for reading 245 that generates an electric charge when irradiated with the first electrode layer 246 and the second electrode layer 246 are laminated in this order. In the vicinity of the interface between the recording photoconductive layer 242 and the charge transport layer 244, a power storage unit 243 that accumulates charges generated in the recording photoconductive layer 242 is formed. Note that each of the above layers is formed on the glass substrate 247 in order from the second electrode layer 246.

The first electrode layer 241 only needs to transmit X-rays. For example, Nesa film (SnO ₂ ), ITO (Indium Tin Oxide), IZO (Indium Zinc Oxide), amorphous light-transmitting oxide film IDIXO (Idemitsu Indium X-metal Oxide; Idemitsu Kosan Co., Ltd.) having a thickness of 50 to 200 nm can be used, and Al or Au having a thickness of 100 nm can also be used.

The recording photoconductive layer 242 only needs to generate charge when irradiated with X-rays, and is excellent in that it has relatively high quantum efficiency and high dark resistance with respect to X-rays. A material mainly composed of a-Se is used. The thickness is suitably 10 μm or more and 1500 μm or less. In particular, when it is used for mammography, it is preferably 150 μm or more and 250 μm or less, and when used for general photographing, it is preferably 500 μm or more and 1200 μm or less.

As the charge transport layer 244, for example, the larger the difference between the mobility of charges charged in the first electrode layer 241 during recording of an X-ray image and the mobility of charges having the opposite polarity, the better (for example, 102 Or more, preferably 103 or more), for example, poly N-vinylcarbazole (PVK), N, N′-diphenyl-N, N′-bis (3-methylphenyl)-[1,1′-biphenyl] -4, Organic compounds such as 4'-diamine (TPD) and discotic liquid crystal, or TPD polymer (polycarbonate, polystyrene, PVK) dispersion, semiconductor materials such as a-Se and As ₂ Se ₃ doped with 10 to 200 ppm of Cl Is appropriate. A thickness of about 0.2 to 2 μm is appropriate.

The reading photoconductive layer 245 may be any material that exhibits conductivity when irradiated with reading light. For example, a-Se, Se-Te, Se-As-Te, metal-free phthalocyanine, metal phthalocyanine, A photoconductive substance mainly composed of at least one of MgPc (Magnesium phthalocyanine), VoPc (phase II of Vanadyl phthalocyanine), CuPc (Cupper phthalocyanine), and the like is preferable. A thickness of about 5 to 20 μm is appropriate.

The second electrode layer 246 includes a plurality of transparent linear electrodes 246a that transmit the reading light and a plurality of light shielding linear electrodes 246b that shield the reading light. The transparent linear electrode 246a and the light-shielding linear electrode 246b extend linearly continuously from one end of the image forming area of the X-ray image detector 240 to the other end. The transparent linear electrodes 246a and the light-shielding linear electrodes 246b are alternately arranged in parallel at predetermined intervals (FIG. 24A, 24B).

The transparent linear electrode 246a is made of a conductive material while transmitting reading light. For example, as with the first electrode layer 241, ITO, IZO, or IDIXO can be used. The thickness is about 100 to 200 nm.

The light shielding linear electrode 246b shields the reading light and is made of a conductive material. For example, the above transparent conductive material and a color filter can be used in combination. The thickness of the transparent conductive material is about 100 to 200 nm.

In the X-ray image detector 240 of this example, as will be described in detail later, an image signal is read out using a pair of the adjacent transparent linear electrode 246a and the light shielding linear electrode 246b. That is, an image signal of one pixel is read out by one set of the transparent linear electrode 246a and the light shielding linear electrode 246b (FIG. 24B). In this example, the transparent linear electrode 246a and the light shielding linear electrode 246b are arranged so that one pixel is approximately 50 μm.

The X-ray imaging system of this example includes a linear reading light source 250 extending in a direction (x direction) orthogonal to the extending direction of the transparent linear electrode 246a and the light shielding linear electrode 246b (FIG. 24A). The linear reading light source 250 of this example is composed of a light source such as an LED (Light Emitting Diode) or LD (Laser Diode) and a predetermined optical system, and the linear reading light having a width of about 10 μm is detected as an X-ray image detector. It is comprised so that 240 may be irradiated. The linear reading light source 250 is moved in the extending direction (y direction) of the transparent linear electrode 246a and the light shielding linear electrode 246b by a predetermined moving mechanism (not shown). The X-ray image detector 240 is scanned by the linear reading light emitted from the light source 250 and the image signal is read out. The operation of reading the image signal will be described in detail later.

In order for the configuration including the X-ray source 11, the first grating 131, the second grating 132, and the X-ray image detector 240 to function as a Talbot interferometer, some conditions must be substantially satisfied. The conditions will be described below.

First, the grid surfaces of the first grating 131 and the second grating 132 must be parallel to the xy plane shown in FIG.

Further, the distance Z ₂ (Talbot interference distance Z) between the first grating 131 and the second grating 132 is the following when the first grating 131 is a phase modulation type grating that applies 90 ° phase modulation. When Expression (20) is substantially satisfied and the first grating 131 is a phase modulation type grating that gives 180 ° phase modulation, the following Expression (21) must be approximately satisfied.

Where λ is the X-ray wavelength (usually effective wavelength), m is 0 or a positive integer, p ₁ is the lattice pitch of the first grating 131 described above, and p ₂ is the grating pitch of the second grating 132 described above. It is.

Further, when the first grating 131 is an amplitude modulation type grating, the above formula (2) must be substantially satisfied with respect to the Talbot interference distance Z.

Further, the thicknesses h ₁ and h ₂ of the first and second gratings 131 and 132 are also set so as to satisfy the expressions (6) and (7) described above with respect to the first and second gratings 31 and 32. Must be set.

Further, in the X-ray imaging system of the present example, as shown in FIG. 25, the first grating 131 and the second grating 132 are formed by extending the first grating 131 and the second grating 132. It is arranged so that the direction is relatively inclined. Then, with respect to the first grating 131 and the second grating 132 arranged in this way, the main scanning direction (x direction in FIG. 24) of each pixel of the image signal detected by the X-ray image detector 240. The main pixel size Dx and the sub-pixel size Dy in the sub-scanning direction have a relationship as shown in FIG.

As described above, the main pixel size Dx is determined by the arrangement pitch of the transparent linear electrodes 246a and the light shielding linear electrodes 246b of the X-ray image detector 240, and is set to 50 μm in this example. . The sub-pixel size Dy is determined by the width of the linear reading light irradiated to the X-ray image detector 240 by the linear reading light source 250, and is set to 10 μm in this example. .

