JP2012110395A - Radiographic system - Google Patents

Abstract

PROBLEM TO BE SOLVED: To prevent the deterioration of the image quality of an obtained radiation phase contrast image caused by the vibration of a radiation source generated by air-cooling of the radiation source such as an X-ray source in phase imaging by radiation such as X-rays etc.SOLUTION: The radiographic system includes: an X-ray source 11; a first absorption lattice 31 set in a progress direction of an X-ray emitted from the X-ray source 11; a second absorption lattice 32 having a period substantially agreeing with a pattern period of an X-ray image formed by the X-ray having passed through the first absorption lattice 31; an FPD (Flat Panel Detector) 33 to detect an X-ray image masked by the second absorption lattice 32; an X-ray tube cooler 27 to air-cool the X-ray source 11; and a controller 20 to stop air-cooling by the X-ray tube cooler 27 when imaging an object, in a first imaging mode wherein the object is disposed, on the upstream side of the first absorption lattice 31 in a progress direction of an X-ray or between the first absorption lattice 31 and the second absorption lattice 32, and is imaged.

Description

  The present invention relates to a radiation imaging system.

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

  In a general X-ray imaging system, a subject is placed between an X-ray source that emits X-rays and an X-ray image detector that detects 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, in addition to a combination of an X-ray intensifying screen and a film and a stimulable phosphor, a flat panel detector (FPD) using a semiconductor circuit is widely used. As an X-ray source used in such an X-ray imaging system, an X-ray source including a so-called anode rotating type X-ray tube is widely used.

  However, since the X-ray absorption ability is lower as a substance composed of an element having a smaller atomic number, a sufficient softness or contrast (contrast) of an X-ray absorption image cannot be obtained in a soft body tissue or soft material. 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 this background, in recent years, instead of X-ray intensity changes by the subject, an image based on the X-ray phase change (angle change) by the subject (hereinafter referred to as a phase contrast image) is obtained. Research on line phase imaging has been actively conducted. In general, it is known that when X-rays are incident on an object, the interaction is higher in phase than in X-ray intensity. For this reason, in the X-ray phase imaging using the phase difference, a high-contrast image can be obtained even for a weakly absorbing object having a low X-ray absorption capability. As a kind of such X-ray phase imaging, in recent years, an X-ray imaging system using an X-ray Talbot interferometer comprising two transmission diffraction gratings (phase grating and absorption grating) and an X-ray image detector has been proposed. It has been devised (for example, see Patent Document 1).

  In the X-ray Talbot interferometer, a first diffraction grating (phase type grating or absorption type grating) is arranged behind the subject, and a specific distance (Talbot interference distance) determined by the grating pitch of the first diffraction grating and the X-ray wavelength. ) Is disposed downstream, and an X-ray image detector is disposed behind the second diffraction grating (absorption type 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 X-ray Talbot interferometer, the moiré fringes generated by superposition (intensity modulation) of the self-image of the first diffraction grating and the second diffraction grating are detected, and the change in the moire fringes caused by the subject is analyzed. Obtain sample phase information. As a method for analyzing moire fringes, for example, a fringe scanning method is known. According to this fringe scanning method, the second diffraction grating is substantially parallel to the surface of the first diffraction grating with respect to the first diffraction grating and substantially in the grating direction (strip direction) of the first diffraction grating. The angle of X-rays refracted by the subject from a change in the signal value of each pixel obtained by the X-ray image detector, which is taken multiple times while being translated in the vertical direction at a scanning pitch obtained by equally dividing the lattice pitch. A distribution (differential image of phase shift) is obtained, and a phase contrast image of the subject can be obtained based on this angular distribution.

JP 2008-200320 A

  However, the refraction angle of X-rays generated by passing through the subject is only a few μrad, and it is obtained by detecting the intensity modulation signal of the moire fringe generated by this refraction angle and the modulated moire fringe by the fringe scanning method. The change in signal is also slight. When such a slight change is measured, the relative position shift between the X-ray focal point and the first and second diffraction gratings affects the detection accuracy of the phase information of the subject.

  The relative position shift between the X-ray focal point and the first and second diffraction gratings can occur due to vibrations applied to the first and second diffraction gratings. Japanese Patent Application Laid-Open No. 2004-228688 discloses that vibration from a subject table is absorbed by a buffer material and transmission of vibration to the first diffraction grating and the second diffraction grating is blocked. However, the relative position shift between the X-ray focal point and the first and second diffraction gratings can also be caused by the vibration of the X-ray source. Patent Document 2 does not consider the vibration of the X-ray source and the shift of the relative position between the first diffraction grating and the second diffraction grating caused by the vibration.

  This is because a cooling means having a fan for cooling the rotary anode type X-ray source is generally provided in the X-ray imaging system. That is, the rotation of the fan of the cooling means is transmitted to the X-ray source, or the X-ray source is vibrated by the wind from the fan.

When the X-ray focal point and the first diffraction grating are relatively displaced in the translational movement direction (x direction) of the second diffraction grating due to the vibration of the X-ray source, the first diffraction grating is positioned at the position of the second diffraction grating. The self-image by the diffraction grating is blurred in the translational movement direction, the contrast of the intensity modulation signal change detected by scanning the second diffraction grating is lowered, and the phase detection accuracy is lowered.
Note that the appropriate grating pitch and grating interval of the first diffraction grating and the second diffraction grating are related to the distance from the focal point (z direction) and are determined by equations (1) and (2) described later.

  In addition, when the X-ray focal point, the first diffraction grating, and the second diffraction grating are displaced relative to each other in the z direction due to the vibration of the X-ray source, the enlargement ratio changes, so that the second grating The pitch of the second grating is shifted relative to the pitch of the incident radiation image, and a moire with a spatial frequency corresponding to the shift occurs. Although it is possible to design correction processing to such an extent that moire on the image does not cause image problems by using a separately acquired image, such as immediately before or after shooting, or by applying a suitable filter process. It is very difficult to correct the moire in which the spatial frequency varies according to the relative shift amount between the X-ray focal point and each grating, and as a result, the image quality of the phase contrast image is deteriorated.

  The present invention has been made in view of the above-described circumstances, and is obtained due to vibration of a radiation source caused by air-cooling a radiation source such as an X-ray source in phase imaging using radiation such as X-rays. The object is to prevent the quality of the radiation phase contrast image from deteriorating.

  A radiation source, a first grating arranged in the traveling direction of the radiation emitted from the radiation source, and a period substantially matching a pattern period of a radiation image formed by the radiation that has passed through the first grating A radiation pattern detector that detects the radiation image masked by the grating pattern, a cooling unit that air-cools the radiation source, and upstream of the first grating in the radiation traveling direction, or the A radiation imaging system comprising: a control unit configured to stop air cooling by the cooling unit during imaging of a subject in a first imaging mode in which imaging is performed by placing an object between a first grid and the grid pattern.

  According to the present invention, in the X-ray phase imaging using radiation, the focus of the radiation source, the first grating, and the grating pattern are stopped by stopping the air cooling by the cooling unit that causes the radiation source to vibrate when photographing the subject. Can be reduced or prevented, and the quality of the obtained radiation phase contrast image can be improved.

It is a schematic diagram which shows the structure of an example of the radiography system for describing embodiment of this invention. It is a control block diagram of the radiography system of FIG. It is a schematic diagram which shows the structure of the radiographic image detector of the radiography system of FIG. It is a perspective view of the imaging part of the radiography system of FIG. It is a side view of the imaging part of the radiography system of FIG. It is a schematic diagram which shows the mechanism for changing the period of the moire fringe by superimposition of the 1st and 2nd grating | lattice. It is a schematic diagram for demonstrating the refraction | bending of the radiation by a to-be-photographed object. It is a schematic diagram for demonstrating the fringe scanning method. It is a graph which shows the signal of the pixel of the radiographic image detector accompanying a fringe scanning. It is a graph for demonstrating the prediction method of the temperature of a radiation source. It is a flow which shows operation | movement 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 perspective view of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the radiographic image detector regarding the other example of the radiography system for describing embodiment of this invention. It is the control block diagram regarding the other example of the radiography system for describing embodiment of this invention. FIG. 20 is a flowchart showing an operation of the radiation imaging system of FIG. 19. FIG. It is a block diagram which shows the structure of the calculating part which produces | generates a radiographic image regarding the other example of the radiography system for describing embodiment of this invention. It is a graph which shows the signal of the pixel of the radiographic image detector for demonstrating the process in the calculating part of the radiography system of FIG. It is a schematic 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 a 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 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 for demonstrating the effect | action which acquires a some fringe image based on the image signal read from the radiographic image detector of an optical reading system. It is a figure which shows the arrangement | positioning relationship between the radiographic image detector using a TFT switch, and the 1st and 2nd grating | lattice. It is a figure which shows schematic structure of the radiographic image detector using CMOS. It is a figure which shows the structure of one pixel circuit of the radiographic image detector using CMOS. It is a figure which shows the arrangement | positioning relationship between the radiographic image detector using CMOS, and the 1st and 2nd grating | lattice. It is a schematic block diagram of the other example of the radiography system for describing embodiment of this invention. It is a figure which shows schematic structure of one Embodiment of a radiographic image detector. It is a figure for demonstrating the effect | action of recording of one Embodiment of a radiographic image detector. It is a figure for demonstrating the effect | action of reading of one Embodiment of a radiographic image detector. It is a figure which shows other embodiment of a radiographic image detector. It is a figure for demonstrating the effect | action of recording of other embodiment of a radiographic image detector. It is a figure for demonstrating the effect | action of the reading of other embodiment of a radiographic image detector. It is a figure which shows an example of the grating | lattice which made the grating | lattice surface into the concave surface.

