WO2012057278A1

WO2012057278A1 - Radiation imaging system and radiation imaging method

Info

Publication number: WO2012057278A1
Application number: PCT/JP2011/074840
Authority: WO
Inventors: 村越　大; 裕康石井
Original assignee: 富士フイルム株式会社
Priority date: 2010-10-27
Filing date: 2011-10-27
Publication date: 2012-05-03
Also published as: JP2014014379A

Abstract

High-precision phase contrast radiological images are generated by appropriate exposure control. A radiation imaging system (10) is provided with: a first grid (31); a second grid (32) having periods that substantially match the pattern periods of a radiological image formed by radiation that has passed through the first grid, and disposed in a plurality of relative positions where the phases differ to each other with respect to the radiological image; a radiation image detector (30) that detects the radiation image masked by the second grid; a dose detection unit (35) that is positioned downstream from the second grid, and that detects the dose of incident radiation; and a control unit (20) that controls exposure. The control unit (20) controls exposure such that, in a plurality of exposures in which the second grid (32) is disposed in the relative positions that differ to each other, exposure continues in the first exposure until the dose of radiation detected by the dose detection unit (35) reaches a preset threshold amount, and such that exposure continues in the second exposure onward until the exposure time required in the first exposure has elapsed.

Description

Radiographic system and radiographic method

The present invention relates to a radiation imaging system and a radiation imaging method.

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

In a general X-ray imaging system, a subject is placed between an X-ray source that emits X-rays and an X-ray image detector that detects an X-ray image, and a transmission image of the subject is captured. In this case, each X-ray radiated from the X-ray source toward the X-ray image detector has characteristics (atomic number, density, thickness) of the substance constituting the subject existing on the path to the X-ray image detector. ), The light is incident on the X-ray image detector. As a result, an X-ray transmission image of the subject is detected and imaged by the X-ray image detector. As an X-ray image detector, there is a flat panel detector (FPD: Flat Panel Detector) using a semiconductor circuit in addition to a combination of an X-ray intensifying screen and a film, a stimulable phosphor (accumulating phosphor), and so on. Widely used. In addition, automatic exposure is used to stabilize the density of the image obtained by the X-ray image detector with respect to the required exposure amount that varies depending on the subject, or to prevent excessive exposure of the subject due to excessive exposure. Control is taking place. In the automatic exposure control, generally, the dose of X-rays transmitted through the subject is detected by a dose detector, and the X-ray irradiation is stopped when the dose detected by the dose detector reaches a preset threshold dose. .

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

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

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

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

Patent Document 1 describes that the above automatic exposure control is performed in the X-ray phase imaging by the fringe scanning method using the first and second diffraction gratings.

International Publication No. 08/102598

In X-ray phase imaging by the fringe scanning method using the first and second diffraction gratings, the radiation image formed by the X-rays that have passed through the first diffraction grating is superimposed on the second diffraction grating as described above. Thus, moire fringes are formed on the image receiving surface of the X-ray image detector downstream of the second diffraction grating. Therefore, although the image detected by the X-ray image detector includes moire fringes, an X-ray phase contrast image from which the influence of moire fringes has been finally removed is presented in medical diagnosis and non-destructive inspection. It is preferable. However, depending on the degree of the grating pitch error of the first and second diffraction gratings, the scanning error of the second diffraction grating, etc., the moire fringes from the X-ray phase contrast image become shorter as the period of the moire fringes becomes shorter. It may be difficult to remove the influence of the image quality, and the image quality may deteriorate. Therefore, it is preferable that the period of the moire fringes is long.

Patent Document 1 does not particularly describe the position of the dose detector, but generally the dose detector is disposed behind the X-ray image detector. In that case, in the X-ray phase imaging based on the fringe scanning method using the first and second diffraction gratings, the dose detector is positioned downstream of the second diffraction grating, and thus on the light receiving portion of the dose detector. In this case, moire fringes are formed. This moire fringe moves with the scanning of the second diffraction grating, and when the period of the moire fringe becomes longer than the dimension of the light receiving part of the dose detector in the periodic direction, when the dark part of the moire fringe overlaps the light receiving part, When the dark part does not overlap the light receiving part, the dose of X-rays incident on the dose detector per unit time varies greatly.