Here, in this example, a plurality of fringe images are acquired, and a phase contrast image is generated based on the plurality of fringe images. If the number of acquired fringe images is M, M subpixel sizes are obtained. The first grating 131 is tilted with respect to the second grating 132 so that Dy becomes one image resolution D in the sub-scanning direction of the phase contrast image.

Specifically, as shown in FIG. 26, a periodic pattern image formed on the position of the second grating 132 by the pitch of the second grating 132 and the first grating 131 (hereinafter referred to as self of the first grating 131). The pitch of the image G1) is p, the relative rotation angle of the self-image of the first grating 131 with respect to the second grating 132 in the xy plane is θ, and the image resolution in the sub-scanning direction of the phase contrast image is D (= Dy × M), the rotation angle θ is set so as to satisfy the following expression (22), so that the self-image G1 of the first grating 131 and the first image G1 with respect to the length of the image resolution D in the sub scanning direction The phase of the second grating 132 is shifted by n periods. FIG. 26 shows a case where M = 5 and n = 1.

Therefore, an image signal obtained by dividing the intensity modulation for n periods of the self-image of the first grating 131 by M can be detected by each pixel of Dx × Dy obtained by dividing the image resolution D of the phase contrast image in the sub-scanning direction by M. Become. In the example shown in FIG. 26, since n = 1, the phase of the self-image G1 of the first grating 131 and the second grating 132 is shifted by one period with respect to the length of the image resolution D in the sub-scanning direction. It will be. More simply, the range that passes through the second grating 132 for one period of the self-image G1 of the first grating 131 changes over the length of the image resolution D in the sub-scanning direction.

Since M = 5, an image signal obtained by dividing the intensity modulation of one period of the self-image of the first grating 131 into five can be detected by each pixel of Dx × Dy, that is, each pixel of Dx × Dy. Thus, it is possible to detect image signals of five different fringe images. The method for acquiring the image signals of the five striped images will be described in detail later.

In this example, as described above, since Dx = 50 μm, Dy = 10 μm, and M = 5, the image resolution Dx in the main scanning direction and the image resolution D = Dy × M in the sub-scanning direction of the phase contrast image are obtained. Although it is the same, it is not always necessary to match the image resolution Dx in the main scanning direction and the image resolution D in the sub scanning direction, and an arbitrary main / sub ratio may be used.

Furthermore, in this example, M = 5, but M may be 3 or more and may be other than 5. In the above description, n = 1, but n may be an integer other than 1 as long as n is an integer other than 0. That is, when n is a negative integer, the rotation is opposite to that in the above-described example, and n may be an intensity modulation for n periods with n being an integer other than ± 1. However, when n is a multiple of M, the phases of the self-image G1 of the first grating 131 and the second grating 132 are equal between one set of M sub-scanning direction pixels Dy, and M different numbers Since it is not a striped image, it is excluded.

Regarding the rotation angle θ of the self-image of the first grating 131 with respect to the second grating 132, for example, after fixing the relative rotation angle of the X-ray image detector 240 and the second grating 132, the first grating 131. This can be done by rotating 131.

For example, when p = 5 μm, D = 50 μm, and n = 1 in the equation (22), the theoretical rotation angle θ is about 5.7 °. Then, the actual rotation angle θ ′ of the self-image of the first grating 131 relative to the second grating 132 can be detected by, for example, the self-image of the first grating and the moire pitch by the second grating 132. .

Specifically, as shown in FIG. 27, if the actual rotation angle is θ ′ and the pitch of the apparent self-image in the x direction generated by the rotation is P ′, the observed moire pitch Pm is 1 Since / Pm = | 1 / P′−1 / p |, the actual rotation angle θ ′ can be obtained by substituting P ′ = p / cos θ ′ into the above equation. The moire pitch Pm may be obtained based on the image signal detected by the X-ray image detector 240.

Then, the theoretical rotation angle θ and the actual rotation angle θ ′ may be compared, and the rotation angle of the first grating 131 may be adjusted automatically or manually based on the difference.

The phase contrast image generation unit 260 generates an X-ray phase contrast image based on image signals of M kinds of different fringe images detected by the X-ray image detector 240.

Next, the operation of the X-ray imaging system of this example will be described.

First, as shown in FIG. 23, after the subject H is arranged between the X-ray source 11 and the first grating 131, X-rays are emitted from the X-ray source 11. Then, the X-ray passes through the subject H and is then irradiated on the first grating 131. The X-rays irradiated to the first grating 131 are diffracted by the first grating 131 to form a Talbot interference image at a predetermined distance from the first grating 131 in the optical axis direction of the X-ray.

This is called the Talbot effect. When a light wave passes through the first grating 131, a self-image of the first grating 131 is formed at a predetermined distance from the first grating 131. For example, when the first grating 131 is a phase modulation type grating that gives 90 ° phase modulation, it is given by Expression (20) (Expression (21) in the case of a 180 ° phase modulation type grating or intensity modulation type grating). While the self-image of the first grating 131 is formed at a given distance, the wavefront of the X-ray incident on the first grating 131 is distorted by the subject H, so that the self-image of the first grating 131 is deformed accordingly. Yes.

Subsequently, the X-ray passes through the second grating 132. As a result, the self-image of the deformed first grating 131 is intensity-modulated by being superimposed on the second grating 132, and is detected by the X-ray image detector 240 as an image signal reflecting the wavefront distortion. Is done.

Here, with reference to FIGS. 28 and 29, the operation of image detection and readout in the X-ray image detector 240 will be described.

First, in a state where a negative voltage is applied to the first electrode layer 241 of the X-ray image detector 240 by the high-voltage power supply 400, intensity modulation is performed by superposition of the self-image of the first grating 131 and the second grating 132. The X-rays thus emitted are irradiated from the first electrode layer 241 side of the X-ray image detector 240 (FIG. 28A).

Then, the X-rays irradiated to the X-ray image detector 240 are transmitted through the first electrode layer 241 and irradiated to the recording photoconductive layer 242. The X-ray irradiation generates a charge pair in the recording photoconductive layer 242, and the positive charge is combined with the negative charge charged in the first electrode layer 241 and disappears, and the negative charge is latent. The image charges are accumulated in the power storage unit 243 formed at the interface between the recording photoconductive layer 242 and the charge transport layer 244 (FIG. 28B).

Next, as shown in FIG. 29, in the state where the first electrode layer 241 is grounded, the linear reading light L1 emitted from the linear reading light source 250 is irradiated from the second electrode layer 246 side. . The reading light L1 passes through the transparent linear electrode 246a and is applied to the reading photoconductive layer 245, and the positive charge generated in the reading photoconductive layer 245 by the irradiation of the reading light L1 passes through the charge transport layer 244. The negative charge is combined with the positive charge charged on the light shielding linear electrode 246b through the charge amplifier 200 connected to the transparent linear electrode 246a.