  FIG. 1 shows a configuration of an example of a radiation imaging system for explaining an embodiment of the present invention, and FIG. 2 shows a control block of the radiation imaging system of FIG.

  The X-ray imaging system 10 is an X-ray diagnostic apparatus that images a subject (patient) H in an upright position, and is disposed opposite to the X-ray source 11 and an X-ray source 11 that irradiates the subject H with X-rays. The imaging unit 12 detects X-rays transmitted through the subject H from the X-ray source 11 and generates image data, and the exposure operation of the X-ray source 11 and the imaging unit 12 based on the operation of the operator. The console is roughly divided into a console 13 that controls the operation and generates a phase contrast image by performing arithmetic processing on the image data acquired by the photographing unit 12.

  The X-ray source 11 is held movably in the vertical direction (x direction) by an X-ray source holding device 14 suspended from the ceiling. The photographing unit 12 is held by a standing stand 15 installed on the floor so as to be movable in the vertical direction.

  The X-ray source 11 emits from the housing 18 having an X-ray tube that generates X-rays according to the high voltage applied from the high voltage generator 16 based on the control of the X-ray source control unit 17, and the housing 18. And a collimator unit 19 including a movable collimator 19a that limits the irradiation field so as to shield a portion of the X-rays that do not contribute to the examination region of the subject H. The X-ray tube of the housing 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. X-rays are generated. The colliding portion of the rotating anode 18a with the electron beam becomes the X-ray focal point 18b.

  The X-ray source 11 includes an X-ray tube cooler 27 having a fan, an X-ray tube thermometer 28, and an X-ray tube vibrometer 29, each operating based on the control of the X-ray source control unit 17. Also have.

  The X-ray tube cooler 27 drives or stops the fan based on the control of the X-ray source control unit 17. The X-ray tube cooler 27 transmits a drive signal to the X-ray source control unit 17 when the fan is driven, and transmits a stop signal to the X-ray source control unit 17 when the fan is stopped.

  The X-ray tube thermometer 28 measures the temperature of the X-ray tube based on the control of the X-ray source control unit 17, and transmits the measured temperature value to the X-ray source control unit 17.

  The X-ray tube vibrometer 29 measures the vibration generated in the X-ray tube based on the control of the X-ray source control unit 17 and transmits the measured vibration value to the X-ray source control unit 17.

  Further, the X-ray tube cooler 27 may not be provided in the X-ray source 11 as long as the X-ray tube can be air-cooled. For example, the X-ray tube cooler 27 may be provided in the X-ray source holding device 14.

  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.

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

  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 that controls the entire X-ray imaging system 10 including a CPU, a ROM, a 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 An image 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. Yes.

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

  The imaging unit 12 includes a flat panel detector (FPD) 30 made of a semiconductor circuit, a first absorption grating 31 for detecting phase change (angle 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 FPD 30 is disposed so that the detection surface is orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11. Although described in detail later, the first and second absorption gratings 31 and 32 are disposed between the FPD 30 and the X-ray source 11. The imaging unit 12 also has a scanning mechanism 33 that changes the relative position of the first absorption type grating 31 relative to the second absorption type grating 32 by translating the second absorption type grating 32 in the vertical direction. Is provided. The scanning mechanism 33 is configured by an actuator such as a piezoelectric element, for example.

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

  The FPD 30 serving as a radiation image detector includes an image receiving unit 41 in which a plurality of pixels 40 that convert X-rays into electric charges and accumulate them in a two-dimensional array on an active matrix substrate, and electric charges from the image receiving unit 41. A scanning circuit 42 that controls the readout timing of the data, a readout circuit 43 that reads out the charges accumulated in each pixel 40, converts the charges into image data and stores them, and calculates the image data via the I / F 25 of the console 13. The data transmission circuit 44 is configured to transmit to the processing 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.

  The pixel 40 directly converts X-rays into electric charges by a conversion layer (not shown) such as amorphous selenium, and directly stores the converted electric charges in a capacitor (not shown) connected to an electrode below the conversion layer. It can be configured as a conversion type X-ray detection 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.

Note that the pixel 40 temporarily converts X-rays into visible light using a scintillator (not shown) made of gadolinium oxide (Gd ₂ O ₃ ), cesium iodide (CsI), or the like, and converts the converted visible light into a photodiode ( It is also possible to configure as an indirect conversion type X-ray detection element that converts into electric charge and accumulates it (not shown). In this embodiment, an FPD based on a TFT panel is used as a radiation image detector. However, the present invention is not limited to this, and various types of radiation image detectors based on a solid-state imaging device such as a CCD sensor or a CMOS sensor. It is also possible to use.

  The readout circuit 43 includes an integration amplifier circuit, an A / D converter, a correction circuit, and an image memory (all not shown). The integrating amplifier circuit integrates the charges output from each pixel 40 via the signal line 46, converts them into a voltage signal (image signal), and inputs it to the A / D converter. The A / D converter converts the input image signal into digital image data and inputs the digital image data to the correction circuit. The correction circuit performs offset correction, gain correction, and linearity correction on the image data, and stores the corrected image data in the image memory. As correction processing by the correction circuit, correction of X-ray exposure amount and exposure distribution (so-called shading) and pattern noise depending on FPD 30 control conditions (drive frequency and readout period) (for example, leak signal of TFT switch) May be included.

  4 and 5 show an imaging unit of the radiation imaging system of FIG.

  The first absorption type grating 31 is composed of a substrate 31a and a plurality of X-ray shielding portions 31b arranged on the substrate 31a. Similarly, the second absorption type grating 32 includes a substrate 32a and a plurality of X-ray shielding portions 32b arranged on the substrate 32a. The substrates 31a and 31b are both made of an X-ray transparent member such as glass that transmits X-rays.

  Each of the X-ray shielding portions 31b and 32b is in one direction in a plane orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11 (in this embodiment, the y direction orthogonal to the x direction and the z direction). It is the linear member extended | stretched. As a material of each X-ray shielding part 31b, 32b, a material excellent in X-ray absorption is preferable, and for example, a 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 portion 31b has a predetermined interval d ₁ with a constant period p ₁ in a direction orthogonal to the one direction (in the x direction in the present embodiment) in a plane orthogonal to the optical axis A of the X-ray. Are arranged with a space between them. Similarly, X-ray shielding portion 32b, in the plane orthogonal to the optical axis A of the X-ray (in the present embodiments, x direction) direction perpendicular to the direction of constant at a period p _2, a predetermined one another They are arranged at intervals d _2. Since the first and second absorption gratings 31 and 32 do not give a phase difference to incident X-rays but give an intensity difference, they are also called amplitude gratings. The slit portions (regions with 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 light metal.

The first and second absorption gratings 31 and 32 are configured to geometrically project the X-rays that have passed through the slit portion regardless of the presence or absence of the Talbot interference effect. Specifically, by setting the distances d ₁ and d ₂ to a value sufficiently larger than the peak wavelength of X-rays emitted from the X-ray source 11, most of the X-rays included in the irradiated X-rays are slit at the slit portion. It is configured to pass through without being diffracted while maintaining straightness. For example, when tungsten is used as the rotary anode 18a described above and the tube voltage is 50 kV, the peak wavelength of the X-ray is about 0.4 mm. In this case, if the distances d ₁ and d ₂ are about ₁ to 10 μm, most of the X-rays are geometrically projected without being diffracted by the slit portion.

The X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam having the X-ray focal point 18b as a light emitting point, and therefore a projected image projected through the first absorption type grating 31 (hereinafter referred to as a projection image). The projection image is referred to as a G1 image) and is enlarged in proportion to the distance from the X-ray focal point 18b. The grating pitch p ₂ and the interval d ₂ of the second absorption type grating 32 are determined so that the slit portions thereof substantially coincide with the periodic pattern of the bright part of the G1 image at the position of the second absorption type grating 32. Yes. That is, when the distance from the X-ray focal point 18b to the first absorption grating 31 is L ₁ and the distance from the first absorption grating 31 to the second absorption grating 32 is L ₂ , the grating pitch p ₂ and the distance d ₂ are determined so as to satisfy the relationship of the following expressions (1) and (2).

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. In the imaging unit 12 of the present embodiment, the first absorption type grating 31 is configured to project incident X-rays without diffracting, and the G1 image of the first absorption type grating 31 is the first absorption type. because similarly obtained at all positions of the back of the 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. the grating pitch _{p 1} of 31, the grating pitch _p 2, X-ray wavelength of the second absorption-type grating 32 (peak wavelength) lambda, and using the positive integer m, is expressed by the following equation (3).

