WO2012070662A1 - Radiographic image detection apparatus, radiography apparatus, and radiography system - Google Patents

Abstract

Provided are a radiographic image detector, radiography apparatus, and radiography system with which phase contrast images of stable image quality can be obtained using gratings that can be stably produced. The radiography system is provided with a first grating (31) having multiple bands (31b) disposed in a first direction (x direction), a second grating (32) having multiple bands (32b) disposed so as to mask periodically the radiographic image formed by radiation passing through the first grating (31), and a radiographic image detector (30) that detects, using multiple pixels (40), the radiographic image masked by the second grating (32). The first and second gratings (31, 32) each contain multiple grating pieces (311, 321) disposed at least in the first direction, x, in a plane intersecting with the optical axis of the radiation and the pitch P_C of grating piece (311, 321) arrangement in the x-direction is two or more times the pixel pitch P_D of the radiographic image detector (30) in the x-direction.

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, and a radiation imaging system that enable phase imaging of a subject using radiation such as X-rays.

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

In a general X-ray imaging system, a subject is placed between an X-ray source that emits X-rays and an X-ray image detector that detects X-rays, and a transmission image of the subject is taken. In this case, each X-ray emitted from the X-ray source toward the X-ray image detector is caused by a difference in characteristics (atomic number, density, thickness) of the substance existing on the path to the X-ray image detector. After receiving a corresponding amount of attenuation (absorption), it enters each pixel of the X-ray image detector. As a result, an X-ray absorption image of the subject is detected and imaged by the X-ray image detector. As an X-ray image detector, a flat panel detector (FPD: Flat Panel Detector) using a semiconductor circuit is widely used in addition to a combination of an X-ray intensifying screen and a film and a stimulable phosphor.

However, since the X-ray absorption ability is lower as a substance composed of an element having a smaller atomic number, a problem that a sufficient softness (contrast) of an X-ray absorption image cannot be obtained with a soft tissue or a soft material of a living body. There is. For example, most of the components of the cartilage part constituting the joint of the human body and the joint fluid in the vicinity thereof are water, and there is little difference in the amount of X-ray absorption between them, so that it is difficult to obtain a difference in light and shade.

Against the background of such problems, in recent years, X-ray phase imaging that obtains an image (hereinafter referred to as a phase contrast image) based on the X-ray phase change by the subject instead of the X-ray intensity change by the subject is proposed. There is a lot of research. 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 above X-ray Talbot interferometer, a first diffraction grating (phase type grating or absorption type grating) is arranged behind the subject, and a specific distance (Talbot) determined by the grating pitch of the first diffraction grating and the X-ray wavelength. The second diffraction grating (absorption type grating) is disposed downstream by the (interference distance), 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 placed subject and X-rays.

In the above X-ray Talbot interferometer, the phase of the subject is detected by detecting the moire caused by the superposition of the self-image of the first diffraction grating and the second diffraction grating, and analyzing the change in the moire caused by the subject. Get information. As a moire analysis method, 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. An angle distribution of X-rays refracted by the subject from a change in each pixel value obtained by performing X-ray imaging while performing translational movement in the vertical direction at a scanning pitch equally divided by the lattice pitch. This is a method for obtaining (differential image of phase shift), and a phase contrast image of the subject can be obtained based on this angular distribution.

The X-ray grating used for the phase imaging as described above needs to have a size corresponding to, for example, a viewing angle of several tens of centimeters from the size of the measurement target. Regarding this grid size, in Patent Document 2, a large number of modules in which the first grid, the second grid, and the X-ray image detector are integrated are necessary as a whole even if the individual module size is small. Realizes a large visual field size.

Japanese Unexamined Patent Publication No. 2008-200399 Japanese Laid-Open Patent Publication No. 2007-203061

Incidentally, an X-ray grating, particularly an absorption type grating, requires a high aspect ratio grating pattern in which the ratio of the thickness of the strip that shields X-rays is high with respect to the opening width of the grating. The opening width of the lattice and the thickness of the stripe are preferably set according to the energy of X-rays used for imaging, but the energy region of X-rays as used in normal medical image diagnosis is 30 to 30. Considering that it is 120 keV, even if the strip is made of gold having high X-ray absorption, a thickness of 10 to 300 μm is required for a lattice opening width of several μm. As a method capable of forming a high aspect ratio lattice pattern, there is an etching process in a photolithography technique. For example, a heavy metal pattern (strip) such as gold or platinum is formed on a substrate with a lattice pitch of several μm using deep etching. In addition to the difficulty of forming a high aspect ratio in this way, in addition to the difficulty of forming a lattice pattern in this way, in-plane uniformity on the order of submicron order is also required over a wide area corresponding to the visual field size. Manufacturing is very difficult.

For this reason, it is conceivable to arrange a plurality of small-sized lattice pieces to be used as a single lattice. However, there is a moire that is not necessary for measurement due to the correlation between the lattice pitch and the pixel pitch of the X-ray image detector. Can occur. That is, since the joint between the lattice pieces has an arrangement pitch with a period corresponding to the dimension of the lattice pieces, an image is obtained in the X-ray image detector by overlapping each periodicity of the arrangement pitch and the pixel pitch. Moire can occur during acquisition.

In the X-ray imaging system using the above-described X-ray Talbot interferometer, a self-image (hereinafter referred to as a self-image G1) formed by X-rays that have passed through the first grating is periodically generated by the second grating. To mask. When the first and second gratings are relatively scanned in the arrangement direction of the grating pattern, the intensity modulation obtained as the moire formed by the superposition of the self-image G1 and the second grating moves. The signal is detected by an X-ray image detector, and a phase contrast image is formed from the phase shift between the intensity modulation signal when the X-ray does not pass through the subject and the intensity modulation signal when the X-ray passes through the subject. To do.

In the case where a plurality of grid pieces are arranged and the first and second grids are respectively configured as described above, in addition to the moire generated in relation to the self-image G1 and the second grid, Moire generated by each cycle must be taken into account. That is, the moire generated due to the relationship between the self-image G1 and the second grating and the moire generated due to the relationship between the lattice piece and the pixel are superimposed on the detection surface of the X-ray image detector.

Here, in order to detect the intensity modulation signal, it is required that the frequency of the moire on the detection surface of the X-ray image detector and the frequency of the pixel period do not coincide with each other and do not have an integer multiple relationship with each other. . Further, in order to detect the phase shift of the intensity modulation signal with high sensitivity, it is necessary to obtain the intensity modulation signal with high contrast, and for this purpose, it is necessary to make the moire frequency lower than the frequency of the pixel period. .
In addition, the phase shift of the intensity modulation signal is obtained by comparing the intensity modulation signal formed from the moire image in the so-called pre-photographing when the subject is not present and the so-called main photographing when the subject is present. When the moire frequency fluctuates between pre-photographing and main-photographing, artifacts are generated and the image quality of the phase contrast image is degraded.
Note that Patent Document 2 mentions artifacts, but does not mention specific countermeasures.

As described above, from the principle of detecting the phase shift of the intensity modulation signal, it is required that the moire frequency does not deviate from a suitable value range that does not match the pixel frequency (not an integer multiple). That is, in order to stably obtain a phase contrast image having an image quality suitable for image measurement, it is important that the moire frequency is stable.

The inventor of the present invention diligently researched and found new knowledge that the moire frequency can fluctuate greatly with respect to the ratio change between the arrangement pitch of the lattice pieces and the pixel pitch. Such a change in the ratio between the arrangement pitch of the grating pieces and the pixel pitch is easily caused by, for example, a relative distance between the grating and the X-ray image detector due to vibration or temperature change, a relative rotation angle shift, and the like. The image quality is degraded.

As described above, an object of the present invention is to provide a radiation image detector, a radiation imaging apparatus, and a radiation imaging system that can obtain a stable phase contrast image using a grating that can be manufactured stably.

(1) A first grating having a plurality of strips arranged in a first direction, and a grating arranged so as to periodically mask a radiation image formed by radiation that has passed through the first grating. A radiation image detector that detects the radiation image periodically masked by the lattice pattern using a plurality of pixels, and the first lattice is radiation that passes through the first lattice. A plurality of first grid pieces arranged in the first direction in a plane intersecting the optical axis of the first grid piece, and the arrangement pitch of the first grid pieces in the first direction is the radiation image detection A radiological image detection apparatus characterized by being at least twice the pixel pitch in the first direction of the vessel.
(2) A first grating having a plurality of strips arranged in a first direction and a plurality arranged in such a manner as to periodically mask a radiation image formed by radiation that has passed through the first grating. A radiation image detector that detects, using a plurality of pixels, the radiation image periodically masked by the second lattice, and the second lattice comprises: A plurality of second grating pieces arranged in at least the first direction in a plane intersecting the optical axis of the radiation passing through the first grating and the second grating, and the second grating An arrangement pitch of the pieces in the first direction is at least twice as large as a pixel pitch in the first direction of the radiological image detector.
(3) A first grating having a plurality of strips arranged in a first direction and a plurality arranged in such a manner as to periodically mask a radiation image formed by radiation that has passed through the first grating. And a radiation image detector that detects the radiation image periodically masked by the second grating using a plurality of pixels, the first grating comprising: A plurality of first lattice pieces arranged in at least the first direction in a plane intersecting the optical axis of the radiation passing through the first lattice, and the first lattice piece includes the first lattice pieces. The arrangement pitch in the direction is at least twice the pixel pitch in the first direction of the radiation image detector, and the second grating is at least the first in a plane intersecting the optical axis of the radiation. A plurality of second grids arranged in a direction A radiation image detection apparatus including a piece, wherein an arrangement pitch of the second lattice pieces in the first direction is twice or more a pixel pitch of the radiation image detector in the first direction. .
(4) A radiation imaging apparatus comprising: the radiation image detection apparatus according to any one of (1) to (3) above; and a radiation source that irradiates radiation toward the first grating.
(5) The distribution of the refraction angle of the radiation incident on the radiation image detector is calculated from the radiation imaging device of (4) above and the image detected by the radiation image detector of the radiation imaging device, and this refraction is calculated. A radiation imaging system comprising: an arithmetic processing unit that generates a phase contrast image of a subject based on a distribution of angles.