A current flows through the charge amplifier 200 due to the combination of the negative charge generated in the reading photoconductive layer 245 and the positive charge charged in the light shielding linear electrode 246b, and this current is integrated and detected as an image signal. The

Then, the linear reading light source 250 moves in the sub-scanning direction to scan the X-ray image detector 240 with the linear reading light L1, and the above-described reading lines are irradiated with the linear reading light L1. The image signals are sequentially detected by the action, and the detected image signals for each reading line are sequentially input and stored in the phase contrast image generation unit 260.

Then, after the entire surface of the X-ray image detector 240 is scanned with the reading light L1 and the image signal of one frame is stored in the phase contrast image generation unit 260, the phase contrast image generation unit 260 stores the stored image. Based on the signal, image signals of five different fringe images are acquired.

Specifically, in this example, as shown in FIG. 26, the image resolution D in the sub-scanning direction of the phase contrast image is divided into five, and the intensity modulation of one period of the self image of the first grating 131 is divided into five. Since the first grating 131 is inclined with respect to the second grating 132 so that the detected image signal can be detected, as shown in FIG. 30, the image signal read from the first reading line is the first The image signal acquired as the fringe image signal M1 and read from the second reading line is acquired as the second fringe image signal M2, and the image signal read from the third reading line is the third fringe image signal M3. And the image signal read from the fourth reading line is acquired as the fourth fringe image signal M4, and the image signal read from the fifth reading line is acquired as the fifth fringe image signal M5. . Note that the first to fifth reading lines shown in FIG. 30 correspond to the sub-pixel size Dy shown in FIG.

In FIG. 30, only the reading range of Dx × (Dy × 5) is shown, but the first to fifth fringe image signals are acquired in the same manner as described above for the other reading ranges. That is, as shown in FIG. 31, an image signal of a pixel row group composed of pixel rows (reading lines) every four pixel intervals in the sub-scanning direction is acquired, and one stripe image signal of one frame is acquired. More specifically, the image signal of the pixel row group of the first reading line is acquired to acquire the first stripe image signal of one frame, and the image signal of the pixel row group of the second reading line is acquired to 1 The second stripe image signal of the frame is acquired, the image signal of the pixel row group of the third reading line is acquired, the third stripe image signal of one frame is acquired, and the image of the pixel row group of the fourth reading line A signal is acquired to acquire a fourth stripe image signal of one frame, an image signal of a pixel row group of the fifth reading line is acquired, and a fifth stripe image signal of one frame is acquired.

As described above, the first to fifth fringe image signals different from each other are acquired, and the phase contrast image generation unit 260 generates a phase contrast image based on the first to fifth fringe image signals.

The method for generating the phase contrast image in this example is the same as that already described with reference to the equations (11) to (17), and thus the description thereof is omitted.

In the above-described configuration in which the first grating 131 and the second grating 132 are tilted, both the first grating 131 and the second grating 132 are configured as absorption (amplitude modulation type) gratings, and the Talbot interference effect is obtained. Irrespective of the presence or absence of, it is good also as a structure which projects the radiation which passed the slit part substantially geometrically. In this case, by a distance d ₁ of the first grating 131 and a distance d ₂ of the second grating 132, and sufficiently larger than the effective wavelength of the X-rays emitted from the X-ray source 11, illumination Most of the X-rays are configured to pass through while maintaining straightness without being diffracted by the slit portion. For example, when tungsten is used as the target of the X-ray source and the tube voltage is 50 kV, the effective wavelength of X-ray is about 0.4 mm. In this case, the distance d ₁ of the first grating 131 spacing d ₂ of the second grating 132, substantially geometrically without being most of the radiation is diffracted by the slit portion be about 1 [mu] m ~ 10 [mu] m Projected. The grating pitch p ₁ of the first grating 131 for the relationship between the lattice pitch p ₂ of the second grating 132 is the same as when the first grating 131 described above is a phase modulation type grating. Further, the inclination of the first grating 131 with respect to the second grating 132 is also the same as in the above example, and the generation of the phase contrast image is performed in the same manner as in the above example.

In the above example, as the X-ray image detector 240, a so-called optical reading type X-ray image detector in which an image signal is read out by scanning linear reading light emitted from the linear reading light source 250 is used. However, the present invention is not limited to this. For example, as described in Japanese Patent Application Laid-Open No. 2002-26300, a large number of TFT switches are arranged two-dimensionally, and image signals are read by turning on and off the TFT switches. An X-ray image detector using a TFT switch or an X-ray image detector using a CMOS may be used.

Specifically, an X-ray image detector using a TFT switch includes, for example, a pixel electrode 271 and a pixel electrode 271 that collect charges photoelectrically converted in a semiconductor film by X-ray irradiation as shown in FIG. A number of pixel circuits 270 each including a TFT switch 272 for reading out the collected charges as an image signal are arranged in a two-dimensional manner. An X-ray image detector using a TFT switch is provided for each pixel circuit row, and is provided for each pixel circuit column and a large number of gate electrodes 273 from which a gate scanning signal for turning on and off the TFT switch 272 is output. And a plurality of data electrodes 274 from which the charge signal read from each pixel circuit 270 is output. The detailed layer configuration of each pixel circuit 270 is the same as the layer configuration described in Japanese Patent Laid-Open No. 2002-26300.

For example, when the second grid 132 and the pixel circuit array (data electrode) are installed in parallel, one pixel circuit array corresponds to the main pixel size Dx described in the above example, The pixel circuit row corresponds to the sub-pixel size Dy described in the above example. Note that the main pixel size Dx and the sub-pixel size Dy can be set to 50 μm, for example.

Similarly to the above example, when M striped images are used to generate the phase contrast image, the M pixel circuit rows have one image resolution D in the sub-scanning direction of the phase contrast image. The first grating 131 is inclined with respect to the second grating 132. The specific rotation angle of the first grating 131 is calculated by the equation (22) as in the above example.

In Expression (22), for example, when the rotation angle θ of the first grating 131 is set with M = 5 and n = 1, one period of the self-image of the first grating 131 by one pixel circuit 270 in FIG. In other words, image signals of five different fringe images can be detected by the five pixel circuit rows connected to the five gate electrodes 273 shown in FIG. 32, respectively. It will be possible to detect. In FIG. 32, one second grating 132 and the self-image G1 are shown corresponding to one pixel circuit array, but in actuality, one pixel circuit array corresponds to one pixel circuit array. A large number of second gratings 132 and self-images G1 may exist, and FIG. 32 is not shown.