The above automatic exposure control extends or shortens the X-ray irradiation time so as to cancel the fluctuation of the X-ray dose incident on the light receiving unit of the dose detector per unit time. As a result, there is a variation in the irradiation dose between photographings, and the variation in the irradiation dose between photographings causes a change in the signal value of each pixel separately from the scanning of the second diffraction grating. For example, in the pixel of the X-ray image detector in which the dark portion of the moire fringe has the same phase as the light receiving unit of the dose detector, the change in the signal value that should originally be obtained by scanning the second diffraction grating is attenuated, In the pixel of the X-ray image detector in which the dark portion of the moire fringe overlaps with the light receiving portion of the dose detector in an opposite phase, the change in the signal value that should be originally obtained by scanning the second diffraction grating is emphasized.

As described above, the X-ray phase imaging by the fringe scanning method using the first and second diffraction gratings detects the phase information of the subject from the change in the signal value of each pixel accompanying the scanning of the second diffraction grating. Thus, a change in the signal value of each pixel due to a factor different from the scanning of the second diffraction grating reduces the detection accuracy of the phase information of the subject.

The present invention has been made in view of the above-described circumstances, and an object thereof is to generate a highly accurate radiation phase contrast image by appropriate exposure control.

(1) A plurality of first gratings and a plurality of phases having a period substantially coincident with a pattern period of a radiation image formed by radiation passing through the first grating and having phases different from each other with respect to the radiation image A second grating placed at a relative position; a radiation image detector for detecting the radiation image masked by the second grating; and the second grating in a traveling direction of the radiation passing through the second grating. A dose detection unit that detects the amount of incident radiation and a control unit that controls exposure, and the control unit is configured to perform a plurality of times in which the second grating is placed at different relative positions. In the first shooting, the exposure is controlled to continue until the dose detected by the dose detection unit reaches a preset threshold dose, and the second and subsequent shootings are performed in the first shooting. Radiographic system for controlling to continue the exposure until the exposure time required has elapsed.
(2) using the first grating and the second grating having a period substantially matching the pattern period of the radiation image formed by the radiation that has passed through the first grating; Is taken at a relative position with a phase different from each other with respect to the radiographic image, and the radiographic image masked by the second grating is detected for each radiographing. In this imaging, the exposure is continued until the dose detected downstream of the second grating reaches a preset threshold dose in the traveling direction of the radiation passing through the second grating, and the second and subsequent imaging is performed. Then, the radiation imaging method which continues exposure until the exposure time required in the said 1st imaging | photography has passed.

According to the present invention, in the first shooting, exposure control is performed based on the dose detected by the dose detector, and in the second and subsequent shootings, exposure control is performed based on the exposure time required for the first shooting. As a result, the necessary exposure amount that varies depending on the subject is ensured, and variations in the irradiation dose during photographing are prevented. Thereby, a highly accurate radiation phase contrast image can be generated.

It is a schematic diagram which shows the structure of an example of the radiography system for describing embodiment of this invention. It is a control block diagram of the radiography system of FIG. It is a schematic diagram which shows the structure of the radiographic image detector of the radiography system of FIG. It is a perspective view of the imaging part of the radiography system of FIG. It is a side view of the imaging part of the radiography system of FIG. It is a schematic diagram which shows the mechanism for changing the period of the moire fringe by superimposition of the 1st and 2nd grating | lattice. It is a schematic diagram for demonstrating the refraction | bending of the radiation by a to-be-photographed object. It is a schematic diagram for demonstrating the fringe scanning method. It is a graph which shows the signal of the pixel of the radiographic image detector accompanying a fringe scanning. It is a flowchart which shows the imaging method by the radiography system of FIG. It is a schematic diagram which shows the structure of the radiographic image detector regarding the modification of the radiography system of FIG. It is a flowchart which shows the imaging | photography method regarding the modification of the radiography system of FIG. It is a flowchart which shows the imaging | photography method regarding the other modification of the radiography system of FIG. It is a schematic diagram which shows the imaging | photography method regarding the other modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a schematic diagram which shows the structure of the modification of the radiography system of FIG. It is a schematic diagram which shows the structure of the other example of the radiography system for describing embodiment of this invention. It is a 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 of the calculating part of the radiography system of FIG.