According to the present invention, it is possible to stably manufacture a grating having a size corresponding to a required visual field size, and it is possible to stably produce a phase contrast without degrading the image quality of the phase contrast image due to the arrangement of the grating pieces. An image is obtained.

It is a side view which shows typically 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 of the 1st, 2nd grating | lattice. It is a side view of the 1st, 2nd grating | lattice. It is a schematic diagram which shows the mechanism for changing the period of the interference fringe (moire) by interaction of the 1st and 2nd grating | lattice. It is a side view which shows the 1st, 2nd grating | lattice, and FPD typically. It is a graph which shows the characteristic about the fluctuation | variation of the moire frequency with respect to ratio of a pixel pitch and arrangement pitch. It is a graph which shows the fluctuation | variation characteristic of the moire frequency of a harmonic component. It is a schematic diagram which shows the dimensional relationship between the arrangement pitch of a 1st, 2nd lattice piece, and a pixel pitch. It is a schematic diagram geometrically showing the upper limit of the arrangement pitch. 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 top view which shows typically the grating | lattice which has the some grating | lattice piece arranged in the x direction and the y direction. 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 top view which shows a two-dimensional lattice typically. 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 figure which shows the structure of the radiographic image detector regarding the other example of the radiography system for describing embodiment of this invention. It is a schematic 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 a CMOS sensor. It is a figure which shows the structure of one pixel circuit of the radiographic image detector using a CMOS sensor. It is a figure which shows the arrangement | positioning relationship between the radiographic image detector using a CMOS sensor, 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 an example of the grating | lattice which made the grating | lattice surface into the concave surface. 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.

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 H in a standing position, and the subject H is interposed between the X-ray source 11 that irradiates the subject H with X-rays and the X-ray source 11. In this state, the radiographing unit 12 is disposed as opposed to the X-ray source 11 and detects X-rays transmitted through the subject H from the X-ray source 11 to generate image data. Based on this, the console 13 (FIG. 2) that controls the exposure operation of the X-ray source 11 and the imaging operation of the imaging unit 12, and generates a phase contrast image by processing the image data acquired by the imaging unit 12. Broadly divided.

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 as a radiological image detector made of a semiconductor circuit, a first absorption grating 31 for detecting phase change of X-rays by the subject H and performing phase imaging, and a first absorption grating 31. Two absorption gratings 32 are provided.

The imaging unit 12 includes scanning means 33 that relatively moves the first absorption grating 31 and the second absorption grating 32 by translating the second absorption grating 32 in the vertical direction (x direction). 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.

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

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

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

Each pixel 40 once converts X-rays into visible light 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 first and second absorption type gratings 31 and 32. FIG.

The first absorption type grating 31 is configured by arranging a plurality of first grating pieces 311. Each of the first grating pieces 311 includes a substrate 31a and a plurality of strips arranged on the substrate 31a. It has an X-ray shielding part 31b as a band. The second absorptive grating 32, which is a grating pattern, is also configured by arranging a plurality of second grating pieces 321. Each of the second grating pieces 321 is arranged on the substrate 32a and the substrate 32a. And a plurality of X-ray shielding portions 32b as a plurality of strips. The substrates 31a and 32a are both made of an X-ray transparent member such as glass that transmits X-rays. Both the first and second grating pieces 311 and 321 are arranged in the x direction as the first direction in a plane orthogonal to the optical axis A of the X-ray (in the xy plane).

Each of the X-ray shielding portions 31b and 32b is one direction in the plane orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11 (one of the x direction and the y direction). In the example, it is composed of a linear member extending in the y direction). As a material of each X-ray shielding part 31b, 32b, a material excellent in X-ray absorption is preferable, and for example, a heavy metal such as gold or platinum is preferable. These X-ray shielding portions 31b and 32b can be formed by a metal plating method or a vapor deposition method.

The X-ray shielding part 31b has a constant lattice pitch p ₁ in a direction orthogonal to the one direction (x direction as the first direction in the example of FIG. 4) in a plane orthogonal to the optical axis A of the X-ray. Are arranged at a predetermined interval d ₁ from each other. Similarly, the X-ray shielding portion 32b, in a plane perpendicular to the optical axis A of the X-ray, with grating pitch p ₂ of the constant in the x-direction, are arranged at a predetermined interval d ₂ from each other.
Such first and second absorption gratings 31 and 32 do not mainly give a phase difference to incident X-rays but give an intensity difference, and 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.

Since the first and second absorption type gratings 31 and 32 include a plurality of grating pieces 311 and 321, respectively, the grating pieces 311 are manufactured when the first and second absorption type gratings 31 and 32 are manufactured. , 321 units. The first and second absorptive gratings 31 and 32 are required to have a so-called high aspect ratio in which the thickness of the X-ray shielding portion is large with respect to the opening width of the grating, and have a high degree of in-plane uniformity over an area corresponding to the field size. However, since it is difficult to stably manufacture the structure, it is possible to manufacture in units of lattice areas having a small area by using the configuration of this example. The mold lattices 31 and 32 can be easily manufactured and can be stably manufactured.

The first absorption type grating 31 is configured to geometrically project X-rays that have passed through the slit portion regardless of the presence or absence of the Talbot interference effect. Specifically, by setting the interval d ₁ to a value sufficiently larger than the effective wavelength of the X-rays emitted from the X-ray source 11, most of the X-rays irradiated do not undergo diffraction at the slit portion. In addition, a self-projected image (hereinafter also referred to as a self-image G1) can be formed behind the first absorption grating 31. For example, when tungsten is used as the target material of the rotary anode 18a and the tube voltage is 50 kV, the effective wavelength of X-ray is about 0.4 mm. In this case, if the distance d ₁ is set to about 1 to 10 μm, the X-ray image formed by the X-rays that have passed through the slit portion is negligible for the diffraction effect. A self-image G1 is formed behind.

In general, the X-ray radiated from the X-ray source 11 is not a parallel beam but a cone beam with the X-ray focal point 18b as a light emission point, and the self-image G1 is at a distance from the X-ray focal point 18b. Scaled proportionally. On the other hand, the grating pitch p ₂ of the second absorption type grating 32 is determined so that the slit part substantially coincides with the periodic pattern of the bright part of the self-image G 1 at the position of the second absorption type grating 32. ing. That is, when the distance from the X-ray focal point 18b to the first absorption-type grating 31 is L ₁ and the distance from the first absorption-type grating 31 to the second absorption-type grating 32 is L ₂ , the second The pitch p ₁ ′, the first grating pitch p ₁ , and the second grating pitch p ₂ of the self-image G1 at the position of the absorption grating 32 are 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 X-ray imaging system 10 has a configuration in which the first absorption grating 31 projects incident X-rays without diffracting the self-image G1 of the first absorption grating 31. because similarly obtained behind the position of the first absorption-type grating 31, 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. 31 grating pitch p ₁ , second absorbing grating 32 grating pitch p ₂ , X-ray wavelength (usually the effective wavelength of X-rays incident on first absorbing grating 31) λ, and positive integer m. And is represented by the following formula (2).

Expression (2) is an expression that represents the Talbot interference distance when the X-ray irradiated from the X-ray source 11 is a cone beam. “Atsushi Momose, et al., Japan Journal of Applied Physics, Vol. 47, No. 10, October 2008, page 8077 ”and“ Timm Weitkamp, et al., Proc. Of SPIE, Vol. 6318, 2006, 63180S-1 ”can be easily derived.

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).

In the case where X-rays emitted from the X-ray source 11 can be regarded as substantially parallel beams, Talbot interference distance Z by the following equation (4), and the distance L _2, the range satisfying the following equation (5) Can be set to a value.

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. The X-ray shielding portions 31b and 32b preferably shield 90% or more of the irradiated X-rays, and the thickness thereof is set according to the energy of the irradiated X-rays. For example, when tungsten is used as the target material of the X-ray tube 18 and the tube voltage is 50 kV, the thicknesses h ₁ and h ₂ are preferably 100 μm or more in terms of gold (Au).

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. For this reason, it is preferable to limit the thicknesses h ₁ and h ₂ from the viewpoint of securing a visual field. Specifically, if the length V of the effective visual field in the x direction on the detection surface of the FPD 30 and the distance from the X-ray focal point 18b to the detection surface of the FPD 30 are L, the thicknesses h ₁ and h ₂ are as shown in FIG. It is preferable to set so that following Formula (6) and (7) may be satisfy | filled from a scientific relationship.

For example, when d ₁ = 2.5 μm and d ₂ = 3.0 μm and L = 2 m is set as a size that can be installed in a normal hospital, the effective visual field length V in the x direction is 10 cm long. In order to ensure the thickness, 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, a moiré is formed by superimposing the self-image G1 of the first absorption type grating 31 and the second absorption type grating 32, and is imaged by the FPD 30. There is a slight difference between the pattern period p ₁ ′ of the self-image G1 at the position of the second absorption type grating 32 and the substantial grating pitch p ₂ ′ of the second absorption type grating 32 due to manufacturing errors and arrangement errors. Arise. 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.

Moire occurs due to a minute difference between the pattern period p ₁ ′ of the self-image G1 and the grating pitch p ₂ ′. The moire period T is expressed by the following equation (8).

The moiré period T in the above equation (8) is actually further expanded by the distance from the second absorption type grating 32 to the detection surface of the FPD 30, so the moiré period on the detection surface of the FPD 30 is T ′. , in order to detect the moire FPD 30, the arrangement pitch P _D x direction of the pixel 40 is required to be not an integral multiple of the moire cycle T 'on the detection surface of at least FPD 30, the following equation ( 9) must be satisfied (where n is a positive integer).

Further, in the range satisfying equation (9), 'it is possible to detect the moire be greater than the arrangement pitch P _D moire cycle T' arrangement pitch P _D moire cycle T preferably smaller than, It is preferable to satisfy the following formula (10). This is because, in order to obtain a high-quality phase contrast image, it is preferable that moire is detected with high contrast in the phase contrast image generation process described later.