Therefore, the image signal read from the pixel circuit row connected to the first read line gate electrode G11 is acquired as the first stripe image signal M1, and the pixel circuit connected to the second read line gate electrode G12. The image signal read from the row is acquired as the second stripe image signal M2, and the image signal read from the pixel circuit row connected to the third read line gate electrode G13 is the third stripe image signal M3. The image signal read from the pixel circuit row connected to the fourth read line gate electrode G14 is acquired as the fourth stripe image signal M4 and connected to the fifth read line gate electrode G15. The image signal read from the pixel circuit row is acquired as the fifth fringe image signal M5.

The method for generating the phase contrast image based on the first to fifth fringe image signals is the same as the above example. As described above, when the size of one pixel circuit 270 in the main scanning direction and the sub scanning direction is 50 μm, the image resolution in the main scanning direction of the phase contrast image is 50 μm, and the image resolution in the sub scanning direction is 50 μm × 5 = 250 μm.

As an X-ray image detector using CMOS, for example, a pixel circuit 280 that generates visible light upon receiving X-ray irradiation and photoelectrically converts the visible light to detect a charge signal is shown in FIG. As shown in FIG. 3, a plurality of two-dimensional arrays can be used. The X-ray image detector using CMOS is provided for each pixel circuit row, and includes a large number of gate electrodes 282 and reset electrodes from which a drive signal for driving a signal readout circuit included in the pixel circuit 280 is output. 284 and a plurality of data electrodes 283 that are provided for each pixel circuit column and output a charge signal read from the signal reading circuit of each pixel circuit 280. The gate electrode 282 and the reset electrode 284 are connected to a row selection scanning unit 285 that outputs a drive signal to the signal readout circuit, and the data electrode 283 performs predetermined processing on the charge signal output from each pixel circuit. A signal processing unit 286 to be applied is connected.

As shown in FIG. 34, each pixel circuit 280 includes a lower electrode 806 formed above the substrate 800 via an insulating film 803, a photoelectric conversion film 807 formed on the lower electrode 806, and a photoelectric conversion film 807. An upper electrode 808 formed above, a protective film 809 formed on the upper electrode 808, and an X-ray conversion film 810 formed on the protective film 809 are provided.

The X-ray conversion film 810 is made of, for example, CsI: TI that emits light having a wavelength of 550 nm when irradiated with X-rays. The thickness is preferably about 500 μm.

The upper electrode 808 is made of a conductive material that is transparent to the incident light because it is necessary to make light having a wavelength of 550 nm incident on the photoelectric conversion film 807. The lower electrode 806 is a thin film divided for each pixel circuit 280 and is formed of a transparent or opaque conductive material.

The photoelectric conversion film 807 is formed of, for example, a photoelectric conversion material that absorbs light having a wavelength of 550 nm and generates a charge corresponding to the light. As such a photoelectric conversion material, for example, an organic semiconductor, an organic material containing an organic dye, a material in which an inorganic semiconductor crystal having a direct transition type band gap and a large absorption coefficient is used alone or in combination are used.

Then, by applying a predetermined bias voltage between the upper electrode 808 and the lower electrode 806, one of the charges generated in the photoelectric conversion film 807 moves to the upper electrode 808 and the other moves to the lower electrode 806. .

In the substrate 800 below the lower electrode 806, a charge accumulating portion 802 for accumulating the charges transferred to the lower electrode 806 corresponding to the lower electrode 806, and the charges accumulated in the charge accumulating portion 802. And a signal readout circuit 801 for converting the signal into a voltage signal and outputting it.

The charge storage portion 802 is electrically connected to the lower electrode 806 by a conductive material plug 804 formed through the insulating film 803. The signal readout circuit 801 is configured by a known CMOS circuit.

When an X-ray image detector using a CMOS as described above is installed so that the second grating 132 and the pixel circuit array (data electrode) are parallel as shown in FIG. The pixel circuit column corresponds to the main pixel size Dx described in the above example, and one pixel circuit row corresponds to the sub pixel size Dy described in the above example. Note that the main pixel size Dx and the sub-pixel size Dy can be set to 10 μm, for example, in the case of an X-ray image detector using CMOS.

Similarly to the above example, when M striped images are used to generate the phase contrast image, the M pixel circuit rows have one image resolution D in the sub-scanning direction of the phase contrast image. The first grating 131 is inclined with respect to the second grating 132. The specific rotation angle of the first grating 131 is calculated by the equation (22) as in the above example.

In Expression (22), for example, when M = 5 and n = 1 and the rotation angle θ of the first grating 131 is set, one pixel circuit 280 in FIG. In other words, the image signals of five different fringe images can be detected by the five pixel circuit rows connected to the five gate electrodes 282 shown in FIG. 33, respectively. It will be possible to detect. In FIG. 35, one second grating 132 and the self-image G1 are shown corresponding to one pixel circuit array. However, in actuality, one pixel circuit array corresponds to one pixel circuit array. Many second gratings 132 and self-images G1 may exist, and FIG. 35 is not shown.

Therefore, as in the case of the X-ray image detector using the TFT switch, the image signal read from the pixel circuit row connected to the first read line gate electrode G11 is acquired as the first fringe image signal M1. Then, the image signal read from the pixel circuit row connected to the second read line gate electrode G12 is acquired as the second stripe image signal M2, and the pixel circuit connected to the third read line gate electrode G13. The image signal read from the row is acquired as the third stripe image signal M3, and the image signal read from the pixel circuit row connected to the fourth read line gate electrode G14 is the fourth stripe image signal M4. And the image signal read from the pixel circuit row connected to the fifth read line gate electrode G15 is acquired as the fifth fringe image signal M5.

The method for generating the phase contrast image based on the first to fifth fringe image signals is the same as the above example. As described above, when the size of one pixel circuit 280 in the main scanning direction and the sub scanning direction is 10 μm, the image resolution in the main scanning direction of the phase contrast image is 10 μm, and the image resolution in the sub scanning direction is 10 μm × 5 = 50 μm.

As described above, an X-ray image detector using a TFT switch or an X-ray image detector using a CMOS can be used. However, these X-ray image detectors have square pixels. When the present invention is applied, the resolution in the sub-scanning direction becomes worse than the resolution in the main scanning direction. On the other hand, in the optical reading type X-ray image detector described in the above example, the resolution Dx is limited in the main scanning direction by the width of the linear electrode (direction perpendicular to the extending direction). Regarding the direction, the resolution Dy is determined by the product of the width of the reading light of the linear reading light source 250 in the sub-scanning direction, the accumulation time of the charge amplifier 200 per line, and the moving speed of the linear reading light source 250. Both the main and sub resolutions are typically several tens of μm, but it is possible to increase the sub scanning direction resolution while maintaining the main scanning direction resolution. For example, the X-ray image detector of the optical reading system can be realized by reducing the width of the linear reading light source 250 or reducing the moving speed, and has a more advantageous configuration.