  Equation (3) is an equation representing the Talbot interference distance when the X-rays emitted from the X-ray source 11 are cone beams. “Atsushi Momose, et al., Japanese Journal of Applied Physics, Vol. 47, No. 10, October 2008, p. 8077 ”.

In the present X-ray imaging system 10, the distance L ₂ is set to a value shorter than the minimum Talbot interference distance Z when m = 1 for the purpose of reducing the thickness of the imaging unit 12. That is, the distance L ₂ is set to a value in the range satisfying the following equation (4).

Incidentally, Talbot distance Z by the following equation (5) and in the case of X-rays emitted from the X-ray source 11 can be regarded as substantially parallel beams, the distance L _2, the value of the range that satisfies the following equation (6) Set to.

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

On the other hand, if the thicknesses h ₁ and h ₂ of the X-ray shielding portions 31b and 32b are excessively increased, X-rays incident obliquely do not easily pass through the slit portion, so-called vignetting occurs, and the X-ray shielding portions 31b and 32b are generated. There is a problem that the effective visual field in the direction (x direction) perpendicular to the stretching direction (strand direction) of the film becomes narrow. Therefore, in view of the field of view secured to define the upper limit of the thickness _h 1, _{h 2.} In order to secure the effective field length V in the x direction on the detection surface of the FPD 30, assuming that the distance from the X-ray focal point 18 b to the detection surface of the FPD 30 is L, the thicknesses h ₁ and h ₂ are shown in FIG. It is necessary to set so that following Formula (7) and (8) may be satisfy | filled from a scientific relationship.

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

  In the imaging unit 12 configured as described above, an intensity-modulated image is formed by superimposing the G1 image of the first absorption-type grating 31 and the second absorption-type grating 32 and is captured by the FPD 30. . The pattern period p1 ′ of the G1 image at the position of the second absorption type grating 32 and the substantial grating pitch p2 ′ (substantial pitch after production) of the second absorption type grating 32 are manufacturing errors and arrangements. Some differences occur due to errors. Among these, the arrangement error means that the substantial pitch in the x direction changes due to the relative inclination and rotation of the first and second absorption gratings 31 and 32 and the distance between the two changes. I mean.

Due to the minute difference between the pattern period p ₁ ′ of the G1 image and the grating pitch p ₂ ′, the image contrast becomes moire fringes. The period T of the moire fringes is expressed by the following equation (9).

  In order to detect the moire fringes with the FPD 30, the arrangement pitch P of the pixels 40 in the x direction needs to satisfy at least the following expression (10), and more preferably satisfies the following expression (11) (here , N is a positive integer).

  Expression (10) means that the arrangement pitch P is not an integral multiple of the moire period T, and it is possible in principle to detect moire fringes even when n ≧ 2. Expression (11) means that the arrangement pitch P is made smaller than the moire period T.

Since the arrangement pitch P of the pixels 40 of the FPD 30 is a value determined by design (generally about 100 μm) and is difficult to change, the magnitude relationship between the arrangement pitch P and the moire period T is adjusted. Adjusts the positions of the first and second absorption gratings 31 and 32 and changes the moire period T by changing at least one of the pattern period p ₁ ′ and the grating pitch p ₂ ′ of the G1 image. It is preferable to do.

  FIG. 6 shows a method of changing the moire cycle T.

The moire period T can be changed by relatively rotating one of the first and second absorption gratings 31 and 32 around the optical axis A. For example, a relative rotation mechanism 50 that rotates the second absorption grating 32 relative to the first absorption grating 31 relative to the optical axis A is provided. When the second absorption type grating 32 is rotated by the angle θ by the relative rotation mechanism 50, the substantial grating pitch in the x direction changes from “p ₂ ′” → “p ₂ ′ / cos θ”. As a result, the moire cycle T changes (FIG. 6A).

As another example, the change of the moire period T is such that either one of the first and second absorption type gratings 31 and 32 is relatively centered about an axis perpendicular to the optical axis A and along the y direction. It can be performed by inclining. For example, a relative tilt mechanism 51 that tilts the second absorption type grating 32 relative to the first absorption type grating 31 about an axis perpendicular to the optical axis A and along the y direction is provided. Provide. When the second absorption type grating 32 is inclined by the angle α by the relative inclination mechanism 51, the substantial lattice pitch in the x direction changes from “p ₂ ′” → “p ₂ ′ × cos α”. As a result, the moire cycle T changes (FIG. 6B).

As another example, the moire period T can be changed by relatively moving one of the first and second absorption gratings 31 and 32 along the direction of the optical axis A. For example, with respect to the first absorption type grating 31, the second absorption type grating 32 is changed so as to change the distance L ₂ between the first absorption type grating 31 and the second absorption type grating 32. A relative movement mechanism 52 that relatively moves along the direction of the optical axis A is provided. When the second absorption type grating 32 is moved to the optical axis A by the movement amount δ by the relative movement mechanism 52, the G1 image of the first absorption type grating 31 projected onto the position of the second absorption type grating 32. The pattern period of “p ₁ ′” → “p ₁ ′ × (L ₁ + L ₂ + δ) / (L ₁ + L ₂ )” changes, and as a result, the moire period T changes (FIG. 6C).

In the X-ray imaging system 10, imaging unit 12 is not the Talbot interferometer as described above, since the distance L ₂ can be freely set, moire by changing the distance L ₂ as relative movement mechanism 52 A mechanism for changing the period T can be suitably employed. The change mechanism (relative rotation mechanism 50, relative tilt mechanism 51, and relative movement mechanism 52) of the first and second absorption gratings 31 and 32 for changing the moiré period T is constituted by an actuator such as a piezoelectric element. Is possible.

  When the subject H is disposed between the X-ray source 11 and the first absorption type grating 31, the moire fringes detected by the FPD 30 are modulated by the subject H. This modulation amount is proportional to the angle of the X-ray deflected by the refraction effect by the subject H. Therefore, the phase contrast image of the subject H can be generated by analyzing the moire fringes detected by the FPD 30.

  Next, a method for analyzing moire fringes will be described.

  FIG. 7 shows one X-ray refracted according to the phase shift distribution Φ (x) of the subject H in the x direction.

  Reference numeral 55 indicates an X-ray path that travels straight when the subject H is not present. The X-ray that travels along the path 55 passes through the first and second absorption gratings 31 and 32 and enters the FPD 30. To do. Reference numeral 56 indicates an X-ray path refracted and deflected by the subject H when the subject H exists. X-rays traveling along this path 56 are shielded by the second absorption type grating 32 after passing through the first absorption type grating 31.

  The phase shift distribution Φ (x) of the subject H is expressed by the following equation (12), where n (x, z) is the refractive index distribution of the subject H, and z is the direction in which the X-ray proceeds.

  The G1 image projected from the first absorptive grating 31 to the position of the second absorptive grating 32 is displaced in the x direction by an amount corresponding to the refraction angle φ due to refraction of X-rays at the subject H. become. This displacement amount Δx is approximately expressed by the following equation (13) based on the fact that the X-ray refraction angle φ is very small.

  Here, the refraction angle φ is expressed by Expression (14) using the X-ray wavelength λ and the phase shift distribution Φ (x) of the subject H.

  Thus, the displacement amount Δx of the G1 image due to the refraction of X-rays at the subject H is related to the phase shift distribution Φ (x) of the subject H. The amount of displacement Δx is expressed by the following equation with the phase shift amount ψ of the signal output from each pixel 40 of the FPD 30 (the phase shift amount of the signal of each pixel 40 with and without the subject H): It is related as shown in (15).

  Therefore, by obtaining the phase shift amount ψ of the signal of each pixel 40, the refraction angle φ is obtained from the equation (15), and the differential amount of the phase shift distribution Φ (x) is obtained using the equation (14). Is integrated with respect to x, a phase shift distribution Φ (x) of the subject H, that is, a phase contrast image of the subject H can be generated. In the present X-ray imaging system 10, the phase shift amount ψ is calculated using a fringe scanning method described below.

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

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

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

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

When imaging 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 (16).

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

  Here, arg [] means 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 (18).

  FIG. 9 shows the 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. 10 indicates a change in the signal value when the subject H does not exist, and a solid line in FIG. 10 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 (14), the phase shift is obtained by integrating the refraction angle φ (x) along the x-axis. A distribution Φ (x) is obtained. In the above description, the y coordinate in the y direction of the pixel 40 is not taken into consideration. However, by performing the same calculation for each y coordinate, a two-dimensional phase shift distribution Φ (x , Y).

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

  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.

  FIG. 10 is a flowchart showing the operation of the radiation imaging system of FIG.

  This operation flow is based on the premise that the X-ray tube is being air-cooled by the X-ray tube cooling unit 27 at the start of the operation flow.

  In the X-ray imaging system 10, imaging conditions (imaging menu) are input. When an imaging instruction for a subject is issued after the imaging conditions are input, the control device 20 registers the input imaging conditions (S110).