The arrangement pitch P _D of the pixels 40 of FPD30 is the magnitude relation between it is difficult to change a design to a defined value (usually about 100 [mu] m), the arrangement pitch P _D and the moire cycle T ' For the adjustment, the positions of the first and second absorption gratings 31 and 32 are adjusted, and the moire pattern is changed by changing at least one of the pattern period p ₁ ′ and the grating pitch p ₂ ′ of the self-image G1. It is preferable to change the period T.

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 one of the first and second absorption gratings 31 and 32 is relatively centered about an axis perpendicular to the optical axis A and along the y direction. 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 yet 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 self-image of the first absorption type grating 31 projected on the position of the second absorption type grating 32. The pattern period of G1 changes as “p ₁ ′” → “p ₁ ′ × (L ₁ + L ₂ + δ) / (L ₁ + L ₂ )”, 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 type gratings 31 and 32 for changing the moire period T ′ is an actuator such as a piezoelectric element. It is possible to configure.

When the subject H is arranged between the X-ray source 11 and the first absorption type grating 31, the moire detected by the FPD 30 is 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 detected by the FPD 30.

FIG. 7 is a side view schematically showing the first and second absorption type gratings 31 and 32 and the FPD 30.

Adjacent first lattice pieces 311 and 311 are connected by adhesion, fusion, pressure bonding or the like. Note that the first lattice piece 311 shown in this example is an independent small piece having a substrate 31a and an X-ray shielding part 31b for each lattice piece 311, but the first lattice piece 311 has 1 It may be formed by sequentially exposing and etching each of a plurality of regions partitioned on a single substrate. Between the adjacent first grid pieces 311 and 311, a boundary portion 315 which is a joint of the grid pieces by adhesion, fusion, pressure bonding or the like is formed in a band shape in a plan view.

Adjacent second grid pieces 321 and 321 are also connected by adhesion, fusion, pressure bonding, or the like, similarly to the first grid pieces 311 and 311. The second lattice piece 321 may also be formed by sequentially performing exposure, etching, and the like on each of a plurality of regions partitioned on one substrate. A band-shaped boundary portion 325 is also formed between the adjacent second lattice pieces 312 and 312 in plan view.

Even if the boundary portions 315 and 325 of the first and second lattice pieces 311 and 321 are X-ray shielding regions using an X-ray shielding material, X-ray transmission is possible like glass and general adhesives. It may be a region, or may be a region in which an X-ray shielding part and an X-ray transmission part are mixed. The boundary portions 315 and 325 in this example are grasped as band-like regions, but the boundary portions of the lattice pieces are not necessarily band-like. For example, when a plurality of lattice pieces are arranged by divided exposure and divided etching on a plurality of areas partitioned on a single substrate, or when a plurality of lattice pieces are connected by adhesion, the boundary portion is a strip shape. However, in the case where the lattice pieces are connected by fusion, pressure bonding, or the like, the boundary portion may be a random shape such as an intermittent line shape instead of a strip shape.

Because of these boundary portions 315 and 325, the X-ray dose transmitted through each of the first absorption-type grating 31 and the second absorption-type grating 32 varies depending on where X-rays are transmitted. Such boundary portions 315 and 325 have periodicity at a constant arrangement pitch according to the respective dimensions of the first and second lattice pieces 311 and 321. That is, the first and second gratings 31 and 32 are configured by arranging a plurality of grating pieces, respectively, in addition to the periodicity of the pitches p ₁ and p ₂ of the X-ray shielding portions 31b and 32b, The first and second grating pieces 311 and 321 have periodicity.

For this reason, the arrangement pitch of the lattice pieces is added to the moire generating element detected by the FPD 30. That is, on the detection surface on which the pixels of the FPD 30 are arranged, the pattern period p ₁ ′ of the self-image G1 of the first absorption grating 31, the grating pitch p _{2 of} the second absorption grating 32, the first and Depending on the relative relationship between the arrangement pitch of each of the second lattice pieces 311 and 321 and the arrangement pitch (pixel pitch) of the pixels 40 of the FPD 30 in the x direction, moiré of several types or a plurality of different periods is superimposed.

At this time, the arrangement pitch of the first and second grating pieces 311 and 321 on the detection surface of the FPD 30 is set so that each of the first and second grating pieces 311 and 321 has a parallel beam when the X-ray is a parallel beam. Although the actual arrangement pitches P _B1 and P _B2 , the X-ray in this example is a cone beam whose irradiation range is expanded in proportion to the distance from the focal point 18b, and therefore the distance from the focal point 18b to each grating, The arrangement pitches P _C1 and P _C2 corrected by multiplying the actual arrangement pitches P _B1 and P _B2 by the magnification ratio, which is the ratio between the focal point 18b and the distance to the FPD 30, respectively, are the first and second gratings at the position of the FPD 30. It becomes the arrangement pitch of each of the pieces 311 and 321. The enlargement factor used for correcting the arrangement pitch P _B1 is L / L1 with reference to FIG. 5, and the enlargement factor used for correcting the arrangement pitch P _B2 is L / (L1 + L2) with reference to FIG.

By the way, in the image detected by the FPD 30, in the portions corresponding to the boundary portions 315 and 325 of the first and second grating pieces 311 and 321, the self-image G1 and the grating pattern of the second absorption grating 32 are Moire due to overlaying cannot be obtained. Although the image data of these boundary portions can be complemented based on the image data of the adjacent portions, the boundary portion 315 and the boundary portion 325 serve as detection surfaces of the FPD 30 in order to minimize the image range to be complemented. Are preferably overlapped. For this reason, the corrected arrangement pitches P _C1 and P _C2 are determined to be equal to each other. Hereinafter, it may be collectively referred to as the arrangement pitch _{P C} that arrangement pitch _{P C1, P C2.}

Incidentally, equally-determined in the arrangement pitch _{P C1, P C2,} manufacturing errors, can tolerate pitch difference of several μm or less by an assembly error or the like, the arrangement pitch _{P C1, P C2,} the range the difference is less than a few μm Is almost equal within. The small differences in such arrangement pitch _{P C1, P C2,} also superimposed moire and arrangement pitch _{P C1} and the arrangement pitch _{P C2} is related to the detection surface of the FPD 30.

Moire frequency for the ratio of the pixel pitch P _D and the arrangement pitch P _C1 is possible to calculate analytically. FIG. 8 shows the fluctuation characteristics of the moire frequency calculated analytically for the moire detected by the FPD 30. Moire frequency shown in FIG. 8 is a frequency of the fundamental wave of the moire, half the sampling frequency f _S by the detector pixel is a Nyquist frequency.
In the analysis of FIG. 8, the array pitches P _C1 and P _C2 corrected with the enlargement ratio are used. Figure 8 is a sequence pitch equal-determined one another of the first and second grating pieces 311, 321 marked with _{P C.}

From Figure 8, when the value of the arrangement pitch P _C is close to the value of the pixel pitch P _D, the moire frequency f _m varies greatly with respect to subtle changes in the ratio of the pixel pitch P _D and the arrangement pitch P _C, the If the ratio is increased to 2 or more, variations of the moire frequency f _m it can be seen that toward the convergence. Change in the ratio of such arrangement pitch P _C and the pixel pitch P _D is, for example, displacement of the relative distance and the relative rotation angle between the first and second absorption type gratings 31 and FPD30 due to vibration or a temperature change This easily occurs and degrades the image quality of the phase contrast image.

Moire generation elements detected by the FPD 30 include the respective grating pitches p ₁ and p ₂ of the first and _second absorption type gratings 31 and 32, and the respective arrangements of the first and second grating pieces 311 and 321. pitch _{P C1, P C2,} and includes a pixel pitch _{P D.} Here, the grating pitch p _1, p ₂ is about several [mu] m, the pixel pitch P _D is about 100μm for a large period difference of each other, detection of moire occurring in relation to these grating pitch and the pixel pitch The intensity is small enough not to affect the image measurement.

On the other hand, with respect to the arrangement pitch P _C of larger lattice piece than the grating pitch p _1, p _2, moire is generated in relation to the arrangement pitch P _C and the pixel pitch P _D, in particular, the arrangement pitch P _C and the pixel pitch P _{When D} is close, strong moire can be detected. Furthermore, as shown in FIG. 8, it is considered that varies greatly even moire frequency f _m with respect to the ratio between the arrangement pitch P _C and the pixel pitch P _D.

Also, there are cases where it is not sufficient to simply consider the fundamental frequency shown in FIG. That is, when an image detector having a high spatial frequency response up to a high frequency is used, it is necessary to consider the moire harmonics that fold back within the spatial frequency band of the detector. A typical example of such an image detector is a direct conversion type FPD. In the direct conversion type FPD, the X-rays incident on the detector are directly converted into electric signals. Therefore, the direct conversion FPD is converted into an electric signal using a photoelectric conversion element such as a photodiode after being converted into light once by a scintillator or the like. It is generally known that there is less spatial blur in the signal detection process compared to a conversion type FPD, and a high spatial frequency response is exhibited over a high frequency.

9, each of the moire frequency component f _m from the first-order fundamental wave to the fourth harmonic, showing a variation characteristic of the ratio between the in arrangement pitch P _C and the pixel pitch P _D. More and more is the order of the moire frequency f _m, the variation of the ratio between the in arrangement pitch P _C and the pixel pitch P _D increases. Further, there is more, the ratio of the pixel pitch P _D when variations of the moire frequency f _m is directed to converge arrangement pitch P _C is increased at higher order. Moire frequency of the fourth harmonic, the arrangement pitch P _C is directed to converge becomes more than 6 times the pixel pitch P _D.

By the way, in order to detect an intensity modulation signal, the moire frequency on the detection surface of the FPD 30 and the frequency of the pixel period do not coincide with each other based on the above equation (9) and are not in an integral multiple relationship with each other. Is requested. Further, in order to detect the phase shift of the intensity modulation signal with high sensitivity, it is necessary to obtain the intensity modulation signal with a high contrast. For this reason, it is necessary to make the moire frequency lower than the frequency of the pixel period. In order to satisfy such a condition, there is a range of suitable values for the moire frequency.