In addition, since a plurality of fringe image signals can be acquired in one shooting, not only the semiconductor detector that can be used immediately and repeatedly as described above, but also a stimulable phosphor sheet or silver salt film can be used. can do. In this case, the reading pixel when reading the stimulable phosphor sheet or the developed silver salt film corresponds to the pixel in the claims.

FIG. 36 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

The X-ray imaging system shown in FIG. 36 detects the periodic pattern image formed by the grating 131 that forms the periodic pattern image by passing the X-rays emitted from the X-ray source 11, and the period. An X-ray image detector (radiation image detector) 340 that performs intensity modulation on the pattern image, a moving mechanism 333 that moves the X-ray image detector 340 in a direction orthogonal to the extending direction of the linear electrode, and X The line image detector 340 includes a phase contrast image generation unit 260 that generates a phase contrast image based on a fringe image obtained by intensity modulation of the periodic pattern image.

Also in this example, a multi-slit (for example, the multi-slit 103 described above) having a predetermined pitch can be installed between the X-ray source 11 and the first grating 131 and used.

The X-ray image detector 340 detects a self-image of the grating 131 formed by the grating 131 when the X-rays pass through the grating 131 and divides a charge signal corresponding to the self-image into a lattice shape to be described later. By accumulating in the charge storage layer, intensity modulation is performed on the self-image to generate a fringe image, and the generated fringe image is output as an image signal. In this example, the X-ray image detector 340 is a direct conversion type X-ray image detector that reads an image signal by scanning with a linear reading light. X-ray image detector.

FIG. 37 shows the external appearance (FIG. 37A), xz plane cross section (FIG. 37B), and yz plane cross section (FIG. 37C) of the X-ray image detector 340.

The X-ray image detector 340 includes a first electrode layer 241 that transmits X-rays, a recording photoconductive layer 242 that generates charges when irradiated with X-rays transmitted through the first electrode layer 241, and recording A charge storage layer 343 that acts as an insulator for charges of one polarity of the charges generated in the photoconductive layer 242 and acts as a conductor for charges of the other polarity, and is irradiated with reading light. A photoconductive layer for reading 245 that generates electric charges when received and a second electrode layer 246 are stacked in this order. Note that each of the above layers is formed on the glass substrate 247 in order from the second electrode layer 246.

The charge storage layer 343 may be any film that is insulative with respect to the polar charge to be stored, such as an acrylic organic resin, polyimide, BCB, PVA, acrylic, polyethylene, polycarbonate, polyetherimide, or the like, or As ₂ S. ₃ , sulfides such as Sb ₂ S ₃ and ZnS, oxides and fluorides. Furthermore, it is more preferable that it is insulative with respect to the charge of the polarity to be accumulated and that it is conductive with respect to the charge of the opposite polarity, and the product of mobility × life is 3 digits or more depending on the polarity of the charge. Substances with differences are preferred.

Preferred compounds include As ₂ Se ₃ , As ₂ Se ₃ doped with Cl, Br, and I from 500 ppm to 20000 ppm, and As ₂ Se ₃ with Se ₂ substituted to about 50% _by Te. _1-x ) ₃ (0.5 <x <1), As ₂ Se ₃ with Se replaced to about 50%, As _x Se with As concentration changed by about ± 15% from As ₂ Se _{3 y} (x + y = 100, 34 ≦ x ≦ 46), amorphous Se—Te system and Te of 5-30 wt%.

Note that as a material of the charge storage layer 343, in order to prevent bending of electric lines of force formed between the first electrode layer 241 and the second electrode layer 246, the dielectric constant thereof is a recording light. It is desirable to use a conductive layer 242 and a photoconductive layer for reading 245 having a dielectric constant that is 1/2 times or more and 2 times or less.

The charge storage layer 343 in this example is divided into lines so as to be parallel to the extending direction of the transparent linear electrode 246a and the light shielding linear electrode 246b of the second electrode layer 246.

The charge storage layer 343 is divided at a pitch finer than the arrangement pitch of the transparent linear electrodes 246 a or the light shielding linear electrodes 246 b, and the arrangement pitch p ₂ and the interval d ₂ are different depending on the combination of the grating 131. It is determined so that imaging can be performed. Note that the arrangement pitch p ₂ and the interval d ₂ of the transparent linear electrodes 246 a or the light shielding linear electrodes 246 b are determined in the same manner as the pitch p ₂ and the interval d ₂ for the _second grating 132 described above, and therefore the same reference numerals are used. I will explain.

Specifically, when the X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam, the self-image of the grating 131 formed through the grating 131 is the X-ray source 11. Enlarged in proportion to the distance from Then, in this embodiment, the arrangement pitch p ₂ of the charge storage layer 343, the portion of the linear charge accumulation layer 343 is approximately coincident with the periodic pattern of the light area of the self-image of the grating 131 at the position of the charge accumulation layer 343 To be decided. That is, the grating pitch of the grating 131 is p ₁ , the distance between the X-ray shielding portions of the grating 131 is d ₁ , the distance from the focal point of the X-ray source 11 to the grating 131 is L ₁ , and the grating 131 to the X-ray image detector 340 When the distance to the detection surface is L ₂ , the arrangement pitch p ₂ of the charge storage layer 343 is determined so as to satisfy the relationship of the above formula (1).

The charge storage layer 343 is formed with a thickness of 2 μm or less in the stacking direction (z direction).

The charge storage layer 343 can be formed by resistance heating vapor deposition using, for example, the above-described material and a mask formed of a metal mask or a fiber having a hole in a metal plate. Further, it may be formed using photolithography.

In the X-ray image detector 340 of this example, as will be described in detail later, an image signal is read out using a pair of the adjacent transparent linear electrode 246a and the light shielding linear electrode 246b. That is, an image signal of one pixel is read out by one set of the transparent linear electrode 246a and the light shielding linear electrode 246b (FIG. 37B). In this example, the transparent linear electrode 246a and the light shielding linear electrode 246b are arranged so that one pixel is approximately 50 μm.

The X-ray imaging system of this example includes a linear reading light source 250 extending in a direction (x direction) orthogonal to the extending direction of the transparent linear electrode 246a and the light shielding linear electrode 246b (FIG. 37A).