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

The X-ray imaging system 10 is an X-ray diagnostic apparatus that images a subject (patient) H in a standing position, and is disposed opposite to the X-ray source 11 that emits X-rays to the subject H, and the X-ray source 11. An imaging unit 12 that detects X-rays transmitted through the subject H from the X-ray source 11 and generates image data, and controls the exposure operation of the X-ray source 11 and the imaging operation of the imaging unit 12 based on the operation of the operator. At the same time, it is roughly divided into a console 13 that generates a phase contrast image by calculating 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.

Based on the control of the X-ray source control unit 17, the X-ray source 11 is emitted from the X-ray tube 18 that generates X-rays according to the high voltage applied from the high voltage generator 16, and the X-ray tube 18. The X-ray includes a collimator unit 19 including a movable collimator 19a that limits an irradiation field so as to shield a portion of the X-ray that does not contribute to imaging of the inspection area of the subject H. The X-ray tube 18 is of an anode rotating type, and emits an electron beam from a filament (not shown) as an electron emission source (cathode) and collides with a rotating anode 18a rotating at a predetermined speed, thereby causing X-rays. Is generated. The colliding portion of the rotating anode 18a with the electron beam becomes the X-ray focal point 18b.

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

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

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

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

As the input device 21, for example, a switch, a touch panel, a mouse, a keyboard, or the like can be used. By the operation of the input device 21, an X-ray tube voltage or an X-ray dose detected by a dose detector described later can be used. X-ray imaging conditions such as a threshold dose, imaging timing, and the like are input. The monitor 24 includes a liquid crystal display or the like, and displays characters such as X-ray imaging conditions and X-ray images under the control of the control device 20.

The imaging unit 12 includes a flat panel detector (FPD) 30 made of a semiconductor circuit, a first absorption type grating 31 and a second absorption type for detecting phase change (angle change) of X-rays by the subject H and performing phase imaging. An absorption grating 32 and a dose detector 35 are provided.

The FPD 30 is arranged so that the detection surface is orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11. Although described in detail later, the first and

second absorption gratings

31 and 32 are disposed between the FPD 30 and the X-ray source 11.

The dose detector 35 is disposed behind the FPD 30 and is located downstream of the subject H, and detects the dose of X-rays transmitted through the subject H. As the X-ray light receiving unit of the dose detector 35, for example, a combination of a phosphor and a photomultiplier tube, an ion chamber, an X-ray detector using a semiconductor circuit, or the like is used.

The imaging unit 12 changes the relative positional relationship of the second absorption type grating 32 with respect to the first absorption type grating 31 by translating the second absorption type grating 32 in the vertical direction (x direction). A scanning mechanism 33 is provided. The scanning mechanism 33 is configured by an actuator such as a piezoelectric element, for example.

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

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

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

Each pixel 40 once converts X-rays into visible light by a scintillator (not shown) made of terbium activated gadolinium oxide (Gd ₂ O ₂ S: Tb), thallium activated cesium iodide (CsI: Tl), or the like. It is also possible to configure as an indirect conversion type X-ray detection element that converts the converted visible light into a charge by a photodiode (not shown) and accumulates it. The X-ray image detector is not limited to an FPD based on a TFT panel, and various X-ray image detectors based on a solid-state imaging device such as a CCD sensor or a CMOS sensor can also be used.

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

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

The first absorption-type grating 31 includes a substrate 31a and a plurality of X-ray shielding portions 31b arranged on the substrate 31a. Similarly, the second absorption type grating 32 includes a substrate 32a and a plurality of X-ray shielding portions 32b arranged on the substrate 32a. The

substrates

31a and 32a are both made of an X-ray transparent member such as glass that transmits X-rays.