  Next, the control device 20 transmits a request to measure the temperature of the X-ray tube to the X-ray source control unit 17. When receiving this request, the X-ray source control unit 17 causes the X-ray tube thermometer 28 to measure the temperature, and acquires the measured temperature value. The X-ray source control unit 17 transmits the acquired temperature value of the X-ray tube to the control device 20 (S111).

  The control device 20 predicts the temperature of the X-ray tube after imaging the subject using the received X-ray tube temperature value and the registered imaging conditions (S112). A method for predicting the temperature of the X-ray tube will be described later.

  The control device 20 determines whether or not the predicted temperature of the X-ray tube exceeds the allowable temperature (S113).

  When the predicted temperature of the X-ray tube exceeds the allowable temperature (S113-NO), the photographing of the subject is not permitted, the process returns to S111, and the temperature of the X-ray tube is measured again (S111). The tube tube temperature prediction (S112) is compared with the allowable temperature (S113).

  When the predicted temperature of the X-ray tube exceeds the allowable temperature, the control device 20 as an example of the warning information unit displays on the monitor 24 or the like that there is a possibility of failure of the X-ray tube. May be. As a result, the user can grasp the state of the X-ray tube instantaneously, and can efficiently recompose the scheduled imaging time of the subject.

  If the predicted temperature of the X-ray tube is below the allowable temperature (S113—YES), the control device 20 transmits a request to stop the air cooling of the X-ray tube to the X-ray source control unit 17. When receiving this request, the X-ray source control unit 17 stops the fan of the X-ray tube cooling unit 27 and acquires a stop signal (S114). Based on the acquired stop signal, the X-ray source control unit 17 transmits an air cooling stop completion notification indicating that the air cooling of the X-ray tube has been stopped to the control device 20.

  When receiving the air cooling stop completion notification, the control device 20 transmits a request to measure the vibration of the X-ray tube to the X-ray source control unit 17. When receiving this request, the X-ray source control unit 17 causes the X-ray tube vibrometer 29 to measure the vibration of the X-ray tube, and acquires the measured vibration value. The X-ray source control unit 17 transmits the acquired X-ray tube vibration value to the control device 20 (S115).

  The control device 20 determines whether or not the received vibration value of the X-ray tube exceeds a threshold value (S116).

  If the received vibration value of the X-ray tube exceeds the threshold value (NO in S116), the photographing of the subject is not permitted, and the process returns to S115 and X-ray tube vibration measurement is performed again.

  If the received vibration value of the X-ray tube is below the threshold value (S116-YES), the control device 20 determines that a phase image can be captured (S117), and the X-ray imaging system 10 Then, phase image shooting is performed (S118).

  FIG. 11 is a graph for explaining a method of predicting the temperature of the radiation source.

  The X-ray tube of the X-ray source 11 is an anode rotating type, and a large amount of heat is accumulated when an electron beam collides with the rotating anode 18a. That is, if the temperature of the rotary anode 18a rises and this temperature exceeds an allowable value, it causes a failure of the X-ray tube. In particular, in the fringe scanning method, as shown in FIG. 8, M times of imaging are performed, so that the temperature of the rotating anode 18a (X-ray tube) is likely to rise. Therefore, in the X-ray imaging system 10, the control device 20 as an example of a temperature prediction unit predicts the temperature of the X-ray tube after imaging of the subject.

  In the graph, the vertical axis indicates the temperature of the X-ray tube, and the horizontal axis indicates time. Here, a process in which the control device 20 transmits a temperature measurement request to the X-ray source control unit 17 at time ta and predicts the temperature of the X-ray tube will be described.

  First, the control device 20 transmits a temperature measurement request to the X-ray source control unit 17 and acquires the temperature Ta of the X-ray tube through the X-ray tube thermometer 28.

  Next, the control device 20 acquires shooting conditions used for shooting the subject. The imaging conditions include the value of the tube voltage applied to the X-ray tube, the value of the tube current flowing through the X-ray tube, and the irradiation time tc for irradiating the subject with X-rays during imaging (M times for phase imaging) Irradiation time required for photographing). The control device 20 obtains the temperature Tc of the X-ray tube that rises at the irradiation time tc based on the temperature rise characteristic according to the imaging condition, and is the temperature of the X-ray tube after the imaging of the subject (time ta + tc). The predicted temperature Ta + Tc is obtained.

  In this case, the predicted temperature Ta + Tc exceeds the allowable temperature Tl set as a temperature value at which the X-ray tube does not fail. For this reason, the control device 20 determines that there is a possibility of failure of the X-ray tube due to imaging of the subject, and prohibits imaging by the X-ray imaging system 10.

  On the other hand, it is assumed that the temperature Tb of the X-ray tube is acquired at time tb, and the predicted temperature Tb + Tc is obtained using the same imaging conditions as described above. In this case, since the predicted temperature Tb + Tc is lower than the allowable temperature Tl, the control device 20 determines that there is no possibility of failure of the X-ray tube by imaging the subject, and permits imaging by the X-ray imaging system 20. To do.

  As described above, in the present X-ray imaging system 10, in the phase image capturing, the air cooling of the X-ray tube cooler 27 is stopped before the subject is imaged. Thereby, the vibration of the X-ray source due to the air cooling can be prevented, and the image quality of the phase contrast image by imaging can be improved. In addition to simply stopping the air cooling of the X-ray tube cooler 27, the X-ray tube vibration meter 29 measures vibrations of the X-ray tube in advance. Thereby, after confirming that the vibration of the X-ray source due to air cooling has stopped, the subject can be photographed, and the image quality of the phase contrast image by the photographing can be improved. Further, before the air cooling of the X-ray tube cooler 27 is stopped, the temperature of the X-ray tube is measured. Thereby, even if the subject is photographed in a state where the air cooling of the X-ray tube cooler 27 is stopped, the possibility that the X-ray tube is broken can be reduced.

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

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

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

  In the X-ray imaging system 10, the X-ray source control unit 17 cools the X-ray tube when the temperature value of the X-ray tube obtained from the X-ray tube thermometer 28 is not more than a predetermined value. The fan of the device 27 may be automatically stopped.

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

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

  The X-ray imaging system 60 is an X-ray diagnostic apparatus that images a subject (patient) H in a lying position, and includes a bed 61 on which the subject H is placed in addition to the X-ray source 11 and the imaging unit 12. Prepare. Since the X-ray source 11 and the imaging unit 12 have the same configuration as that of the above embodiment, the same reference numerals as those of the above embodiment are given to the respective components. Only differences from the above embodiment will be described below. Since other configurations and operations are the same as those in the above embodiment, the description thereof is omitted.

  In the present embodiment, the imaging unit 12 is attached to the lower surface side of the top plate 62 so as to face the X-ray source 11 with the subject H interposed therebetween. One X-ray source 11 is held by an X-ray source holding device 14, and the X-ray irradiation direction is set downward by an angle changing mechanism (not shown) of the X-ray source 11. In this state, the X-ray source 11 irradiates the subject H lying on the top plate 62 of the bed 61 with X-rays. Since the X-ray source holding device 14 enables the X-ray source 11 to move up and down by extending and contracting the support 14b, the distance from the X-ray focal point 18b to the detection surface of the FPD 30 can be adjusted by this up and down movement. .

As described above, the imaging unit 12 includes a first absorption type grating 31 it is possible to shorten the distance L ₂ between the second absorption-type grating 32, since it can be thinned, the bed 61 The leg part 63 which supports the top plate 62 can be shortened, and the position of the top plate 62 can be lowered. For example, the imaging unit 12 is preferably thinned, and the position of the top plate 62 is preferably set to a height that allows the subject (patient) H to sit down easily (for example, about 40 cm above the floor). In addition, it is preferable to lower the position of the top plate 62 in order to secure a sufficient distance from the X-ray source 11 to the imaging unit 12.

  Contrary to the positional relationship between the X-ray source 11 and the imaging unit 12, the X-ray source 11 is attached to the bed 61, and the imaging unit 12 is installed on the ceiling, so that the subject H can be photographed in the supine position. It is also possible to do this.

  13 and 14 show the configuration of another example of a radiation imaging system for explaining an embodiment of the present invention.

    The X-ray imaging system 70 is an X-ray diagnostic apparatus that enables imaging of a subject (patient) H in a standing position and a standing position, and the X-ray source 11 and the imaging unit 12 include a swivel arm 71. Is held by. The turning arm 71 is connected to the base 72 so as to be turnable. Since the X-ray source 11 and the imaging unit 12 have the same configuration as that of the above embodiment, the same reference numerals as those of the first embodiment are given to the respective components. Only differences from the above embodiment will be described below. Since other configurations and operations are the same as those in the above embodiment, the description thereof is omitted.

  The swivel arm 71 includes a U-shaped portion 71a having a substantially U-shape and a straight linear portion 71b connected to one end of the U-shaped portion 71a. The photographing part 12 is attached to the other end of the U-shaped part 71a. A first groove 73 is formed in the linear portion 71 b along the extending direction, and the X-ray source 11 is slidably attached to the first groove 73. The X-ray source 11 and the imaging unit 12 face each other, and the distance from the X-ray focal point 18b to the detection surface of the FPD 30 can be adjusted by moving the X-ray source 11 along the first groove 73. it can.