In order to allow the configuration including the X-ray source 11, the grating 131, and the X-ray image detector 340 having the charge storage layer 343 divided as described above to function as a Talbot interferometer, several conditions are further satisfied. Must be almost satisfied. The conditions will be described below.

First, it is necessary that the detection surfaces of the grating 131 and the X-ray image detector 340 are parallel to the xy plane shown in FIG.

Further, the distance Z ₂ (Talbot interference distance Z) between the grating 131 and the detection surface of the X-ray image detector 340 is equal to the above formula when the grating 131 is a phase modulation type grating that applies 90 ° phase modulation. When (20) is substantially satisfied and the grating 131 is a phase modulation type grating that gives 180 ° phase modulation, the above equation (21) must be substantially satisfied.

When the grating 131 is an amplitude modulation type grating, the above formula (2) must be substantially satisfied with respect to the Talbot interference distance Z.

As described above, the moving mechanism 333 changes the relative position between the grating 131 and the X-ray image detector 340 by translating the X-ray image detector 340 in a direction orthogonal to the extending direction of the linear electrode. It is something to be made. The moving mechanism 333 is configured by an actuator such as a piezoelectric element, for example.

Next, the operation of the X-ray imaging system of this example will be described.

X-rays pass through the subject H and then irradiate the grating 131. The X-rays irradiated to the grating 131 are diffracted by the grating 131, thereby forming a Talbot interference image at a predetermined distance from the grating 131 in the optical axis direction of the X-ray.

The self-image of the grating 131 is incident from the first electrode layer 241 side of the X-ray image detector 340, undergoes intensity modulation by the charge storage layer 343 of the X-ray image detector 340, and reflects only the wavefront. An X-ray image detector 340 detects the image signal of the fringe image.

Here, with reference to FIGS. 38 and 39, the operation of detecting and reading out the fringe image in the X-ray image detector 340 will be described in more detail.

First, in a state where a negative voltage is applied to the first electrode layer 241 of the X-ray image detector 340 by the high-voltage power supply 400, the X-ray carrying the self-image of the grating 131 is the first X-ray image detector 340. Is irradiated from the electrode layer 241 side (FIG. 38A).

Then, the X-rays irradiated to the X-ray image detector 340 are transmitted through the first electrode layer 241 and irradiated to the recording photoconductive layer 242. The X-ray irradiation generates a charge pair in the recording photoconductive layer 242, and the positive charge is combined with the negative charge charged in the first electrode layer 241 and disappears, and the negative charge is latent. The image charge is stored in the charge storage layer 343 (FIG. 38B).

Here, since the charge storage layer 343 in this example is linearly divided at the arrangement pitch as described above, the charge storage layer 343 is directly below the charge generated according to the self-image of the lattice 131 in the recording photoconductive layer 242. Only the charges in which the charge storage layer 343 exists are trapped and stored by the charge storage layer 343, and other charges pass between the linear charge storage layers 343 (hereinafter referred to as non-charge storage regions), After passing through the reading photoconductive layer 245, it flows out to the transparent linear electrode 246a and the light shielding linear electrode 246b.

As described above, by accumulating only the electric charge generated in the recording photoconductive layer 242 and the linear electric charge accumulation layer 343 immediately below the electric charge, the self-image of the lattice 131 is changed into the linear shape of the electric charge accumulation layer 343. An image signal of a fringe image that is subjected to intensity modulation by superimposing the pattern and the distortion of the wavefront of the self-image by the subject H is accumulated in the charge accumulation layer 343. That is, the charge storage layer 343 of this example performs the same function as the second grating in phase imaging using two conventional gratings.

Next, as shown in FIG. 39, in the state where the first electrode layer 241 is grounded, the linear reading light L1 emitted from the linear reading light source 250 is irradiated from the second electrode layer 246 side. Is done. The reading light L1 passes through the transparent linear electrode 246a and is irradiated to the reading photoconductive layer 245, and the positive charge generated in the reading photoconductive layer 245 by the irradiation of the reading light L1 is a latent image in the charge storage layer 343. The negative charge is combined with the positive charge charged to the light shielding linear electrode 246b through the charge amplifier 200 connected to the transparent linear electrode 246a while being combined with the charge.

A current flows through the charge amplifier 200 due to the combination of the negative charge generated in the reading photoconductive layer 245 and the positive charge charged in the light shielding linear electrode 246b, and this current is integrated and detected as an image signal. The

Then, the linear reading light source 250 moves in the sub-scanning direction (y direction), the X-ray image detector 340 is scanned with the linear reading light L1, and the reading line irradiated with the linear reading light L1. The image signal is sequentially detected by the above-described operation every time, and the detected image signal for each reading line is sequentially input to the phase contrast image generation unit 260 and stored.

Then, the entire surface of the X-ray image detector 340 is scanned with the reading light L 1, and the image signal of the entire frame is stored in the phase contrast image generation unit 260.

The principle of the method for generating the phase contrast image in this example is the same as the content described with reference to the equations (11) to (17), and thus the description thereof is omitted. The phase contrast image generation unit 260 generates a phase contrast image based on the plurality of fringe images.

Incidentally, X-ray imaging system described above, so that the distance Z ₂ from the grating 131 to the detection surface of the X-ray image detector 340 becomes Talbot interference distance, the above equation (20) or formula (21) or formula ( 2) is satisfied, but the grating 131 may be configured to project incident X-rays without being diffracted. According to this configuration, since the projected image projected through the grating 131 is obtained similarly at all positions behind the grating 131, the distance from the grating 131 to the detection surface of the X-ray image detector 340 is obtained. the Z _2, can be set independently of the Talbot interference distance.

Next, a modified example of the above X-ray imaging system will be described.

In the X-ray imaging system described above, the X-ray image detector 340 is translated by the moving mechanism 333, and X-ray images are captured at each position to acquire M fringe image signals. The X-ray imaging system of the example is configured to be able to acquire M striped image signals by capturing one X-ray image without requiring the moving mechanism 333 as described above.

That is, as described with reference to FIGS. 25 to 31 and the like, also in this example, the grating 131 and the X-ray image detector 340 include the extension direction of the grating 131 and the charge storage layer of the X-ray image detector 340. It arrange | positions so that the extending | stretching direction of 343 may incline relatively. The main pixel size Dx in the main scanning direction (x direction in FIG. 37) of each pixel of the image signal detected by the X-ray image detector 340 with respect to the lattice 131 and the charge storage layer 343 thus arranged. And the sub-pixel size Dy in the sub-scanning direction have a relationship as shown in FIG.