Each of the

X-ray shielding portions

31b and 32b is in one direction in a plane orthogonal to the optical axis A of the X-rays emitted from the X-ray source 11 (in the illustrated example, the y direction orthogonal to the x direction and the z direction). It is comprised by the linear member extended | stretched. As a material of each

X-ray shielding part

31b, 32b, a material excellent in X-ray absorption is preferable, and for example, a heavy metal such as gold or platinum is preferable. These

X-ray shielding portions

31b and 32b can be formed by a metal plating method or a vapor deposition method.

X-ray shielding portion 31b is in a plane perpendicular to the optical axis A of the X-ray, at a predetermined period p ₁ in a direction (x-direction) orthogonal to the one direction, are arranged at a predetermined interval d ₁ from each other ing. Similarly, X-ray shielding portion 32b, in the plane orthogonal to the optical axis A of the X-ray, at a predetermined period p ₂ in a direction (x-direction) orthogonal to the one direction, at a predetermined interval d ₂ from each other Are arranged. Since the first and

second absorption gratings

31 and 32 do not give a phase difference to incident X-rays but give an intensity difference, they are also called amplitude gratings. Note that the slit portions (regions having the distances d ₁ and d ₂ ) may not be voids, and the voids may be filled with an X-ray low-absorbing material such as a polymer or a light metal.

The first and

second absorption 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 effective wavelength of the X-rays emitted from the X-ray source 11, most of the X-rays included in the irradiated X-rays can be obtained at the slit portion. It is configured to pass through without being diffracted while maintaining straightness. For example, when tungsten is used as the rotary anode 18a described above and the tube voltage is 50 kV, the effective wavelength of X-ray is about 0.4 mm. In this case, if the distances d ₁ and d ₂ are about 1 to 10 μm, most of the X-rays are geometrically projected without being diffracted at the slit portion.

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

In the Talbot interferometer, the distance L ₂ from the first absorption type grating 31 to the second absorption type grating 32 is limited to the Talbot interference distance determined by the grating pitch of the first diffraction grating and the X-ray wavelength. The imaging unit 12 of the present X-ray imaging system 10 has a configuration in which the first absorption grating 31 projects incident X-rays without diffracting, and the G1 image of the first absorption grating 31 is the first. because at every position of the rear absorption type grating 31 similarly obtained, the distance L _2, can be set independently of the Talbot distance.

As described above, the imaging unit 12 does not constitute a Talbot interferometer, but the Talbot interference distance Z when it is assumed that X-rays are diffracted by the first absorption type grating 31 is the first absorption type grating. the grating pitch _{p 1} of 31, the grating pitch _p 2, X-ray wavelength of the second absorption-type grating 32 (effective wavelength) lambda, and using the positive integer m, is expressed by the following equation (2).

Expression (2) is an expression that represents the Talbot interference distance when the X-ray irradiated from the X-ray source 11 is a cone beam. “Atsushi Momose, et al., Japan Journal of Applied Physics, Vol. 47, No. 10, October 2008, page 8077 ”.

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

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

The

X-ray shielding portions

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

However, when the X-rays irradiated from the X-ray source 11 are cone beams, if the thicknesses h ₁ and h ₂ of the

X-ray shielding portions

31b and 32b are too thick, the X-rays incident obliquely enter the slit portion. There is a problem that it becomes difficult to pass, so-called vignetting occurs, and the effective visual field in the direction (x direction) perpendicular to the extending direction (strand direction) of the

X-ray shielding portions

31b and 32b becomes narrow. Therefore, in view of the field of view secured to define the upper limit of the thickness _h 1, _{h 2.} In order to secure the effective field length V in the x direction on the detection surface of the FPD 30, assuming that the distance from the X-ray focal point 18 b to the detection surface of the FPD 30 is L, the thicknesses h ₁ and h ₂ are shown in FIG. From the scientific relationship, it is necessary to set so as to satisfy the following expressions (6) and (7).

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

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

second absorption gratings

31 and 32 and the distance between the two changes. I mean.

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

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

Further, it is possible to detect moire fringes even if the arrangement pitch P is larger than the moire period T within the range satisfying the expression (9), but the arrangement pitch P is preferably smaller than the moire period T. 10) is preferably satisfied. This is because, in order to obtain a high-quality phase contrast image, moire fringes are preferably detected with high contrast in the phase contrast image generation process described later.