Further, as will be described later, the phase contrast image is obtained from the amount of phase shift of the intensity modulation signal formed from the moire image between the pre-photographing without placing the subject and the main photographing with the subject, so the moire frequency As a result of fluctuations, artifacts occur. That is, in order to stably obtain a phase contrast image having an image quality suitable for image measurement, it is important that the moire frequency is stable.

8, as shown in FIG. 9, the moire caused by the relationship between the arrangement pitch P _C and the pixel pitch P _D is likely to greatly change the frequency of the moiré detected by the FPD 30. For example, if the ratio between the pixel pitch and the array pitch changes due to vibration, temperature change, etc., the moire frequency is out of the preferred range and does not satisfy Equation (9), and the intensity modulation signal cannot be detected, or the intensity modulation signal There is a possibility that the detection sensitivity of the phase shift is lowered. Further, since the relative position of the X-ray focal point 18b and the first and second absorption type gratings 31 is magnification of an image on the detection surface of FPD30 displaced in the optical axis direction is changed, the pixel pitch P _D the ratio between the arrangement pitch P _C and changes. As a result, a moire of the spatial frequency corresponding to the ratio is generated. Thus, the spatial frequency is dependent on the relative displacement between the X-ray focal point 18b and each of the first and second absorption gratings 31 and 32. It is very difficult to correct moire that fluctuates.

8, the variation characteristic of the moire frequency shown in FIG. 9, by increasing the arrangement pitch P _C to the pixel pitch P _D, it is possible to stabilize the moire frequency.

Figure 10 illustrates the dimensional relationship between the arrangement pitch P _C and the pixel pitch P _D schematically. In this example, the resolution of the FPD 30, and the like in consideration of the moiré frequency variation characteristic of up to the fourth harmonic in FIG. 9, the arrangement pitch P _C by a more than six times the pixel pitch P _D in the x-direction. Arrangement pitch P _C is preferably 8 or more times the pixel pitch P _D in the x-direction, more preferably 10 times or more, more preferably is 100 times or more. However, the arrangement pitch _{P C} is response performance to the spatial frequency of the FPD 30, typically determined by MTF (Modulation Transfer Function), etc., it may be at least twice the pixel pitch _{P D.}

Incidentally, in FIG. 10, the arrangement pitch P _C where each array pitch of the first and second grating pieces 311 and 321 are equal, but showing the relationship between the pixel pitch P _D, first and second If each pitch of the grating pieces 311 and 321 are different, their arrangement pitch, respectively, it may be at least twice the pixel pitch P _D.

On the other hand, an upper limit can be set for the arrangement pitch in consideration of vignetting of X-rays. Although the formulas (6) and (7) have been described with reference to FIG. 5, the following formulas (11) and (12) are similarly defined from the geometrical relationship shown in FIG.

The dimension of the first absorption-type grating 31 in the direction along the optical axis A is h ₁ , the distance between adjacent X-ray shielding portions 31 b and 31 b, that is, the opening width is d ₁ , and the grating in the first absorption-type grating 31. P _B1 the actual pitch of the _strip, and the distance between the first absorption type grating 31 and the X-ray focal point 18b and R _1, it is preferable to satisfy the expression (11). Further, the dimension of the second absorption type grating 32 in the direction along the optical axis A is h ₂ , the distance between the adjacent X-ray shielding parts 32 b and 32 b, that is, the opening width is d ₂ , and the second absorption type grating 32. It is preferable that Expression (12) is satisfied, where P _B2 is the actual arrangement pitch of the lattice pieces and R ₂ is the distance between the second absorption grating 32 and the X-ray focal point 18b.

By determining the arrangement pitch as described above, the moire frequency can be stabilized against fluctuations in the ratio between the arrangement pitch and the pixel pitch. Here, each of the arrangement pitch _{P C1, P C2} of the first and second grating pieces 311 and 321, so that at least twice the pixel pitch _{P D,} it may be a numerical value away from the pixel pitch _{P D} Therefore, strict dimensional control of the arrangement pitch is unnecessary. Therefore, the first and second absorption type gratings 31 and 32 configured by arranging the first and second grating pieces 311 and 321 can be easily manufactured.

Next, the moire analysis method will be described.

FIG. 12 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 formula (13), 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 self-image G1 projected from the first absorption grating 31 to the position of the second absorption 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. It will be. This amount of displacement Δx is approximately expressed by the following equation (14) based on the small X-ray refraction angle φ.

Here, the refraction angle φ is expressed by Expression (15) using the X-ray wavelength λ and the phase shift distribution Φ (x) of the subject H.

Thus, the displacement amount Δx of the self-image G1 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 (16).

Therefore, by obtaining the phase shift amount ψ of the signal of each pixel 40, the refraction angle φ is obtained from the equation (16), and the differential amount of the phase shift distribution Φ (x) is obtained using the equation (15). 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, one of the first and second absorption type gratings 31 and 32 is translated in a stepwise manner in the x direction relative to the other, that is, the phase between the two grating patterns is changed. While shooting. In the present X-ray imaging system 10, the second absorption type grating 32 is moved by the scanning means 33 described above, but the first absorption type grating 31 may be moved. With the movement of the second absorption type grating 32, the moire moves, and the translational distance (movement amount in the x direction) is one period of the grating period of the second absorption type grating 32 (grating pitch p ₂ ), That is, when the phase change reaches 2π, the moire returns to the original position. A fringe image is photographed by the FPD 30 while moving the second absorption type grating 32 by an integer of the grating pitch p ₂ by such a change in moire, and the signal of each pixel 40 from the plurality of photographed stripe images. And the arithmetic processing unit 22 performs arithmetic processing to obtain the phase shift amount ψ of the signal of each pixel 40.

FIG. 13 schematically shows a state in which the second absorption type grating 32 is moved by the scanning pitch (p ₂ / M) obtained by dividing the grating pitch p ₂ into M (integers of 2 or more).

The scanning unit 33 sequentially translates the second absorption type grating 32 to M scanning positions where k = 0, 1, 2,..., M−1. In the figure, the initial position of the second absorption type grating 32 is set such that the dark part of the self-image G1 at the position of the second absorption type grating 32 when the subject H does not exist is almost at the X-ray shielding part 32b. Although the matching position (k = 0) is assumed, this initial position may be any position among 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 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 (17).

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 (18), the refraction angle φ (x) is expressed as the following expression (19).

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 (19).

FIG. 14 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. 14 indicates a change in the signal value when the subject H does not exist, and a solid line in FIG. 14 indicates a change in the 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 (15), the phase shift is obtained by integrating the refraction angle φ (x) along the x-axis. A distribution Φ (x) is obtained.

The above calculation is performed by the calculation processing unit 22, and the calculation processing unit 22 stores the phase contrast image in the storage unit 23. In the above description, the y coordinate in the y direction of the pixel 40 is not taken into consideration. However, by performing the same calculation for each y coordinate, a two-dimensional phase shift distribution Φ (x , Y). The above calculation is performed by the calculation processing unit 22, and the calculation processing unit 22 stores the calculated phase shift distribution Φ (x, y) in the image storage unit 23 as a phase contrast image.

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, since the moire frequency is stable, it is possible to prevent the occurrence of artifacts due to fluctuations in the moire frequency when the subject H exists and when the subject H does not exist. From the above, the first and second absorption- type gratings 31, 32 are configured by arranging a plurality of grating pieces 311, 321 that are smaller than the visual field size, respectively, thereby forming the first and second absorption- type gratings 31, 32. 32 can be manufactured stably, and the fluctuation of the moire frequency that can be caused by the arrangement of the lattice pieces 311 and 321 is stabilized by defining the ratio between the arrangement pitch and the pixel pitch, so that the stable image quality can be obtained. It is possible to obtain a phase contrast image.

Further, since most of the X-rays are not diffracted by the first absorption type grating 31 and geometrically projected onto the second absorption type grating 32, high spatial coherence is required for the irradiated X-rays. Instead, a general X-ray source used in the medical field can be used as the X-ray source 11. The distance L ₂ from the first absorption type grating 31 to the second absorption type grating 32 can be set to an arbitrary value, and the distance L ₂ is smaller than the minimum Talbot interference distance in the Talbot interferometer. Since it can be set, the photographing unit 12 can be downsized (thinned). Furthermore, in the present X-ray imaging system, almost all wavelength components of irradiated X-rays contribute to the projected image (self-image G1) from the first absorption grating 31, and the moire contrast is improved. Contrast image detection sensitivity can be improved.

Note that the X-ray imaging system 10 performs a fringe scan on the projection image of the first grating to calculate the refraction angle φ. Therefore, both the first and second gratings are absorption type. Although described as being a lattice, the present invention is not limited to this. As described above, when the refraction angle φ is calculated by performing fringe scanning on the Talbot interference image, that is, the Talbot interferometer in which the distance between the first grating and the second grating is set to the Talbot distance. The present invention is also useful when used. Therefore, the first grating is not limited to the absorption type grating but may be a phase type grating.

The analysis method of the moire 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. No. 1 (1982) p. 156 ”, a method using Fourier transform / inverse Fourier transform, and various methods using moire can be applied.
Here, in the Fourier transform method, the moire is directly analyzed by Fourier transform without acquiring an intensity modulation signal by scanning from the moire by the first and second gratings as in the fringe scanning method. Also in such a Fourier transform method, the stability of the moire frequency is important, and the same effect as described above can be obtained.

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, the phase shift distribution Φ corresponds to the refraction angle φ as described above. The differential amount of the phase shift distribution Φ is integrated, and the differential angle 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.

Also, a phase differential image (differential amount of phase shift distribution Φ) is created from a group of images acquired by shooting (pre-shooting) in the absence of a subject, and acquired by shooting (main shooting) in the presence of a subject. A phase differential image created from a group of images may be corrected. The phase differential image acquired by the pre-imaging reflects the phase unevenness inherent to the measurement system, for example, the grating pitch and thickness non-uniformity, and the grating scanning pitch error. On the other hand, the phase differential image created from the group of images acquired by shooting in the presence of the subject (main shooting) also contains phase irregularities unique to the measurement system of the same type as pre-photographing. Acting as an offset. Accordingly, by subtracting the phase differential image obtained by the pre-photographing from the phase differential image obtained by the main imaging, a phase differential image in which the phase unevenness of the measurement system is corrected can be obtained.