Similarly to the configuration and operation described with reference to FIGS. 25 to 31 and the like, after one radiographic image is taken, the entire surface of the X-ray image detector 340 is scanned with the reading light L1. Then, the image signal of the entire frame is stored in the phase contrast image generation unit 260, and the phase contrast image generation unit 260 acquires the image signals of five different fringe images based on the stored image signal. Based on the first to fifth fringe image signals, the phase contrast image generation unit 260 generates a phase contrast image in the same manner as in the above example.

In the above example, the X-ray image detector 340 is provided with three layers of the recording photoconductive layer 242, the charge storage layer 343, and the reading photoconductive layer 245 between the electrodes. However, this layer configuration is not necessarily required. For example, as shown in FIG. 40, the transparent photoelectrode 246a and the light shielding electrode 246b of the second electrode layer are provided without providing the reading photoconductive layer 245. A linear charge storage layer 343 may be provided so as to be in direct contact with the recording medium, and a recording photoconductive layer 242 may be provided on the charge storage layer 343. The recording photoconductive layer 242 also functions as a reading photoconductive layer.

This structure is a structure in which the charge storage layer 343 is provided directly on the second electrode layer 246 without the reading photoconductive layer 245, and the linear charge storage layer 343 can be easily formed. That is, the linear charge storage layer 343 can be formed by vapor deposition. In this vapor deposition step, a metal mask or the like is used to selectively form a linear pattern. However, in the configuration in which the linear charge storage layer 343 is provided on the reading photoconductive layer 245, the reading photoconductive layer 245 is provided. Because of the process of setting the metal mask after vapor deposition, the photoconductive layer 245 for reading is deteriorated by an operation in the atmosphere between the vapor deposition process of the read photoconductive layer 245 and the vapor deposition process of the recording photoconductive layer 242. There is a risk that foreign matter may enter between the photoconductive layers and cause degradation of quality. By adopting a structure in which the above-described reading photoconductive layer 245 is not provided, an operation in the air after the photoconductive layer is deposited can be reduced, so that the above-described fear of quality deterioration can be reduced.

Hereinafter, the operation of recording and reading out the X-ray image of the X-ray image detector 360 shown in FIG. 40 will be described.

First, in the state where a negative voltage is applied to the first electrode layer 241 of the X-ray image detector 360 by the high-voltage power supply 400, the X-ray carrying the self-image of the grating 131 is the first X-ray image detector 360. From the electrode layer 241 side (FIG. 41A).

Then, the X-rays irradiated to the X-ray image detector 340 are transmitted through the first electrode layer 241 and irradiated to the recording photoconductive layer 242. The X-ray irradiation generates a charge pair in the recording photoconductive layer 242, and the positive charge is combined with the negative charge charged in the first electrode layer 241 and disappears, and the negative charge is latent. The image charge is stored in the charge storage layer 343 (FIG. 41B). Note that since the linear charge storage layer 343 in contact with the second electrode layer 246 is an insulating film, charges that have reached the charge storage layer 343 are captured there and go to the second electrode layer 246. Can't, and stays accumulated.

Here, as in the X-ray image detector 340 of the above example, by accumulating only the electric charge generated in the recording photoconductive layer 242 and the electric charge accumulating layer 343 directly below it, The self-image of the lattice 131 is intensity-modulated by being superimposed on the linear pattern of the charge storage layer 343, and an image signal of a fringe image reflecting the distortion of the wavefront of the self-image by the subject H is stored in the charge storage layer 343. Will be.

Then, as shown in FIG. 42, in the state where the first electrode layer 241 is grounded, the linear reading light L1 emitted from the linear reading light source 250 is irradiated from the second electrode layer 246 side. The reading light L1 passes through the transparent linear electrode 246a and is applied to the recording photoconductive layer 242 in the vicinity of the charge storage layer 343. Positive charges generated by the irradiation of the reading light L1 are linear charge storage layer 343. Attracted to recombine. The other negative charge is drawn to the transparent linear electrode 246a, and the light shielding linear electrode is connected to the positive charge charged in the transparent linear electrode 246a and the charge amplifier 200 connected to the transparent linear electrode 246a. It couple | bonds with the positive charge charged to 246b. As a result, a current flows through the charge amplifier 200, and this current is integrated and detected as an image signal.

Even when the above-described X-ray image detector 360 is used, the method for acquiring a plurality of fringe image signals and the method for generating a phase contrast image are the same as those in the above examples.

In each of the above examples, the charge storage layer 343 of the X-ray image detector 340 is formed to be completely separated into a linear shape. However, the present invention is not limited to this, for example, as shown in FIG. You may make it form in a grid | lattice form by forming a linear pattern on flat plate shape.

In the X-ray image detector of the optical reading system described in the above example, the resolution Dx is limited in the main scanning direction by the width of the linear electrode (direction perpendicular to the extending direction), but in the sub-scanning direction. The resolution Dy is determined by the product of the reading light of the linear reading light source 250 in the sub-scanning direction, the accumulation time of the charge amplifier 200 per line, and the moving speed of the linear reading light source 250. Both the main and sub resolutions are typically several tens of μm, but it is possible to increase the sub scanning direction resolution while maintaining the main scanning direction resolution. For example, this can be realized by reducing the width of the linear reading light source 250 or by reducing the moving speed.

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

As described above, the present specification includes a first grating and a grating pattern having a period substantially matching the pattern period of a radiation image formed by radiation that has passed through the first grating; A radiation image detector for detecting the radiation image masked by the grating pattern, wherein the first grating is arranged in a plane intersecting a traveling direction of the radiation passing through the first grating. A radiation image detection device is disclosed that includes a plurality of lattice pieces and a connecting portion that connects adjacent lattice pieces is formed of a radiation shield.

In the radiological image detection apparatus disclosed in this specification, the lattice pattern is a second lattice, and the second lattice intersects with the traveling direction of the radiation passing through the second lattice. A connecting portion that includes a plurality of lattice pieces arranged inside and connects adjacent lattice pieces is formed of a radiation shield.

Further, the radiological image detection apparatus disclosed in the present specification is configured such that the radiographic image detector has a first area in which the connection portion of the first lattice is projected in a projection with a radiographic focal point as a viewpoint, A second region on which a connecting portion of the second lattice is projected, and the first region and the second region coincide with each other, or among the first region and the second region One area is included in the other area.

Further, in the radiological image detection device disclosed in this specification, in each of the first and second grids, the plurality of grid pieces are arranged in a first direction, and the radiographic image detector includes: In projection from the viewpoint of the radiation focal point, a first region where a connecting portion of lattice pieces adjacent to each other in the first direction of the first lattice is projected, and the first direction of the second lattice Each of the pixels belonging to the first region, the second region on which the connecting portion of the lattice pieces adjacent to each other is projected, and the third region excluding the first region and the second region; At least one pixel belonging to the third region is interposed between each pixel belonging to the second region.