Since the arrangement pitch P of the pixels 40 of the FPD 30 is a value determined by design (generally about 100 μm) and is difficult to change, the magnitude relationship between the arrangement pitch P and the moire period T is adjusted. Adjusts the positions of the first and

second absorption gratings

31 and 32 and changes the moire period T by changing at least one of the pattern period p ₁ ′ and the grating pitch p ₂ ′ of the G1 image. It is preferable to do.

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

The moire period T can be changed by relatively rotating one of the first and

second absorption gratings

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

As another example, the change of the moire period T is such that either one of the first and second

absorption type gratings

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

As another example, the moire period T can be changed by relatively moving one of the first and

second absorption gratings

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

In the X-ray imaging system 10, imaging unit 12 is not the Talbot interferometer as described above, since the distance L ₂ can be freely set, moire by changing the distance L ₂ as relative movement mechanism 52 A mechanism for changing the period T can be suitably employed. The change mechanism (relative rotation mechanism 50, relative tilt mechanism 51, and relative movement mechanism 52) of the first and

second absorption gratings

31 and 32 for changing the moiré period T is constituted by an actuator such as a piezoelectric element. Is possible.

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

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

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

Reference numeral 55 indicates an X-ray path that travels straight when the subject H is not present. The X-ray that travels along the path 55 passes through the first and

second absorption gratings

31 and 32 and enters the FPD 30. To do. Reference numeral 56 indicates an X-ray path refracted and deflected by the subject H when the subject H exists. X-rays traveling along this path 56 are shielded by the second absorption type grating 32 after passing through the first absorption type grating 31.

The phase shift distribution Φ (x) of the subject H is expressed by the following equation (11), where n (x, z) is the refractive index distribution of the subject H, and z is the direction in which the X-ray travels.

The G1 image projected from the first absorptive grating 31 to the position of the second absorptive grating 32 is displaced in the x direction by an amount corresponding to the refraction angle φ due to refraction of X-rays at the subject H. become. This amount of displacement Δx is approximately expressed by the following equation (12) based on the small X-ray refraction angle φ.

Here, the refraction angle φ is expressed by Expression (13) using the X-ray wavelength λ and the phase shift distribution Φ (x) of the subject H.

Thus, the displacement amount Δx of the G1 image due to the refraction of X-rays at the subject H is related to the phase shift distribution Φ (x) of the subject H. The amount of displacement Δx is expressed by the following equation with the phase shift amount ψ of the signal output from each pixel 40 of the FPD 30 (the phase shift amount of the signal of each pixel 40 with and without the subject H): It is related as shown in (14).

Therefore, by obtaining the phase shift amount ψ of the signal of each pixel 40, the refraction angle φ is obtained from the equation (14), and the differential amount of the phase shift distribution Φ (x) is obtained using the equation (13). Is integrated with respect to x, a phase shift distribution Φ (x) of the subject H, that is, a phase contrast image of the subject H can be generated. In the present X-ray imaging system 10, the phase shift amount ψ is calculated using a fringe scanning method described below.

In the fringe scanning method, imaging is performed while one of the first and second

absorption type gratings

31 and 32 is translated in a stepwise manner relative to the other in the x direction (that is, the phase of both grating periods is changed). Shoot while changing). In the X-ray imaging system 10, the second absorption type grating 32 is moved by the scanning mechanism 33 described above, but the first absorption type grating 31 may be moved. As the second absorption type grating 32 moves, the moire fringes move, and the translation distance (the amount of movement in the x direction) is one period of the grating period of the second absorption type grating 32 (grating pitch p ₂ ). (Ie, when the phase change reaches 2π), the moire fringes return to their original positions. Such a change in moire fringes is obtained by photographing the moire fringes with the FPD 30 while moving the second absorption grating 32 by an integer of the grating pitch p _2, and from each of the photographed plural fringe images, The signal is acquired and processed by the processing unit 22 to obtain the phase shift amount ψ of the signal of each pixel 40.