FIG. 15 schematically shows a configuration example in which the lattice pieces are arranged in the y direction as the second direction in addition to the x direction as the first direction. Both the first and second gratings are preferably configured to have grating pieces 131 arranged two-dimensionally in the x and y directions as shown in FIG. The lattice pieces 131 are arranged in the x direction via the boundary portion 31cx, and are arranged in the y direction via the boundary portion 31cy. Since the lattice pieces are also arranged in the y direction in this way, the size of the lattice pieces in the y direction is short, so that the lattice can be easily manufactured in terms of in-plane uniformity required for the manufacture of the lattice. . When each of the first and second gratings is composed of a plurality of grating pieces, it is preferable to use the grating pieces arranged two-dimensionally as shown in FIG.

In each of the above examples, the first and second gratings are both constituted by a plurality of grating pieces. However, only the first grating or only the second grating is constituted by a plurality of grating pieces. It may be.

FIG. 16 shows another example of a radiation imaging apparatus for explaining an embodiment of the present invention.

A mammography apparatus 80 shown in FIG. 16 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. 17 shows a modification of the radiation imaging apparatus of FIG.

The mammography apparatus 90 shown in FIG. 17 is different from the mammography apparatus 80 described above in that the first absorption type grating 31 is disposed between the X-ray source 11 and the compression plate 84. The first absorption type lattice 31 is accommodated in a lattice accommodation portion 91 connected to the arm member 81. The imaging unit 92 includes an FPD 30, a second absorption 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 projected image (self-image G1) of the first absorption grating 31 is deformed by the subject B. Accordingly, even in this case, the moiré 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. 18 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

18 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 self-image G1 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 degraded. 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 equation (20), where L ₃ is the distance from the multi-slit 103 to the first absorption-type lattice 31.

In the above equation (20), the projection image (self-image G1) 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 second absorption type grating 32. This is a geometrical condition for matching (overlapping) at the positions.

In addition, since the position of the multi-slit 103 is substantially the X-ray focal position, the first absorption grating 31 is given by assuming that the pattern period of the self-image G1 at the position of the second absorption grating 32 is p ₁ ′. The lattice pitch p ₁ of the second absorption lattice 32 and the lattice pitch p ₂ of the second absorption lattice 32 are determined so as to satisfy the relationship of the following equation (21).

Further, when the X-ray irradiated from the X-ray source 11 is a cone beam, in order to secure the effective field length V in the x direction on the detection surface of the FPD 30, from the multi slit 103 to the detection surface of the FPD 30. When the distance is L ′, the thicknesses h ₁ and h ₂ of the X-ray shielding portions 31b and 32b of the first and second absorption gratings 31 and 32 are determined so as to satisfy the following expressions (22) and (23). Is done.

Thus, in this example, the self-image G1 formed by each of the plurality of small focus light sources formed by the multi-slit 103 is 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 above examples.

FIG. 19 shows another example of a radiation imaging system for explaining an embodiment of the present invention.

As described above, the phase contrast image is based on the X-ray refraction component in the periodic array direction (x direction) of the X-ray shielding portions 31b and 32b of the first and second absorption gratings 31 and 32. For this reason, the refraction | bending component of the extending | stretching direction (y direction) of X-ray shielding part 31b, 32b is not reflected. That is, a part outline along a direction intersecting the x direction (or the y direction when orthogonal) is drawn as a phase contrast image based on a refractive component in the x direction via a lattice plane that is an xy plane. A part contour that does not intersect the direction and extends along the x direction is not drawn as a phase contrast image in the x direction. That is, there is a part that cannot be drawn depending on the shape and orientation of the part to be the subject H. For example, when the direction of the load surface of an articular cartilage such as a knee is matched to the y direction of the x and y directions which are the in-plane directions of the lattice, the region contour in the vicinity of the load surface (yz surface) substantially along the y direction is Although it is sufficiently depicted, it is considered that rendering of the cartilage peripheral tissues (tendons, ligaments, etc.) that intersect the load surface and extend substantially along the x direction is insufficient. By moving the subject H, it is possible to re-photograph a part that is not sufficiently drawn, but in addition to increasing the burden on the subject H and the operator, ensuring position reproducibility with the re-captured image There is a problem that is difficult.

Therefore, as another example, as shown in FIG. 19, the first line is centered on a virtual line (X-ray optical axis A) orthogonal to the center of the lattice plane of the first and second absorption gratings 31 and 32. The second absorption gratings 31 and 32 are integrally rotated at an arbitrary angle from the first direction shown in FIG. 19A (the direction in which the X-ray shielding portions 31b and 32b extend are along the y direction). Then, a rotation mechanism 105 is provided that has a second direction (a direction in which the extending direction of the X-ray shielding portions 31b and 32b extends along the x direction) shown in FIG. 19B, and the first direction and the second direction. It is also preferable to configure so that a phase contrast image is generated in each of the above. By doing so, the above-described problem of position reproducibility can be eliminated. FIG. 19A shows the first orientation of the first and second gratings 31 and 32 such that the extending direction of the X-ray shielding portions 31b and 32b is in the direction along the y direction. In b), the second of the first and second gratings 31 and 32 is rotated 90 degrees from the state of FIG. 19A and the extending direction of the X-ray shielding portions 31b and 32b is in the direction along the x direction. However, the rotation angles of the first and second gratings are arbitrary. Further, in addition to the first direction and the second direction, a phase contrast image in each direction is generated by performing two or more rotation operations such as the third direction and the fourth direction. May be.

In the illustrated example, the first and second absorption gratings 31 and 32 are rotated by 90 °, and the first direction and the second direction are orthogonal to each other. As long as the direction intersects, the rotation angle of the first and second absorption gratings 31 and 32 is not limited to 90 °. Further, the grating rotating mechanism 105 may be configured to rotate only the first and second absorption type gratings 31 and 32 separately from the FPD 30, or the first and second absorption type gratings 31. , 32 and the FPD 30 may be rotated together. Further, when the multi-slit 103 is provided, the multi-slit 103 and the collimator 109 or the radiation source formed integrally with them is rotated so that the rotation coincides with the first and second absorption gratings 31 and 32. . Furthermore, the generation of phase contrast images in the first and second orientations using the grating rotation mechanism 105 can be applied to any of the X-ray imaging systems described above.

FIG. 20 shows an example of a two-dimensional lattice in which X-ray shielding portions are arranged in two directions, the x direction and the y direction. Each lattice piece 231 has an X-ray shielding part 31bx arranged in the x direction and an X-ray shielding part 31by arranged in the y direction. The lattice pieces 231 are arranged in two directions, the x direction and the y direction, the arrangement pitch of the lattice pieces 231 in the x direction is more than twice the pixel pitch in the x direction, and the arrangement pitch of the lattice pieces 231 in the y direction. Is more than twice the pixel pitch in the y direction. When both the first and second gratings are two-dimensional gratings as shown in FIG. 20, the data of the first and second gratings in the relative movement in the x direction, and the y values of the first and second gratings. Data at the time of relative movement in the direction is obtained. That is, even if the lattice is not rotated around the z axis or the direction of the subject H is not changed, there is no direction that is insensitive as the refraction direction of the X-ray in the xy plane, and the image defect is direction-specific. It is possible to provide a diagnostic image that does not occur.

FIG. 21 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. 21, it is also possible to use first and second absorption type gratings 110 and 111 in which the grating surface is formed in a substantially concave curved surface shape.

The first absorption type grating 110 is configured by arranging a plurality of first grating pieces 110A, and each of the first grating pieces 110A is X-ray transmissive on the surface of the planar substrate 110a. , 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. The pieces 110A are arranged, and the first lattice pieces 110A adjacent to each other in the circumferential direction are arranged, so that the lattice surface is formed in a substantially concave curved surface shape. A boundary portion 110c is formed between the adjacent first lattice pieces 110A.

Similarly, the second absorption type grating 111 is configured by arranging 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 in the circumferential direction are arranged, the lattice surface is formed in a substantially concave curved surface shape. A boundary 111c is formed between the adjacent second lattice pieces 111A.

_{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).

As described above, by configuring the first and second absorption type gratings 110 and 111 by arranging a plurality of grating pieces, it is possible to easily configure the grating surfaces in 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. Further, when the multi slit 103 is used, it is also preferable that the multi slit 103 is formed in a concave curved surface like the first and second absorption gratings 110 and 111.

FIG. 22 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 substantial pitch p ₂ ′ of the arrangement pitch p ₂ of the linear electrodes, the pattern period p ₁ ′ of the self-image G1 at the position of the charge collection electrode 121 (position of the X-ray image detector), and the pixel in the x direction The relationship between the pixel pitch P _{D of} 120 and the pixel pitch P _{D of} Equation (9) is based on the moire period T expressed by Equation (8), similarly to the second absorption grating 32 of the X-ray imaging system 10 described above. It is necessary to satisfy | fill, Furthermore, it is preferable to satisfy | fill 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 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.

When the X-ray image detector configured in this way is applied to, for example, the X-ray imaging system 10 described above, the second absorption type grating 32 from the imaging unit 12 becomes unnecessary. In this configuration, only the first lattice 31 is configured by arranging a plurality of lattice pieces 311. In this configuration, since a fringe image of a plurality of phase components can be acquired by one shooting, 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.

Next, another example of the radiation imaging system for describing the embodiment of the present invention will be described.

The X-ray imaging system shown in FIG. 23 is formed by a first grating 131 that passes through the X-rays emitted from the X-ray source 11 and forms a periodic pattern image (self-image G1), and the first grating 131. A second grating 132 that periodically masks the self-image G1; an X-ray image detector 240 that detects a moire formed by superposition of the self-image G1 and the second grating 132; and an X-ray image detector And a phase contrast image generation unit 260 that acquires a fringe image based on the moire detected by 240 and generates a phase contrast image based on the acquired fringe image. The phase contrast image generation unit 260 constitutes part of the processing of the control device 20 in the console 13 (FIG. 2).