In the radiological image detection apparatus disclosed in this specification, in each of the first and second gratings, the plurality of grating pieces are also arranged in a second direction intersecting the first direction. The radiological image detector includes a fourth region in which a connection portion of lattice pieces adjacent to each other in the second direction of the first lattice is projected in a projection with a radiation focus as a viewpoint; A fifth region where a connecting portion of lattice pieces adjacent to each other in the second direction of the second lattice is projected, and a sixth region excluding the fourth region and the fifth region, At least one pixel belonging to the sixth area is interposed between each pixel belonging to the fourth area and each pixel belonging to the fifth area.

In the radiological image detection apparatus disclosed in this specification, in each of the first and second gratings, a surface on which the plurality of grating pieces are arranged is a cylindrical surface, and a central axis thereof has a radiation focus. Pass through.

Further, in the radiological image detection apparatus disclosed in this specification, the connecting portion is formed of an adhesive in which a radiation absorbing material is dispersed.

Further, in the radiological image detection apparatus disclosed in the present specification, the radiation absorbing material is heavy metal particles having an atomic number of 40 or more.

In the radiological image detection apparatus disclosed in this specification, the heavy metal is at least one selected from the group consisting of gold, platinum, and lead.

Also, the present specification discloses a radiation imaging apparatus including any one of the above-described radiation image detection apparatuses and a radiation source that irradiates radiation toward the first grating.

Further, in the present specification, in the image data acquired by the radiation imaging apparatus and the radiation image detector, the data of each pixel belonging to the defect area onto which the connecting portion is projected is displayed around the pixel. A radiation imaging system is disclosed that includes an arithmetic processing unit that complements data of pixels belonging to an area excluding the defective area.

Further, in the radiation imaging system disclosed in this specification, the arithmetic processing unit may calculate a refraction angle of radiation incident on the radiation image detector from image data in which data of each pixel belonging to the defect area is complemented. The distribution is calculated, and a phase contrast image of the subject is generated based on the distribution of the refraction angles.

According to the present invention, a plurality of lattice pieces are connected to form a lattice, thereby obtaining a large-size lattice while maintaining the accuracy by using a relatively small individual lattice piece. it can. Then, by forming the connecting portion of the adjacent lattice pieces with a radiation shield, the pixels belonging to the defect area can be easily and reliably extracted, and the data of the pixels can be complemented with certainty. The image quality can be improved.

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 22, 2010 (Japanese Patent Application No. 2010-260667), the contents of which are incorporated herein by reference.

10 X-ray imaging system 11 X-ray source 12 Imaging unit 13 Console 30 FPD
31 First absorption grating 31A First grating piece 32 Second absorption grating 32A Second grating piece 33 Scanning mechanism 40 pixels

Claims (12)

1. A first lattice;
   A grating pattern having a period substantially matching a pattern period of a radiation image formed by radiation that has passed through the first grating;
   A radiation image detector for detecting the radiation image masked by the lattice pattern;
   With
   The first grating includes a plurality of grating pieces arranged in a plane that intersects the traveling direction of the radiation passing through the first grating, and a connecting portion that connects adjacent grating pieces is a radiation shield. The formed radiographic image detection apparatus.
2. The radiological image detection apparatus according to claim 1,
   The lattice pattern is a second lattice;
   The second grating includes a plurality of grating pieces arranged in a plane that intersects the traveling direction of the radiation passing through the second grating, and a connecting portion that connects adjacent grating pieces is a radiation shield. The formed radiographic image detection apparatus.
3. The radiological image detection apparatus according to claim 2,
   The radiation image detector includes: a first region where the first lattice connection portion is projected and a second region where the second lattice connection portion is projected in a projection with a radiation focus as a viewpoint; And including
   The radiographic image detection apparatus in which the first area and the second area coincide with each other, or one of the first area and the second area is included in the other area.
4. The radiological image detection apparatus according to claim 2,
   In each of the first and second gratings, the plurality of grating pieces are arranged in a first direction,
   The radiation image detector includes: a first region where a connection portion between lattice pieces adjacent to each other in the first direction of the first lattice is projected in a projection with a radiation focus as a viewpoint; and the second region A second region where a connection portion of lattice pieces adjacent to each other in the first direction of the lattice is projected, and a third region excluding the first region and the second region,
   The radiation image detection apparatus in which at least one pixel belonging to the third region is interposed between each pixel belonging to the first region and each pixel belonging to the second region.
5. The radiological image detection apparatus according to claim 4,
   In each of the first and second gratings, the plurality of grating pieces are also arranged in a second direction intersecting the first direction,
   The radiological image detector includes: a fourth region where a connection portion between lattice pieces adjacent to each other in the second direction of the first lattice is projected in the projection with a radiation focus as a viewpoint; and the second region A fifth region where a connecting portion of lattice pieces adjacent to each other in the second direction of the lattice is projected, and a sixth region excluding the fourth region and the fifth region,
   The radiation image detection apparatus in which at least one pixel belonging to the sixth area is interposed between each pixel belonging to the fourth area and each pixel belonging to the fifth area.
6. The radiological image detection apparatus according to any one of claims 2 to 5,
   In each of the first and second gratings, a radiation image detection apparatus in which a surface on which the plurality of grating pieces are arranged is a cylindrical surface, and a central axis thereof passes through a radiation focus.
7. The radiological image detection apparatus according to any one of claims 1 to 5,
   The connection part is a radiographic image detection device formed of an adhesive in which a radiation absorbing material is dispersed.
8. The radiological image detection apparatus according to claim 7,
   The radiation image detecting apparatus, wherein the radiation absorbing material is particles of heavy metal having an atomic number of 40 or more.
9. The radiological image detection apparatus according to claim 8,
   The heavy metal is a radiological image detection apparatus which is at least one selected from the group consisting of gold, platinum and lead.
10. The radiological image detection apparatus according to any one of claims 1 to 9,
    A radiation source for irradiating radiation toward the first grating;
    A radiographic apparatus comprising:
11. The radiation imaging apparatus according to claim 10;
    In the image data acquired by the radiation image detector, the data of each pixel belonging to the defect area projected by the connecting portion is supplemented with the data of the pixels belonging to the area around the pixel and excluding the defect area. A radiation imaging system including an arithmetic processing unit.
12. It is a radiography system of Claim 11, Comprising:
    The calculation processing unit calculates a distribution of refraction angles of radiation incident on the radiation image detector from image data supplemented with data of each pixel belonging to the defect area, and based on the distribution of refraction angles, A radiography system that generates a phase contrast image of a subject.

Images