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

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

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

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

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

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

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

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

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

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

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

In the above calculation, the change of the M signal values of each pixel 40 for calculating the phase shift amount ψ needs to be brought about by the scanning of the second absorption type grating 32. For this purpose, the X-ray irradiation dose irradiated from the X-ray source 11 to the imaging unit 12 is required to be substantially constant during imaging.

FIG. 10 shows an imaging flow by the X-ray imaging system 10.

In the X-ray imaging system 10, in the first imaging in which the second absorption grating 32 is placed at the position of k = 0 (see FIG. 8), exposure control is performed based on the dose detected by the dose detector 35. Is done. In the second and subsequent photographing in which the second absorption type grating 32 is placed at each position of K = 1, 2,..., M−1 (see FIG. 8), the exposure time required for the first photographing is obtained. Based on this, exposure control is performed.

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

The dose detector 35 detects the cumulative X-ray dose incident on the light receiving unit from the start of X-ray irradiation. For the dose detected by the dose detector 35, for example, a threshold dose is set in consideration of a necessary exposure amount that differs depending on the subject H. When the detected X-ray dose reaches the threshold dose, the dose detector 35 sends a control signal indicating that the threshold dose has been reached to the control device 20 (step S2).

The control device 20 that has received the control signal sent from the dose detector 35 sends a control signal for instructing to stop the X-ray irradiation to the X-ray source control unit 17. The X-ray source control unit 17 that has received this control signal controls the high voltage generator 16 so as to stop the supply of power to the X-ray tube 18. Thereby, irradiation of the subject H with X-rays is stopped (step S3).

The control device 20 sends an exposure time T ₀ required for the first imaging, that is, a control signal that instructs the X-ray source controller 17 to stop irradiation after sending a control signal that instructs the X-ray source controller 17 to start X-ray irradiation. Is measured and stored.

Next, in the second and subsequent imaging, the control device 20 sends a control signal instructing the start of X-ray irradiation to the X-ray source control unit 17. The X-ray source control unit 17 that has received this control signal controls the high voltage generator 16 so as to start supplying power to the X-ray tube 18. Thereby, irradiation of the subject H with X-rays is started (step S4).

Here, in the second and subsequent imaging, the moire fringes move with the scanning of the second absorption type grating 32, and the dose depends on the degree of overlap between the light receiving part of the dose detector 35 and the dark part of the moire fringes. The dose incident on the light receiving portion of the detector 35 per unit time changes. Therefore, the time required for the dose detected by the dose detector 35 to reach the above threshold dose in each of the second and subsequent imaging is the dose detected by the dose detector 35 in the first imaging. It differs from the exposure time T ₀ it took to reach the dose. Therefore, in each of the second and subsequent shootings, when exposure control is performed based on the dose detected by the dose detector 35, the exposure time varies between the shootings, and as a result, the irradiation dose varies. It will be.

Therefore, the control device 20 measures the elapsed time T after sending a control signal instructing the X-ray source control unit 17 to start X-ray irradiation, and the elapsed time T is stored in the first time. When the exposure time T ₀ required for imaging is reached, a control signal for instructing to stop X-ray irradiation is sent to the X-ray source control unit 17 (step S5). The X-ray source control unit 17 that has received this control signal controls the high voltage generator 16 so as to stop the supply of power to the X-ray tube 18. Thereby, irradiation of the subject H with X-rays is stopped (step S6).

Thus, in the first imaging, the exposure control is performed based on the dose detected by the dose detector 35, in the second and subsequent shot, the exposure based on the exposure time T ₀ required for the first shooting By performing the control, a necessary exposure amount that is different depending on the subject H is secured, and variations in irradiation dose between photographings are prevented. Thereby, a highly accurate radiation phase contrast image can be generated.

Here, when the dark part of the moire fringes overlaps the light receiving part of the dose detector 35 and when the dark part of the moire fringes does not overlap the light receiving part, the dose incident on the light receiving part per unit time is the latter In many cases, the time required to reach the threshold dose is shorter in the latter case. In the first shooting, not overlapping dark moire fringes on the light receiving portion of the dose detector 35, when the exposure time T ₀ in this state is determined, including the second and subsequent shot, the dark portion of the moire fringes There is a possibility that the dose of X-rays incident on the pixel 40 of the overlapping FPD 30 is insufficient, and the S / N of the signal of the pixel is lowered.