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 The necessary thickness h ₁ in the above-described normal medical diagnostic X-ray energy region is about ₁ μm to 10 μm. Further, an absorption type grating can be used as the first grating 131. On the other hand, the second grating 132 is preferably an absorption grating.

Here, when the X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam, the self-image G1 of the first grating 131 formed through the first grating 131 is , Enlarged in proportion to the distance from the X-ray source 11. In this example, the grating pitch p ₂ of the second grating 132 is substantially the same as the periodic pattern of the bright part of the self-image G 1 of the first grating 131 at the position of the second grating 132. To be decided. That is, when the first grating 131 is a phase modulation type grating or an absorption type grating that applies 90 ° phase modulation, the distance from the focal point of the X-ray source 11 to the first grating 131 is L ₁ , When the distance from the grating 131 to the second grating 132 is L ₂ , the grating pitch p ₂ of the second grating 132 is the above equation (1), and when using the multi slit 103 described above, the multi slit 103 is used. And the distance from the first grating 131 to L ₃ is determined so as to satisfy the relationship of the above formula (21). Further, in the case where the first grating is a phase modulation type grating that applies 180 ° phase modulation, considering that the pitch of the self-image G1 is ½ of the grating pitch p ₁ of the first grating 131, the equation It replaces with (1) and Formula (21), and it determines so that the following Formula (24) and the following Formula (25) may be satisfy | filled.

When the X-rays emitted from the X-ray source 11 are parallel beams, p ₂ = p ₁ when the first grating 131 is a phase modulation type grating or an absorption type grating that applies 90 ° phase modulation. In the case where the first grating is a phase modulation type grating that applies phase modulation of 180 °, it is determined so as to satisfy p ₂ = p _1/2 .

The X-ray image detector 240 detects, as an image signal, an image in which the self-image G1 of the first grating 131 formed by the X-rays incident on the first grating 131 is periodically masked by the second grating 132. To do. 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 schematically shows an external appearance (FIG. 24A), an xz plane cross section (FIG. 24B), and a yz plane cross section (FIG. 24C) of the X-ray image detector 240 of this example.

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. Then, as shown in FIG. 24, the transparent linear electrodes 246a and the light shielding linear electrodes 246b are alternately arranged in parallel at predetermined intervals.

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, as shown in FIG. 24, 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. 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.

In the X-ray imaging system of this example, as shown in FIG. 24, 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. It has. The linear reading light source 250 of this example includes a light source such as an LED (Light Emitting Diode) or LD (Laser Diode) and a predetermined optical system, and is parallel to the extending direction of the transparent linear electrode 246a and the light shielding linear electrode 246b. The X-ray image detector 240 is irradiated with linear reading light having a width of approximately 10 μm in a random direction (y direction). 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 this example, the configuration including the X-ray source 11, the first grating 131, the second grating 132, and the X-ray image detector 240 is caused to function as a Talbot interferometer. For that purpose, some conditions must be almost satisfied. The conditions will be described below. Here, “substantially satisfy” means that the energy of X-rays radiated from the X-ray source 11 under various conditions described later, that is, the wavelength has a width rather than a single one, and therefore has a width relative to the X-ray energy width. This means that there is a permissible width, and that performance such as image quality is inferior because it is not optimal, but in this example, there is a permissible width capable of obtaining at least a phase contrast image.

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. Equation (26) should be nearly satisfied.

Where λ is the wavelength of X-rays (usually the effective wavelength of X-rays incident on the first grating 131), m is 0 or a positive integer, p ₁ is the grating pitch of the first grating 131, and p _2. Is the grating pitch of the second grating 132 described above.

Further, when the first grating 131 is a phase modulation type grating that applies 180 ° phase modulation, the following expression (27) must be substantially satisfied.

Further, when the first grating 131 is an absorption type grating, the following expression (28) must be substantially satisfied. However, m 'is a positive integer.

Further, the thicknesses h ₁ and h ₂ of the first and second gratings 131 and 132 also satisfy the expressions (6) and (7) described above with respect to the first and second absorption gratings 31 and 32. It is preferable to set so.

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 connected to the extending direction of the self-image G1 of the first grating 131 and the second grating 131. The grating 132 is disposed so as to be relatively inclined with respect to the extending direction. 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, the pitch of the second grating 132 and the pattern period of the self-image G1 formed at the position of the second grating 132 by the first grating 131 are p ₁ ′, second When the relative rotation angle in the xy plane of the self-image G1 of the first grating 131 with respect to the first grating 131 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 (29), the phase of the self-image G1 of the first grating 131 and the phase of the second grating 132 for n periods with respect to the length of the image resolution D in the sub-scanning direction. It will shift. FIG. 26 shows a case where M = 5 and n = 1.

Therefore, the image signal obtained by dividing the intensity modulation of n periods of the self-image G1 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 region that passes through the second grating 132 in 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, so that the first The intensity of the self-image G1 of the grating 131 is modulated in the sub-scanning direction.

Since M = 5, an image signal obtained by dividing the intensity modulation of one period of the self-image G1 of the first grating 131 into five by each pixel of Dx × Dy can be detected, that is, each of Dx × Dy. Image signals of five stripe images different from each other can be detected depending on the pixel. 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 G1 of the first grating 131 with respect to the second grating 132, for example, after the relative rotation angle of the X-ray image detector 240 and the second grating 132 is fixed, This can be done by rotating the grating 131.

For example, if p ₁ ′ = 5 μm, D = 50 μm, and n = 1 in the above equation (29), the rotation angle θ is about 5.7 °. The actual rotation angle θ ′ of the self-image G1 of the first grating 131 with respect to the second grating 132 is detected by, for example, the pitch of the moire by the self-image G1 of the first grating and the second grating 132. Can do.

Specifically, as shown in FIG. 27, assuming that the actual rotation angle is θ ′ and the pitch P ′ of the apparent self-image G1 in the x direction generated by the rotation is, the observed moire pitch Pm is 1 Since / Pm = | 1 / P′−1 / p ₁ ′ |, the actual rotation angle θ ′ can be obtained by substituting P ′ = p ₁ ′ / 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 rotation angle θ determined by the above equation (29) is compared with the actual rotation angle θ ′, and the rotation angle of the first grating 131 is adjusted automatically or manually only by the difference. Good.

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, and when a light wave passes through the first grating 131, a self-image G1 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 applies 90 ° phase modulation, the above equation (26) (in the case of a 180 ° phase modulation type grating, the above equation (27), the case of an absorption type grating) Forms a self-image G1 of the first grating 131 at a distance given by the above equation (28). On the other hand, the wavefront of the X-ray incident on the first grating 131 is distorted by the subject H, so that the self-image G1 of the first grating 131 is deformed accordingly.

Subsequently, the X-ray passes through the second grating 132. As a result, the deformed self-image G1 of the first grating 131 forms a moire 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, the operation of image detection and readout in the X-ray image detector 240 will be described.

First, as shown in FIG. 28, 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, the self-image G1 of the first grating 131 and the second grating 132. X-rays with moiré formed by superimposing and 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 of the phase contrast image in the sub-scanning direction is divided into five, and the intensity modulation of one period of the self-image G1 of the first grating 131 is 5 Since the self-image G1 of the first grating 131 is tilted with respect to the second grating 132 so that the divided image signal can be detected, the image read from the first reading line as shown in FIG. The signal is acquired as the first fringe image signal M1, the image signal 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. The image signal acquired as the fringe image signal M3 and 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 the fifth fringe image signal. Acquired as 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 (13) to (19), and thus the description thereof is omitted.

In the above-described configuration in which the first grating 131 and the second grating 132 are inclined, both the first grating 131 and the second grating 132 are configured as absorption gratings, regardless of the presence or absence of the Talbot interference effect. The radiation that has passed through the slit portion may be geometrically projected. 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 can be configured to form the self-image G1 of the first grating behind the first grating 131 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, an image in which the distance d ₁ of the first grating 131 spacing d ₂ of the second grating 132, be about 1 [mu] m ~ 10 [mu] m X-rays passing through the slit portion is formed ignore the effects of diffraction The self-image G1 of the first grating 131 is geometrically projected behind the first grating 131 as much as possible. The grating pitch p ₁ of the first grating 131 for the relationship between the lattice pitch p ₂ of the second grating 132, wherein when the first grating 131 described above is an absorption grating (1), further multi-slit Is used, the distance from the multi-slit 103 to the first grating 131 is L _{3, which} is the same as the equation (21) described above. The inclination of the self-image G1 formed by 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 shown in FIG. 25, the extending direction of the second grating 132 is parallel to the y direction, and the extending direction of the self-image G1 of the first grating 131 is θ relative to the y direction. In contrast, the extending direction of the self-image G1 of the first grating 131 is parallel to the y direction, and the extending direction of the second grating 132 is inclined by θ with respect to the y direction. Good.

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 sensor 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. In addition, the self-image G 1 of the first grating 131 is tilted with respect to the second grating 132. The specific rotation angle of the self-image G1 of the first grating 131 is calculated by the above equation (29) as in the above example.

In the above equation (29), for example, when the rotation angle θ of the self-image G1 of the first grating 131 is set with M = 5 and n = 1, the one grating circuit 270 in FIG. An image signal obtained by dividing the intensity modulation of one period of the self-image G1 into five can be detected, that is, five stripes different from each other depending on five pixel circuit rows connected to the five gate electrodes 273 shown in FIG. Each image signal of the image can be detected. 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 a CMOS sensor, 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 illustrated in FIG. As shown in FIG. 33, a plurality of two-dimensional arrays can be used. The X-ray image detector using the CMOS sensor is provided for each pixel circuit row, and includes a large number of gate electrodes 282 that output a drive signal for driving a signal readout circuit included in the pixel circuit 280 and a reset. An electrode 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 are provided. 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 the X-ray image detector using the CMOS sensor as described above is installed so that the second grating 132 and the pixel circuit array (data electrode) are parallel as shown in FIG. One 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, for example, 10 μm in the case of an X-ray image detector using a CMOS sensor.