  Further, the base 72 is formed with a second groove 74 extending in the vertical direction. The swivel arm 71 is movable in the vertical direction along the second groove 74 by a coupling mechanism 75 provided at a connection portion between the U-shaped portion 71a and the linear portion 71b. Further, the turning arm 71 can be turned around the rotation axis C along the y direction by the connecting mechanism 75. From the standing imaging state shown in FIG. 11, the swivel arm 71 is rotated 90 ° clockwise around the rotation axis C, and the imaging unit 12 is placed under a bed (not shown) on which the subject H is placed. By placing the, it is possible to shoot the supine position. Note that the turning arm 71 is not limited to 90 ° rotation, and can rotate at any angle, and can shoot in directions other than standing shooting (horizontal direction) and lying-down shooting (vertical direction). Is possible.

  In this embodiment, since the X-ray source 11 and the imaging unit 12 are held by the turning arm 71, the distance from the X-ray source 11 to the imaging unit 12 can be set easily and accurately compared to the above embodiment. Can do.

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

  A mammography apparatus 80 in which the present invention is applied to mammography (X-ray mammography) is an apparatus that captures an X-ray image (phase contrast image) of the 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.

  In addition, since the X-ray source 11 and the imaging | photography part 12 are the structures similar to the thing of the said embodiment, the code | symbol same as the said embodiment is attached | subjected to each component. Since other configurations and operations are the same as those in the above embodiment, the description thereof is omitted.

  FIG. 16 shows a configuration of a modified example of the radiation imaging system of FIG.

The mammography apparatus 90 differs from the mammography apparatus 80 of the fourth embodiment only in that the first absorption 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 does not include the first absorption type grating 31, and includes the FPD 30, the second absorption type grating 32, and the scanning mechanism 33.

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

  In this embodiment, since the subject B is irradiated with X-rays whose dose is almost halved by the shielding by the first absorption type grating 31, the exposure amount of the subject B is the same as that of the fourth embodiment. It can be reduced to about half of the case. Note that, as in the present embodiment, disposing the subject between the first absorption type grating 31 and the second absorption type grating 32 is not limited to the mammography apparatus, but can be applied to other X-ray imaging systems. Is possible.

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

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

In the above embodiment, when the distance from the X-ray source 11 to the FPD 30 is set to a distance (1 to 2 m) as set in a general hospital radiographing room, the focal size of the X-ray focal point 18b (general (Between about 0.1 mm and 1 mm), the blur of the G1 image may affect, and the image quality of the phase contrast image may be 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 this embodiment, 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 (that is, a third absorption type grating) having the same configuration as the first and second absorption type gratings 31 and 32 provided in the photographing unit 12, and is unidirectional (this embodiment). In the embodiment, the plurality of X-ray shielding portions extending in the y direction) are in the same direction as the X-ray shielding portions 31b and 32b of the first and second absorption gratings 31 and 32 (in the present embodiment, the x direction). Arranged periodically. The multi-slit 103 partially shields the radiation from the X-ray source 11, thereby reducing the effective focal size in the x direction and forming a large number of point light sources (dispersed light sources) in the x direction. It is aimed.

  The lattice pitch p3 of the multi-slit 103 needs to be set so as to satisfy the following equation (19), where the distance from the multi-slit 103 to the first absorption-type lattice 31 is L3.

  In this embodiment, since the position of the multi slit 103 is substantially the X-ray focal position, the grating pitch p2 and the interval d2 of the second absorption grating 32 are expressed by the following equations (20) and (21). Determined to satisfy the relationship.

  Further, in this embodiment, in order to secure the effective field length V in the x direction on the detection surface of the FPD 30, if the distance from the multi slit 103 to the detection surface of the FPD 30 is L ′, the first and second The thicknesses h1 and h2 of the X-ray shielding portions 31b and 32b of the absorption gratings 31 and 32 are determined so as to satisfy the following expressions (22) and (23).

  Expression (19) indicates that the projection image (G1 image) of the X-rays emitted from the point light sources dispersedly formed by the multi-slit 103 by the first absorption type grating 31 is the position of the second absorption type grating 32. This is a geometric condition for matching (overlapping). As described above, in the present embodiment, the G1 images based on the plurality of point light sources formed by the multi-slit 103 are superimposed, thereby improving the image quality of the phase contrast image without reducing the X-ray intensity. it can.

  The multi slit 103 described above can be applied to any of the above embodiments.

  FIG. 18 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, the second absorption grating 32 is provided independently of the FPD 30, but by using an X-ray image detector having a configuration disclosed in Japanese Patent Application Laid-Open No. 2009-133823. The second absorption type grating can be eliminated. 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 the phases thereof 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.

Each of the first to sixth linear electrode groups 122 to 127 is formed by periodically arranging linear electrodes extending in the y direction at a predetermined pitch p ₂ in the x direction. The pattern period p of the G1 image at the substantial pitch p ₂ ′ (substantial pitch after manufacture) of the arrangement pitch p ₂ of the linear electrodes and the position of the charge collection electrode 121 (position of the X-ray image detector). The relationship between ₁ ′ and the arrangement pitch P of the pixels 120 in the x direction is similar to the second absorption grating 32 of the X-ray imaging system 10 described above, and the period T of the moire fringes represented by the equation (7). Therefore, it is necessary to satisfy the formula (8), and it is preferable to satisfy the formula (9).

  Furthermore, each pixel 120 is provided with a switch group 128 for reading out the charges collected by the charge collection electrode 121. The switch group 128 includes TFT switches provided in each of the first to sixth linear electrode groups 121 to 126. 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-time shooting. A phase contrast image can be generated based on these six types of fringe images.

  For example, when the X-ray image detector configured as described above is applied to the X-ray imaging system 10 described above, the second absorption type grating 32 is not required from the imaging unit 12, and a plurality of images can be obtained by one imaging. Since a phase component fringe image can be acquired, physical scanning for fringe scanning becomes unnecessary, and the scanning mechanism 33 can be eliminated. Thereby, it is possible to reduce the cost and further reduce the thickness of the photographing unit. In addition, it is also possible to use the other structure of Unexamined-Japanese-Patent No. 2009-133823 for the structure of an electric charge collection electrode instead of the said structure.

  FIG. 19 shows a control block of another example of the radiation imaging system for explaining the embodiment of the present invention, and FIG. 20 is a flowchart showing the operation of the radiation imaging system of FIG.

  The X-ray imaging system 130 is different from the X-ray imaging system 10 of the above-described embodiment in that a retraction mechanism 34 is provided in the imaging unit 12. Since other configurations are the same as those in the above embodiment, the description thereof is omitted.

  The retracting mechanism 34 retracts the first absorption grating 31 of the imaging unit 12 from the X-ray irradiation field emitted from the X-ray source 11. Thus, the X-ray imaging system 130 is a dual-purpose system that can acquire not only the phase contrast image of the subject but also the X-ray absorption image of the subject.

  The operation of the X-ray imaging system 130 will be described. As described above, this operation flow is based on the premise that the X-ray tube is cooled by the X-ray tube cooling unit 27 at the start of the operation flow.

  First, in the X-ray imaging system 130, in addition to the imaging conditions, the phase imaging mode for acquiring the phase contrast image of the subject or the normal imaging mode for acquiring the absorption image of the subject is selected. When a shooting instruction is issued after a shooting condition is input and a shooting mode is selected, the control device 20 registers the input shooting condition (S200), and the selected shooting mode is a phase shooting. It is determined whether or not the mode is selected (S210).

  When the selected shooting mode is not the phase shooting mode (S210-NO), it is determined that the shooting mode is the normal shooting mode, and the process proceeds to S220. From S220 to S240, the temperature of the X-ray tube is measured in the same manner as S110 to S113 described above (S220), and the temperature of the X-ray tube after subject photographing in normal image photographing is predicted (S230). The predicted temperature of the X-ray tube is compared with the allowable temperature of the X-ray tube (S240). When the predicted temperature of the X-ray tube exceeds the allowable temperature (S240-NO), the photographing of the subject is not permitted, the process returns to S220, and the temperature of the X-ray tube is measured again (S220). The tube tube temperature prediction (S230) is compared with the allowable temperature (S240). If the predicted temperature of the X-ray tube does not exceed the allowable temperature (S240-YES), the control device 20 determines that an absorption image (normal image) can be taken (S250), and X-rays are obtained. The imaging system 10 performs normal image capturing after the first absorption grating 31 is retracted from the X-ray irradiation field by the retracting mechanism 34 (S260).

  When the selected imaging mode is the phase imaging mode (S210-YES), the processing from S111 to S118 is performed as described above, assuming that it is the phase imaging mode.

  Thus, at the time of taking an absorption image, the air cooling by the X-ray tube cooler 27 is continued without stopping the air cooling of the X-ray tube by the X-ray tube cooler 27. This is because the vibration caused by air cooling by the X-ray tube cooler 27 hardly affects the image quality of the absorption image. That is, in the present X-ray imaging system 130, the possibility of failure of the X-ray tube is reduced by securing the air cooling time of the X-ray tube according to the selected imaging mode.

FIG. 21 is a block diagram illustrating a configuration of a calculation unit that generates a radiographic image regarding another example of the radiation imaging system for describing the embodiment of the present invention. FIG. 21 is a block diagram of the calculation unit of the radiation imaging system. It is a graph which shows the signal of the pixel of the radiographic image detector for demonstrating a process.

  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. Since other configurations are the same as those of the X-ray imaging system 10 described above, description thereof will be omitted. The arithmetic processing unit 190 includes a phase contrast image generation unit 191, an absorption image generation unit 192, and a small angle scattered image generation unit 193. These all perform arithmetic processing based on image data obtained at M scanning positions of k = 0, 1, 2,..., M−1. Among these, the phase contrast image generation unit 191 generates a phase contrast image according to the above-described procedure.

The absorption image generation unit 192 generates an absorption image by averaging the pixel data I _k (x, y) obtained for each pixel with respect to k and calculating an average value as shown in FIG. To do. The average value may be calculated by simply averaging the pixel data I _k (x, y) with respect to k. However, when M is small, the error increases, so that the pixel data Ik (x , Y) may be fitted with a sine wave, and then the 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) without a subject. This absorption image reflects the transmittance unevenness of the detection system (including information such as the transmittance unevenness of the grid and the influence of the absorption of the dose detector). Therefore, a correction coefficient map for correcting the transmittance unevenness of the detection system can be created from this image. Absorption of the subject, in which an absorption image is created from a group of images obtained by shooting in the state of the subject (main shooting), and the above-described correction coefficient is applied to each pixel, thereby correcting the transmittance unevenness of the detection system. An image can be obtained.

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

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

  According to the present X-ray imaging system, an absorption image and a small angle scattered image are generated from a plurality of images acquired for the phase contrast image of the subject. There is no deviation, and it is possible to superimpose the phase contrast image with the absorption image and the small-angle scattered image, and the burden on the subject is reduced as compared with the case of separately shooting for the absorption image and the small-angle scattered image. be able to. Further, unlike the X-ray imaging system 130, the phase contrast image and the absorption image can be acquired without providing the retracting mechanism 34, so that the configuration of the imaging unit 12 can be simplified.

  FIG. 23 is a schematic configuration diagram of another example of a radiation imaging system for describing an embodiment of the invention.

  The X-ray phase imaging system of this example is formed by a first grating 131 that passes X-rays emitted from the X-ray source 11 and forms a first periodic pattern image, and the first grating 131. A second grating 132 that forms a second periodic pattern image by intensity-modulating the first periodic pattern image, and an X-ray image detector that detects the second periodic pattern image formed by the second grating 132 (Radiation image detector) 240 and a phase contrast that obtains a fringe image based on the second periodic pattern image detected by the X-ray image detector 240 and generates a phase contrast image based on the obtained fringe image An image generation unit 260. 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 normal X-ray energy region for medical diagnosis is about 1 μm to several μm. An amplitude modulation type grating can also be used as the first grating 131.
On the other hand, the second grating 132 is preferably an amplitude modulation type grating.

Here, when the X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam, the self-image of the first grating 131 formed through the first grating 131 is: It is enlarged in proportion to the distance from the X-ray source 11. In this example, the grating pitch P ₂ and the interval d ₂ of the second grating 132 are such that the slit portion is a periodic pattern of the bright part of the self-image of the first grating 131 at the position of the second grating 132. It is determined so as to substantially match. That is, when the distance from the focal point of the X-ray source 11 to the first grating 131 is L ₁ and the distance from the first grating 131 to the second grating 132 is L ₂ , the second grating pitch P ₂ and distance d ₂ are determined so as to satisfy the relationship of the above (1) and (2).

Note that when the X-rays emitted from the X-ray source 11 are parallel beams, P ₂ = P ₁ , d
₂ = determined to satisfy d ₁

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

FIG. 24 is a diagram showing a schematic configuration of an optical reading type radiation image detector.
24A is a perspective view of the X-ray image detector 240 of this example, FIG. 24B is a cross-sectional view of the XZ-plane of the X-ray image detector shown in FIG. 24A, and FIG. FIG. 25 is a YZ plane cross-sectional view of the X-ray image detector shown in FIG.

  As shown in FIGS. 24A to 24C, the X-ray image detector 240 of the present example irradiates the first electrode layer 241 that transmits X-rays and the X-ray that transmits through the first electrode layer 241. The photoconductive layer for recording 242 that generates charges by receiving the charge, acts as an insulator for the charges of one polarity among the charges generated in the photoconductive layer for recording 242 and for the charges of the other polarity In this case, a charge transport layer 244 acting as a conductor, a reading photoconductive layer 245 that generates charges when irradiated with reading light, and a 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 may be any material that transmits 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.) can be used with a thickness of 50 to 200 nm, and Al or Au with 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. What has a-Se as a main component 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 containing at least one of MgPc (Magnesium phtalocyanine), VoPc (phase II of Vanadyl phthalocyanine), CuPc (Cupper phtalocyanine) and the like as a main component 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 FIGS. 24A and 24B, 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 material that transmits reading light and has conductivity. For example, as with the first electrode layer 241, ITO, IZO, or IDIXO can be used. And the thickness is about 100-200 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 adjacent transparent linear electrodes 246a and light shielding linear electrodes 246b. That is, as shown in FIG. 24B, 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.

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

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

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

Further, the distance Z ₂ (Talbot interference distance Z) between the first grating 131 and the second grating 132 is the following when the first grating 131 is a phase modulation type grating that applies 90 ° phase modulation. Equation (24) should be almost satisfied.

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

  When the first grating 131 is a phase modulation type grating that applies 180 ° phase modulation, or when the first grating 131 is an amplitude modulation type grating, the above formula ( 3) must be almost satisfied.

Further, the thicknesses h ₁ and h ₂ of the first and second gratings 131 and 132 also satisfy the expressions (7) and (8) described above with respect to the first and second absorption gratings 31 and 32. It is necessary to set as follows.

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

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

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

  Specifically, as shown in FIG. 26, a periodic pattern image formed on the position of the second grating 132 by the pitch of the second grating 132 and the first grating 131 (hereinafter referred to as self of the first grating 131). The pitch of the image G1) is p, the relative rotation angle of the self-image of the first grating 131 with respect to the second grating 132 in the XY plane is θ, and the image resolution in the sub-scanning direction of the phase contrast image is D. Assuming that (= Dy × M), by setting the rotation angle θ to satisfy the following expression (25), the self-image G1 of the first grating 131 with respect to the length of the image resolution D in the sub-scanning direction. And the phase of the second grating 132 are shifted by n periods. FIG. 26 shows a case where M = 5 and n = 1.

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

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

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

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

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

  For example, if p = 5 μm, D = 50 μm, and n = 1 in the above equation (25), the theoretical rotation angle θ is about 5.7 °. Then, the actual rotation angle θ ′ of the self-image of the first grating 131 relative to the second grating 132 can be detected by, for example, the self-image of the first grating and the moire pitch by the second grating 132. .

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

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

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

  Next, the operation of the X-ray phase imaging apparatus 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 radiated to the first grating 131 are diffracted by the first grating 131, so that the first grating 131 self is at a predetermined distance in the optical axis direction of the X-ray from the first grating 131. Form an image.

  For example, when the first grating 131 is a phase modulation type grating that applies 90 ° phase modulation, the above equation (24) (in the case of a 180 ° phase modulation type grating or an intensity modulation type grating, the above equation (3)). While the self-image of the first grating 131 is formed by the subject H, the wavefront of the X-ray incident on the first grating 131 is distorted by the subject H, and the self-image of the first grating 131 is deformed accordingly. is doing.

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

  Here, the operation of image detection and readout in the X-ray image detector 240 will be described.

  First, as shown in FIG. 28A, in the 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 of the first grating 131 and the second image X-rays whose intensity is modulated by superimposing with the grating 132 are irradiated from the first electrode layer 241 side of the X-ray image detector 240.

  The X-rays irradiated to the X-ray image detector 240 pass through the first electrode layer 241 and are 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 portion 243 formed at the interface between the recording photoconductive layer 242 and the charge transport layer 244 (see FIG. 28B).

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

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

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

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

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

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

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

  The method for generating the phase contrast image in this example is the same as the content already described with reference to the equations (12) to (18), and thus description thereof is omitted.

In the above-described configuration in which the first grating 131 and the second grating 132 are tilted, both the first grating 131 and the second grating 132 are configured as absorption (amplitude modulation type) gratings, and the Talbot interference effect is obtained. Irrespective of the presence or absence of, it is good also as a structure which projects the radiation which passed the slit part geometrically. In this case, by a distance d ₁ of the first grating 131 and a distance d ₂ of the second grating 132 is sufficiently larger than a peak wavelength of X-rays emitted from the X-ray source 11, illumination Most of the X-rays are configured to pass through while maintaining straightness without being diffracted by the slit portion. For example, when tungsten is used as the target of the X-ray source and the tube voltage is 50 kV, the peak wavelength of the X-ray is about 0.4 mm. In this case, the distance d ₁ of the first grating 131 spacing d ₂ of the second grating 132, the geometrically not most of the radiation is diffracted by the slit portion be about 1μm~10μm projection Is done. The relationship between the lattice pitch _{P 2} of the first grating pitch _{P 1} and a second grating 132 of the grating 131, for a distance _{d 1} of the first grating 131 and the relationship between the distance _{d 2} of the second grating 132 Is the same as that when the first grating 131 is a phase modulation type grating. Further, the inclination of the first grating 131 with respect to the second grating 132 is also the same as in the above example, and the generation of the phase contrast image is performed in the same manner as in the above example.