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. In addition, the self-image G 1 of the first grating 131 is tilted with respect to the second grating 132. A specific self-image G1 rotation angle of the first grating 131 is calculated by the above equation (29) as in the above example.

In the above equation (29), for example, when M = 5 and n = 1 and the rotation angle θ of the self-image G1 of the first grating 131 is set, one pixel circuit 280 in FIG. An image signal obtained by dividing the intensity modulation of one period of the self-image G1 into five can be detected, that is, five stripes different from each other depending on five pixel circuit rows connected to the five gate electrodes 282 shown in FIG. Each image signal of the image can be detected. 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.

Note that the extending direction of the gate electrode and the data electrode of the X-ray image detector is not limited to the example shown in FIGS. 32 and 35. For example, the gate electrode is in the vertical direction on the paper surface and the data line is in the horizontal direction on the paper surface. An X-ray image detector may be arranged.

Further, the self-image G1 of the first grating 131 and the second grating 132 may be rotated by 90 ° with respect to the arrangement of the X-ray image detectors as shown in FIGS. In this case, by acquiring the image signals read from the pixel circuits 270 and 280 arranged in the direction parallel to the gate electrode, the image signals constituting the different fringe images are acquired as in the above example. be able to.

In addition, the periodic direction of the self-image G1 of the first grating 131 or the periodic direction of the second grating 132, and one of the directions in which the pixel circuits 270 and 280 of the X-ray image detector are arranged Do not necessarily match. Configuration capable of acquiring image signals of pixels arranged in a direction parallel to the periodic direction of moire generated by the self-image G1 of the first grating 131 and the second grating 132, or in an intersecting direction other than the orthogonal direction If so, the relationship between the self-image G1 of the first grating 131 and the periodic direction of the second grating 132 and the arrangement direction of the pixel circuit 270 of the X-ray image detector may be any relationship.

As described above, an X-ray image detector using a TFT switch or an X-ray image detector using a CMOS sensor can be used. However, these X-ray image detectors generally have pixels. Due to the square shape, 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.

Next, a configuration example of another X-ray imaging system for explaining an example of the present invention will be described. FIG. 36 shows a schematic configuration of the X-ray imaging system of this example.
As shown in FIG. 36, the X-ray imaging system has a grating 131 that passes X-rays emitted from the X-ray source 11 to form a periodic pattern image (self-image G1), and a self-image formed by the grating 131. An X-ray image detector 340 that selectively detects G1; and a moving mechanism 333 that moves the X-ray image detector 340 relative to the grating 131 in a direction orthogonal to the extending direction of the linear electrodes; And a phase contrast image generation unit 260 that generates a phase contrast image based on the intensity modulation signal obtained from the output of the X-ray image detector 340 in accordance with the relative movement by the movement mechanism 333.

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 the self-image G1 of the grating 131 formed by the grating 131 by passing the X-rays through the grating 131, and the charge signal corresponding to the self-image G1 in a lattice shape to be described later. By accumulating in the divided charge accumulation layer, the self-image G1 is selectively detected in a region and 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 an appearance FIG. Of the X-ray image detector 340 of this example. 37A), an xz plane cross section (FIG. 37B), and a yz plane cross section (FIG. 37C) are schematically shown.

As shown in FIG. 37, the X-ray image detector 340 in this example generates charges by receiving the first electrode layer 241 that transmits X-rays and the irradiation of X-rays that have transmitted through the first electrode layer 241. Among the charges generated in the recording photoconductive layer 242 and the recording photoconductive layer 242, it acts as an insulator for charges of one polarity, and acts as a conductor for charges of the other polarity. The charge storage layer 343, a reading photoconductive layer 245 that generates charges when irradiated with reading light, 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 (SexTe1-x ) ₃ (0.5 <x <1), As ₂ Se ₃ with Se replaced to about 50%, AsxSey with As concentration changed by about ± 15% from As ₂ Se ₃ (x + y = 100, 34.ltoreq.x.ltoreq.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.

Then, as shown in FIG. 37, the charge storage layer 343 in this example is linearly divided 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. ing.

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. That is, the charge storage layer 343 divided linearly as described above has a function of periodically masking the self-image G1 of the grating 131 in the same manner as the second grating 132 in each example described above. . The arrangement pitch p ₂ and the interval d ₂ of the transparent linear electrodes 246a or the light shielding linear electrodes 246b are determined in the same manner as the lattice pitch p ₂ and the interval d ₂ related to the second grating 132 described above, and therefore the same reference numerals are used. It explains using.

Specifically, when the X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam, the self-image G1 of the grating 131 formed through the grating 131 is expressed by the X-ray source. It is enlarged in proportion to the distance from 11. 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 the periodic pattern of the light area of the self image G1 of the grating 131 at the position of the charge accumulation layer 343 It is determined to match. That is, the grating pitch of the grating 131 is p ₁ , the pattern period p ₁ ′ of the self-image G1 at the position of the detection surface of the X-ray image detector 340, the distance from the focal point of the X-ray source 11 to the grating 131 is L ₁ , and the grating When the distance from 131 to the detection surface of the X-ray image detector 340 is L ₂ , the arrangement pitch p ₂ of the charge storage layer 343 is expressed by the above equation (1), and when using the multi slit 103 described above, the multi pitch the distance from the slit 103 to grating 131 as _{L 3,} is determined to satisfy the above equation (21).

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, as shown in FIG. 37, an image signal of one pixel is read out by one set of transparent linear electrode 246a and light shielding linear electrode 246b. 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.

In the X-ray imaging system of this example, as shown in FIG. 37, a linear reading light source 250 extended in a direction (x direction) orthogonal to the extending direction of the transparent linear electrode 246a and the light shielding linear electrode 246b. It has.

In this example, a 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 is caused to function as a Talbot interferometer. In order to do this, several conditions 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. (26) must be almost satisfied.

When the grating 131 is a phase modulation type grating that applies 180 ° phase modulation, the above expression (27) must be substantially satisfied, and when the grating 131 is an absorption type grating, the above expression (28) must be substantially satisfied. I must.

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 on the grating 131 are diffracted by the grating 131 to form a Talbot interference image (self-image G1 of the grating 131) at a predetermined distance from the grating 131 in the optical axis direction of the X-ray.

Then, the self-image G1 of the grating 131 is incident from the first electrode layer 241 side of the X-ray image detector 340 and is region-selectively stored by the divided charge storage layer 343 of the X-ray image detector 340. An X-ray image detector 340 outputs the image signal.

Here, the operation of detecting and reading out the fringe image in the X-ray image detector 340 will be described in more detail.

First, as shown in FIG. 38, in the 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 G1 of the grating 131 becomes X-ray. Irradiation is performed from the first electrode layer 241 side of the image detector 340 (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. It is stored in the charge storage layer 343 as an image charge (see FIG. 38B).

Here, since the charge storage layer 343 in this example is linearly divided at the arrangement pitch as described above, out of the charges generated according to the self-image G1 of the grating 131 in the recording photoconductive layer 242, Only charges in which the charge storage layer 343 exists immediately below 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). Then, 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 having the linear electric charge accumulation layer 343 immediately below the electric charge, the self-image of the lattice 131 becomes the line of the electric charge accumulation layer 343. By overlapping with the pattern, the region is selectively stored in the charge storage 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 (13) to (19), 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, as the distance Z ₂ to the detection surface of the X-ray image detector 340 from the grating 131 becomes Talbot interference distance, depending on the type of grating 131, the above equation (26) or formula Although either (27) or (28) is satisfied, the grating 131 may be configured to project incident X-rays without diffracting. According to this configuration, since the projected image projected through the grating 131 is obtained similarly at a position behind the grating 131, the distance Z ₂ from the grating 131 to the detection surface of the X-ray image detector 340 is obtained. Can be set regardless 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 described above, also in this example, as shown in FIGS. 25 to 27 and the like, the grating 131 and the X-ray image detector 340 are connected to each other. The extending direction of the self-image G1 and the extending direction of the charge storage layer 343 of the X-ray image detector 340 are arranged so as to be relatively inclined. 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. In the same manner as 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 accumulation layer 343 is provided directly on the second electrode layer 246 without the reading photoconductive layer 245, and the linear charge accumulation layer 343 can be formed by vapor deposition. The charge storage layer 343 can be easily formed. In the vapor deposition process, a metal mask or the like is used to selectively form a linear pattern. In the configuration in which the linear charge storage layer 343 is provided on the reading photoconductive layer 245, a step of setting a metal mask for forming the linear charge storage layer 343 by vapor deposition after vapor deposition of the reading photoconductive layer 245 is performed. Therefore, the reading photoconductive layer 245 is deteriorated or foreign matter is mixed in between the photoconductive layers by an operation in the air between the reading photoconductive layer 245 vapor deposition step and the recording photoconductive layer 242 vapor deposition step. May cause deterioration of quality. On the other hand, by adopting a structure in which the above-described reading photoconductive layer 245 is not provided, operations in the air after the photoconductive layer is deposited can be reduced, so that the above-described concern about 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, as shown in FIG. 41, in a 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 source 400, the X-ray carrying the self-image G1 of the grating 131 is Irradiation is performed from the first electrode layer 241 side of the image detector 360 (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 G1 of the lattice 131 is stored in the charge storage layer 343 in a region-selective manner by overlapping with the linear pattern of the charge storage layer 343.

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.

FIG. 44 shows a configuration of a calculation unit regarding another example of the radiation imaging system for explaining the embodiment of the present invention.

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

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

The absorption image generation unit 192 generates an absorption image by averaging the pixel data I _k (x, y) obtained for each pixel with respect to k and calculating an average value as shown in FIG. 45. 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 detection intensity unevenness of the detection system (including information such as grating transmittance unevenness, in-plane unevenness of irradiation X-ray intensity, and sensitivity unevenness of the X-ray image detector). Therefore, a correction coefficient map for correcting detection intensity 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 acquired by shooting in the presence of the subject (main shooting), and the detection coefficient unevenness of the detection system is corrected by applying the above correction coefficient to each pixel. 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.

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.