  In the above example, as the X-ray image detector 240, a so-called optical reading type X-ray image detector in which an image signal is read out by scanning linear reading light emitted from the linear reading light source 250 is used. However, the present invention is not limited to this. For example, as described in JP-A-2002-26300, a large number of TFT switches are arranged two-dimensionally, and image signals are read by turning on and off the TFT switches. An X-ray image detector using a TFT switch or an X-ray image detector using a CMOS may be used.

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

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

  Similarly to the above example, when M striped images are used to generate the phase contrast image, the M pixel circuit rows have one image resolution D in the sub-scanning direction of the phase contrast image. The first grating 131 is inclined with respect to the second grating 132. The specific rotation angle of the first grating 131 is calculated by the above equation (25) as in the above example.

  In the above equation (25), for example, when M = 5 and n = 1 and the rotation angle θ of the first grating 131 is set, one pixel circuit 270 of FIG. Image signals obtained by dividing the intensity modulation of the period into five can be detected. That is, five pixel circuit rows connected to the five gate electrodes 273 shown in FIG. Each 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 CMOS, for example, a pixel circuit 280 that generates visible light upon receiving X-ray irradiation and photoelectrically converts the visible light to detect a charge signal is shown in FIG. As shown in FIG. 3, a plurality of two-dimensional arrays can be used. The X-ray image detector using CMOS is provided for each pixel circuit row, and includes a large number of gate electrodes 282 and reset electrodes from which a drive signal for driving a signal readout circuit included in the pixel circuit 280 is output. 284 and a plurality of data electrodes 283 that are provided for each pixel circuit column and output a charge signal read from the signal reading circuit of each pixel circuit 280. The gate electrode 282 and the reset electrode 284 are connected to a row selection scanning unit 285 that outputs a drive signal to the signal readout circuit, and the data electrode 283 performs predetermined processing on the charge signal output from each pixel circuit. A signal processing unit 286 to be applied is connected.

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

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

  Since the upper electrode 808 needs to make light having a wavelength of 550 nm incident on the photoelectric conversion film 807, the upper electrode 808 is made of a conductive material transparent to the incident light. 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 so as to penetrate the insulating film 803. The signal readout circuit 801 is configured by a known CMOS circuit.

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

  Similarly to the above example, when M striped images are used to generate the phase contrast image, the M pixel circuit rows have one image resolution D in the sub-scanning direction of the phase contrast image. The first grating 131 is inclined with respect to the second grating 132. The specific rotation angle of the first grating 131 is calculated by the above equation (25) as in the above example.

  In the above formula (25), for example, when M = 5 and n = 1 and the rotation angle θ of the first grating 131 is set, one pixel circuit 280 in FIG. Image signals obtained by dividing the intensity modulation of the period into 5 can be detected. That is, image signals of five different fringe images are obtained by five pixel circuit rows connected to the five gate electrodes 282 shown in FIG. Each 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.

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

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

  Moreover, since a plurality of fringe image signals can be acquired by one imaging, the possibility of failure of the X-ray tube can be reduced.

FIG. 36 is a schematic configuration diagram of another example of a radiation imaging system for describing an embodiment of the invention.

  As shown in FIG. 36, the X-ray phase imaging apparatus detects a grating 131 that forms a periodic pattern image by passing X-rays emitted from the X-ray source 11, and a periodic pattern image formed by the grating 131. At the same time, an X-ray image detector (radiation image detector) 340 that modulates the intensity of the periodic pattern image, and a movement that moves the X-ray image detector 340 in a direction orthogonal to the extending direction of the linear electrode. A mechanism 333 and a phase-contrast image generation unit 260 that generates a phase-contrast image based on a fringe image obtained by intensity-modulating the periodic pattern image in the X-ray image detector 340 are provided.

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

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

  37A is a perspective view of the X-ray image detector 340 of this example, FIG. 37B is a cross-sectional view of the XZ-plane of the X-ray image detector shown in FIG. 37A, and FIG. FIG. 38 is a YZ plane cross-sectional view of the X-ray image detector shown in FIG.

  As shown in FIGS. 37A to 37C, the X-ray image detector 340 of the present example irradiates the first electrode layer 241 that transmits X-rays and the X-ray that transmits through the first electrode layer 241. The photoconductive layer for recording 242 that generates charges by receiving the charge, acts as an insulator for the charges of one polarity among the charges generated in the photoconductive layer for recording 242 and for the charges of the other polarity In this case, a charge storage layer 343 acting as a conductor, a reading photoconductive layer 245 that generates charges when irradiated with reading light, and a second electrode layer 246 are laminated in this order. Note that each of the above layers is formed on the glass substrate 247 in order from the second electrode layer 246.

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

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

  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, the charge storage layer 343 in this example is parallel to the extending direction of the transparent linear electrode 246a and the light shielding linear electrode 246b of the second electrode layer 246, as shown in FIGS. It is divided like a line.

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

Specifically, when the X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam, the self-image of the grating 131 formed through the grating 131 is the X-ray source 11. Enlarged in proportion to the distance from In this example, the arrangement pitch P ₂ and the interval d ₂ of the charge storage layer 343 are such that the portion of the linear charge storage layer 343 is the period of the bright portion of the self-image of the lattice 131 at the position of the charge storage layer 343. It is determined so as to substantially match the pattern. That is, the grating pitch of the grating 131 is P ₁ , the distance between the X-ray shielding portions of the grating 131 is d ₁ , the distance from the focal point of the X-ray source 11 to the grating 131 is L ₁ , and the grating 131 to the X-ray image detector 340 When the distance to the detection surface is L ₂ , the arrangement pitch P ₂ and the interval d ₂ of the charge storage layer 343 are determined so as to satisfy the relationship of the above (1) and Expression (2).

  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 adjacent transparent linear electrodes 246a and light shielding linear electrodes 246b. That is, as shown in FIG. 37B, 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.

  And the X-ray phase imaging device of this example was extended in the direction (X direction) orthogonal to the extending | stretching direction of the transparent linear electrode 246a and the light shielding linear electrode 246b, as shown to FIG. 37 (A). A linear reading light source 250 is provided.

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

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

Further, the distance Z ₂ (Talbot interference distance Z) between the grating 131 and the detection surface of the X-ray image detector 340 is equal to the above formula when the grating 131 is a phase modulation type grating that applies 90 ° phase modulation. (24) must be almost satisfied.

  Further, when the grating 131 is a phase modulation type grating that applies 180 ° phase modulation, or when the grating 131 is an amplitude modulation type grating, the above formula (3) must be substantially satisfied with respect to the Talbot interference distance Z. 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 phase imaging apparatus of this example will be described.

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

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

  Here, 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. 38A, 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 of the grating 131 is The X-ray image detector 340 is irradiated from the first electrode layer 241 side.

  The X-rays applied to the X-ray image detector 340 pass through the first electrode layer 241 and are applied 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 (see FIG. 38B).

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

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

  Then, as shown in FIG. 38, 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 <b> 1, and an image signal of one 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 (12) to (18), 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, the X-ray phase contrast imaging apparatus described above, so that the distance Z ₂ from the grating 131 to the detection surface of the X-ray image detector 340 becomes Talbot interference distance, the above equation (24) and formula (3) However, the grating 131 may project the incident X-rays without diffracting them. According to this configuration, since the projected image projected through the grating 131 is obtained similarly at all positions behind the grating 131, the distance from the grating 131 to the detection surface of the X-ray image detector 340 is obtained. the Z _2, can be set independently of the Talbot interference distance.

Next, a modification of the above-described X-ray phase image capturing apparatus will be described. In the X-ray phase image capturing apparatus described above, the X-ray image detector 340 is translated by the moving mechanism 333, and M striped image signals are acquired by capturing an X-ray image at each position. The X-ray phase imaging apparatus of this example is configured to be able to acquire M striped image signals by capturing an X-ray image once without requiring the moving mechanism 333 as described above. is there.
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 the grating 131. The stretching direction and the stretching 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 in 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, it is not always necessary to have this layer configuration. For example, as shown in FIG. A linear charge storage layer 343 may be provided so as to be in direct contact with the recording medium, and a recording photoconductive layer 242 may be provided on the charge storage layer 343. The recording photoconductive layer 242 also functions as a reading photoconductive layer.

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

  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 below.

  First, as shown in FIG. 41A, 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 supply 400, the X-ray carrying the self-image of the grating 131 is The X-ray image detector 360 is irradiated from the first electrode layer 241 side.

  The X-rays applied to the X-ray image detector 340 pass through the first electrode layer 241 and are applied 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 accumulated in the charge accumulation layer 343 (see FIG. 41B). Note that since the linear charge storage layer 343 in contact with the second electrode layer 246 is an insulating film, charges that have reached the charge storage layer 343 are captured there and go to the second electrode layer 246. Can't, and stays accumulated.

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

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

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

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

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

  As described above, the present specification includes a radiation source, a first grating arranged in a traveling direction of radiation emitted from the radiation source, and radiation that has passed through the first grating. A grating pattern having a period substantially matching a pattern period of the radiation image, a radiation image detector for detecting the radiation image masked by the grating pattern, a cooling unit for air-cooling the radiation source, In the first shooting mode in which shooting is performed by placing a subject upstream of the first grid in the traveling direction or between the first grid and the grid pattern, air cooling by the cooling unit is performed during shooting of the subject. A radiation imaging system including a control unit that stops the operation is disclosed.

  Further, in the radiation imaging system, a radiation imaging system is disclosed in which the lattice pattern is placed at a plurality of relative positions having different phases with respect to the radiation image.

  In the radiation imaging system, the lattice pattern is a second lattice, and one of the first lattice and the second lattice is moved, and the second lattice is moved with respect to the radiation image. A radiation imaging system further comprising a scanning mechanism placed at the plurality of relative positions is disclosed.

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

  In the radiation imaging system, the radiation image detector includes pixels that detect the radiation image masked by the lattice pattern in a two-dimensional array, and the pixel rows are orthogonal to the pixel rows. Scanning is sequentially performed in the column direction, and image signals corresponding to the radiation images of the pixel rows are sequentially read out, and the first grating and the second grating are extending directions of the first grating. And image signals read from groups of different pixel rows based on image signals acquired by the radiological image detector, which are arranged so as to be relatively inclined with respect to the extending direction of the second grating. Radiographic imaging further comprising a phase image generation unit that acquires a phase contrast image based on the acquired image signals of a plurality of fringe images The stem is disclosed.

  The radiation imaging system further includes a linear reading light source extending in a direction in which the pixel row extends, and the radiation image detector is scanned in a direction in which the linear reading light source extends. Discloses a radiation imaging system from which the image signal is read out.

  The radiation imaging system further includes a third grating that selectively transmits the radiation emitted from the radiation source and emits the radiation to the first grating, and the third grating includes the radiation. A radiation imaging system provided at the source is disclosed.

  In the radiation imaging system, a second imaging mode in which imaging is performed by retracting at least the first grating from the radiation field and placing a subject between the radiation source and the radiation image detector. In the above, a radiation imaging system is disclosed in which the control unit continues the air cooling without stopping the air cooling by the cooling unit when the subject is imaged.

  Moreover, the said radiography system is disclosing the radiography system further provided with the vibration detection part which detects the vibration generate | occur | produced in the said radiation source.

  In the radiation imaging system, in the first imaging mode, the control unit permits imaging of an object when the vibration amplitude of vibration detected by the vibration detection unit is a threshold value or less. Is disclosed.

  In the radiation imaging system, the radiation source has an X-ray tube that generates X-rays by colliding an electron beam with a rotating anode rotating at a predetermined speed, and has a temperature rise characteristic according to imaging conditions. Based on this, a radiation imaging system having a temperature prediction unit that predicts the temperature of the X-ray tube when imaging is performed under set imaging conditions is disclosed.

  In the radiation imaging system, a radiation imaging system is disclosed in which the control unit permits imaging of a subject when the temperature predicted by the temperature prediction unit is equal to or lower than a threshold value.

  Further, in the above radiographic system, there is disclosed a radiographic system in which the control unit prohibits imaging of a subject when the temperature predicted by the temperature prediction unit is equal to or higher than a threshold value.

  Further, in the above radiographic system, there is disclosed a radiographic system further including a warning information unit that issues a warning notifying the possibility of failure of the radiation source when the temperature predicted by the temperature prediction unit is equal to or higher than a threshold value. .

  In the radiation imaging system, the imaging condition may be at least one of a value of a tube voltage used for the X-ray tube, a value of a tube current used for the X-ray tube, and an irradiation time for irradiating the subject with X-rays. A radiation imaging system showing one of them is disclosed.

  In the radiation imaging system, a distribution of refraction angles of radiation incident on the radiation image detector is calculated from an image acquired by the radiation image detector, and the phase of the subject is calculated based on the distribution of refraction angles. A radiation imaging system having an arithmetic unit that generates a contrast image is disclosed.

11 X-ray source 20 Controller 27 X-ray tube cooler 28 X-ray tube thermometer 29 X-ray tube vibrometer 30 Flat panel detector (FPD)
31 First absorption type grating 32 Second absorption type grating

Claims (16)

A radiation source;
A first grating disposed in a traveling direction of radiation emitted from the radiation source;
A grating pattern having a period substantially matching a pattern period of a radiation image formed by radiation that has passed through the first grating;
A radiation image detector for detecting the radiation image masked by the lattice pattern;
A cooling section for air-cooling the radiation source;
In the first shooting mode in which shooting is performed by placing a subject upstream of the first grating or between the first grating and the grating pattern in the traveling direction of radiation, the cooling unit The control unit that stops air cooling by
Radiation imaging system provided. The radiation imaging system according to claim 1,
The radiographic system in which the lattice pattern is placed at a plurality of relative positions having different phases from each other with respect to the radiation image. The radiation imaging system according to claim 1,
The lattice pattern is a second lattice;
A radiation imaging system further comprising a scanning mechanism that moves one of the first grating and the second grating and places the second grating at the plurality of relative positions with respect to the radiation image. The radiation imaging system according to claim 1,
The radiation image detector includes, for each pixel, a conversion layer that converts radiation into charges, and a charge collection electrode that collects charges converted in the conversion layer,
The charge collection electrode includes a plurality of linear electrode groups having a period substantially matching the pattern period of the radiation image,
The plurality of linear electrode groups are arranged so that their phases are different from each other,
The grid pattern is a radiation imaging system configured by each of the plurality of linear electrode groups. The radiation imaging system according to claim 1 or 2,
In the radiation image detector, pixels that detect the radiation image masked by the lattice pattern are arranged in a two-dimensional manner, and the pixel rows are sequentially scanned in a pixel column direction orthogonal to the pixel rows. The image signal corresponding to the radiation image for each pixel row is sequentially read out,
The first grating and the second grating are arranged such that the extending direction of the first grating and the extending direction of the second grating are relatively inclined;
Based on image signals acquired by the radiation image detector, image signals read from different groups of pixel rows are acquired as image signals of different stripe images, and the acquired image signals of the plurality of stripe images A radiation imaging system further comprising a phase image generation unit that generates a phase contrast image based on the above. The radiation imaging system according to claim 5,
A linear reading light source extending in the extending direction of the pixel rows;
The radiation imaging system, wherein the radiation image detector reads the image signal by scanning the linear reading light source in a direction in which the pixel row extends. The radiographic system according to any one of claims 1 to 6,
Further comprising a third grating that selectively passes through the radiation emitted from the radiation source and emits the radiation to the first grating;
The third grating is a radiation imaging system provided in the radiation source. The radiographic system according to any one of claims 1 to 7,
In the second photographing mode in which at least the first grating is retracted from the radiation field and the subject is placed between the radiation source and the radiation image detector to perform photographing, The control unit is a radiation imaging system in which air cooling is continued without stopping air cooling by the cooling unit. The radiographic system according to any one of claims 1 to 8,
A radiation imaging system further comprising a vibration detection unit that detects vibration generated in the radiation source. The radiation imaging system according to claim 9,
In the first imaging mode, the control unit permits imaging of a subject when a vibration amplitude of vibration detected by the vibration detection unit is a threshold value or less. It is a radiography system as described in any one of Claim 1 to 10, Comprising:
The radiation source has an X-ray tube that generates X-rays by colliding an electron beam with a rotating anode that rotates at a predetermined speed;
A radiation imaging system including a temperature prediction unit that predicts a temperature of the X-ray tube when imaging is performed under set imaging conditions based on a temperature rise characteristic according to imaging conditions. It is a radiography system of Claim 11, Comprising:
The said control part is a radiography system which permits imaging | photography of a to-be-photographed object, when the temperature estimated by the said temperature estimation part is below a threshold value. The radiation imaging system according to claim 12,
The said control part is a radiography system which prohibits imaging | photography of a to-be-photographed object, when the temperature estimated by the said temperature estimation part is more than a threshold value. The radiation imaging system according to claim 13,
A radiation imaging system further comprising a warning information unit that issues a warning notifying the possibility of failure of the radiation source when the temperature predicted by the temperature prediction unit is equal to or higher than a threshold value. The radiation imaging system according to any one of claims 11 to 14,
The radiographic system showing at least one of a value of a tube voltage used for the X-ray tube, a value of a tube current used for the X-ray tube, and an irradiation time for irradiating the subject with X-rays . The radiation imaging system according to any one of claims 1 to 15,
A calculation unit that calculates a distribution of refraction angles of radiation incident on the radiation image detector from an image acquired by the radiation image detector, and generates a phase contrast image of a subject based on the distribution of refraction angles; ,
A radiation imaging system having

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