In each example described above, the present invention is applied to an apparatus for medical diagnosis. However, the present invention is not limited to medical diagnosis use, and can be applied to other radiation detection apparatuses for industrial use. .

As described above, in this specification, the first grating having a plurality of strips arranged in the first direction and the radiation image formed by the radiation that has passed through the first grating have a period. And a radiological image detector that detects the radiographic image masked by the grid pattern using a plurality of pixels, wherein the first grid includes the first grid A plurality of first grating pieces arranged in the first direction in a plane intersecting an optical axis of radiation passing through one grating, and the arrangement of the first grating pieces in the first direction A radiation image detection apparatus is disclosed, wherein the pitch is at least twice the pixel pitch in the first direction of the radiation image detector.

Further, in the present specification, a first grating having a plurality of strips arranged in a first direction and a radiation image formed by radiation passing through the first grating are periodically masked. A second grating having a plurality of strips arranged in a line, and a radiation image detector for detecting the radiation image masked by the second grating using a plurality of pixels, and The grating includes a plurality of second grating pieces arranged at least in the first direction in a plane intersecting an optical axis of radiation passing through the first grating and the second grating, An arrangement pitch of the lattice pieces in the first direction is at least twice as large as a pixel pitch in the first direction of the radiological image detector.

Further, in the present specification, a first grating having a plurality of strips arranged in a first direction and a radiation image formed by radiation passing through the first grating are periodically masked. A second grating having a plurality of strips arranged in a line, and a radiation image detector for detecting the radiation image masked by the second grating using a plurality of pixels, The grating includes a plurality of first grating pieces arranged in at least the first direction in a plane intersecting the optical axis of the radiation passing through the first grating, and the first grating piece includes the first grating pieces. The arrangement pitch in the direction 1 is at least twice the pixel pitch in the first direction of the radiation image detector, and the second grating is at least in the plane intersecting the optical axis of the radiation. A plurality of second array arranged in one direction The radiation image is characterized in that the arrangement pitch of the second grating pieces in the first direction is not less than twice the pixel pitch in the first direction of the radiation image detector. A detection device is disclosed.

Further, in the radiological image detection apparatus disclosed in this specification, the arrangement pitch of the first grating pieces and the arrangement pitch of the second grating pieces are substantially equal.

In the radiological image detection apparatus disclosed in the present specification, the radiation passing through the first grating is a cone beam whose irradiation range is expanded in proportion to the distance from the radiation focus, and the arrangement pitch is The value corrected based on the enlargement ratio, which is the ratio between the distance between the grating having the arrangement pitch and the radiation focus and the distance between the radiation image detector and the radiation focus.

Further, in the radiological image detection apparatus disclosed in the present specification, the dimension of the grating in the direction along the optical axis of the radiation is h, the distance between the stripes of the grating is d, and the grating Where P _{B is} the arrangement pitch of R and R is the distance between the grating and the radiation focus.

Meet.

In the radiological image detection apparatus disclosed in this specification, the first direction is along the circumferential direction of a cylindrical surface having an axis passing through the radiation focus.

In the radiological image detection apparatus disclosed in this specification, the arrangement pitch is based on the fundamental frequency of moire detected by the radiographic image detector and the frequency of harmonic components up to at least the fourth order. Can be decided.

Further, in the radiological image detection apparatus disclosed in this specification, the arrangement pitch is 6 times or more the pixel pitch.

In the radiological image detection apparatus disclosed in this specification, the arrangement pitch is 8 times or more the pixel pitch.

Further, in the radiological image detection apparatus disclosed in this specification, the arrangement pitch is 10 times or more the pixel pitch.

Further, in the radiological image detection apparatus disclosed in the present specification, in the lattice including the plurality of lattice pieces, a band-shaped boundary portion is formed between adjacent lattice pieces.

Further, in the radiological image detection apparatus disclosed in the present specification, in the lattice including the plurality of lattice pieces, the plurality of lattice pieces are also arranged in a second direction intersecting the first direction. The

Moreover, in the radiological image detection apparatus disclosed in this specification, the first grating further includes a plurality of strips arranged in a second direction intersecting the first direction.

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

Further, in the present specification, the distribution of the refraction angle of the radiation incident on the radiation image detector is calculated from the above-described radiation imaging apparatus and the image detected by the radiation image detector of the radiation imaging apparatus, A radiation imaging system including an arithmetic processing unit that generates a phase contrast image of a subject based on the distribution of refraction angles is disclosed.

According to the present invention, it is possible to stably manufacture a grating having a size corresponding to a required visual field size, and it is possible to stably produce a phase contrast without degrading the image quality of the phase contrast image due to the arrangement of the grating pieces. An image is obtained.

Although the present invention has been described in detail and with reference to specific embodiments, it will be apparent to those skilled in the art that various changes and modifications can be made without departing from the spirit and scope of the invention.
This application is based on a Japanese patent application filed on November 26, 2010 (Japanese Patent Application No. 2010-264242), the contents of which are incorporated herein by reference.

10 X-ray imaging system 11 X-ray source (radiation source)
12 Imaging unit (radiation image detection device)
13 Console (control calculation means)
30 Flat panel detector (FPD)
31 First absorption type grating 31a Substrate 31b X-ray shielding part (strip)
32 Second absorption type grating 32a Substrate 32b X-ray shielding part (strip)
311 1st lattice piece 312 2nd lattice piece 315 Boundary portion 325 Boundary portion P _B1 , P _B2 arrangement pitch P _C1 , P _C2 arrangement pitch P _D image pitch

Claims (16)

1. A first grid having a plurality of strips arranged in a first direction;
   A grating pattern arranged to periodically mask a radiation image formed by radiation passing through the first grating;
   A radiation image detector that detects the radiation image masked by the lattice pattern using a plurality of pixels;
   The first grating includes a plurality of first grating pieces arranged at least in the first direction in a plane intersecting an optical axis of radiation passing through the first grating, and the first grating An arrangement pitch of the pieces in the first direction is at least twice as large as a pixel pitch in the first direction of the radiological image detector.
2. A first grid having a plurality of strips arranged in a first direction;
   A second grating having a plurality of strips arranged to periodically mask a radiation image formed by radiation that has passed through the first grating;
   A radiation image detector that detects the radiation image masked by the second grating using a plurality of pixels;
   The second grating includes a plurality of second grating pieces arranged at least in the first direction in a plane intersecting the optical axis of the radiation passing through the first grating and the second grating. The arrangement pitch of the second lattice pieces in the first direction is at least twice the pixel pitch of the radiographic image detector in the first direction.
3. A first grid having a plurality of strips arranged in a first direction;
   A second grating having a plurality of strips arranged to periodically mask a radiation image formed by radiation that has passed through the first grating;
   A radiation image detector that detects the radiation image masked by the second grating using a plurality of pixels;
   The first grating includes a plurality of first grating pieces arranged at least in the first direction in a plane intersecting an optical axis of radiation passing through the first grating, and the first grating The arrangement pitch of the pieces in the first direction is not less than twice the pixel pitch in the first direction of the radiation image detector,
   The second grating includes a plurality of second grating pieces arranged in at least the first direction in a plane intersecting the optical axis of the radiation,
   An arrangement pitch of the second lattice pieces in the first direction is at least twice as large as a pixel pitch in the first direction of the radiographic image detector.
4. The radiological image detection apparatus according to claim 3,
   The radiological image detection apparatus according to claim 1, wherein the arrangement pitch of the first grating pieces and the arrangement pitch of the second grating pieces are substantially equal.
5. The radiological image detection apparatus according to any one of claims 1 to 4,
   The radiation passing through the first grating is a cone beam whose irradiation range is expanded in proportion to the distance from the radiation focus,
   The arrangement pitch is a value corrected based on an enlargement ratio that is a ratio of a distance between a grating having the arrangement pitch and a radiation focus and a distance between the radiation image detector and the radiation focus. A radiological image detection apparatus characterized by that.
6. The radiological image detection apparatus according to claim 5,
   The dimension of the grating in the direction along the optical axis of the radiation is h, the distance between the stripes of the grating is d, the arrangement pitch of the grating is P _B , and between the grating and the radiation focus If the distance is R,
   

   

   
   The radiation image detection apparatus characterized by satisfy | filling.
7. The radiological image detection apparatus according to claim 5,
   The radiation image detecting apparatus according to claim 1, wherein the first direction is along a circumferential direction of a cylindrical surface having an axis passing through the radiation focus.
8. The radiological image detection apparatus according to any one of claims 1 to 7,
   The arrangement pitch is determined based on a fundamental frequency of moire detected by the radiological image detector and frequencies of harmonic components up to at least the fourth order.
9. The radiological image detection apparatus according to any one of claims 1 to 8,
   The radiological image detection apparatus, wherein the arrangement pitch is 6 times or more the pixel pitch.
10. The radiological image detection apparatus according to claim 9,
    The radiological image detection apparatus, wherein the array pitch is 8 times or more the pixel pitch.
11. It is a radiographic image detection apparatus of Claim 10, Comprising:
    The radiological image detection apparatus, wherein the arrangement pitch is 10 times or more of the pixel pitch.
12. It is a radiographic image detection apparatus as described in any one of Claim 1 to 11, Comprising:
    In the lattice including the plurality of lattice pieces, a boundary portion is formed between adjacent lattice pieces.
13. The radiological image detection apparatus according to any one of claims 1 to 12,
    In the lattice including the plurality of lattice pieces, the plurality of lattice pieces are also arranged in a second direction intersecting with the first direction.
14. The radiological image detection apparatus according to any one of claims 1 to 13,
    The first grating further includes a plurality of strips arranged in a second direction intersecting the first direction.
15. The radiological image detection apparatus according to any one of claims 1 to 14,
    A radiation imaging apparatus comprising: a radiation source that irradiates radiation toward the first grating.
16. A radiographic apparatus according to claim 15;
    From the image detected by the radiation image detector of the radiation imaging apparatus, the distribution of the refraction angle of the radiation incident on the radiation image detector is calculated, and based on the refraction angle distribution, the phase contrast image of the subject is calculated. A radiation imaging system comprising: an arithmetic processing unit to generate.

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