Therefore, in the first imaging, the subject H is allowed within a range in which the pixels 40 of the FPD 30 are not saturated under the assumption that the dark part of the moire fringes does not overlap the light receiving part of the dose detector 35. It is preferable to set the threshold dose slightly higher within the range of the exposure dose. Further, it is preferable that the readout circuit 43 is configured to be designed or set so that the inputs of the integration amplifier circuit, the A / D converter and the like are not saturated with respect to the saturation output of the pixel 40 of the FPD 30. As a result, the X-ray dose incident on the pixel 40 is ensured even in the pixel 40 where the dark part of the moire fringe overlaps, including the second and subsequent imaging, and the S / N of the signal of the pixel is prevented from decreasing. it can.

In addition, although demonstrated as what switches irradiation and a stop of X-ray | X_line by the start and stop of supply of the electric power to the X-ray tube 18, X-ray | X_line 18 is supplied by the opening and closing of the collimator 19a, and X-ray | X_line 18 is supplied. The irradiation and stop of the line may be switched, or a disk-shaped (or slit-shaped) shutter plate in which openings and shields are alternately formed is provided at the exit of the X-ray source 11, and this is irradiated with X-rays. X-ray irradiation and stop may be switched by rotating (or translating) so as to be synchronized with the timing. According to this, it is possible to keep the X-ray tube 18 in a stable state and more reliably prevent variations in irradiation dose.

The above-described fringe scanning and phase contrast image generation processing is automatically performed after the imaging instruction is given by the operator from the input device 21, and the respective units are linked and operated based on the control of the control device 20. The phase contrast image of the subject H is displayed on the monitor 24.

As described above, according to the present X-ray imaging system 10, exposure control is performed based on the dose detected by the dose detector 35 in the first imaging, and the first imaging in the second and subsequent imaging. By performing the exposure control based on the exposure time T ₀ required for the imaging, a necessary exposure amount that varies depending on the subject H is ensured, and variations in the irradiation dose between the imaging are prevented. Thereby, a highly accurate radiation phase contrast image can be generated.

Further, according to the present X-ray imaging system 10, since most of the X-rays are linearly projected onto the second absorption grating 32 without being diffracted by the first absorption grating 31, High spatial coherence is not required, and a general X-ray source used in the medical field as the X-ray source 11 can be used. The distance L ₂ from the first absorption type grating 31 to the second absorption type grating 32 can be set to an arbitrary value, and the distance L ₂ is smaller than the minimum Talbot interference distance in the Talbot interferometer. Since it can be set, the photographing unit 12 can be downsized (thinned). 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. .72, No. 1 (1982) p. 156 ", and 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 present X-ray imaging system 10, the dose detector 35 is disposed behind the FPD 30, but the dose detector 35 may be disposed between the second absorption grating 32 and the FPD 30. Furthermore, the X-ray image detector itself may have a configuration equivalent to that of the dose detector 35. As a specific mode, an X-ray image detector having a configuration disclosed in Japanese Patent Application Laid-Open No. 2004-130058 is disclosed. Can be used, thereby eliminating the dose detector 35. As shown in FIG. 11, the X-ray image detector includes a pixel (photoelectric conversion element) 47 for detecting a dose, in addition to a plurality of pixels 40 for capturing moire fringes. The pixel 47 is connected to a readout circuit 48 different from the group of the pixels 40 and is configured to always output charges according to the amount of incident light without switching. The charge read from the pixel 47 is amplified by the read circuit 48, and the output is added by an adder (not shown), so that it enters the pixel 47 from the start of X-ray irradiation in the first imaging. A cumulative x-ray dose is detected.

FIG. 12 shows an imaging flow in a modification of the radiation imaging system 10.

In this modification, pre-photographing is performed a plurality of times with a constant exposure time in a state where the subject H is not arranged. The initial position of the second absorption type grating 32 in the plurality of times of photographing performed with the subject H being arranged based on the change in the dose detected by the dose detector 35 during the plurality of times of pre-imaging. Is set.

First, in a state where the subject H is not arranged, the exposure time is constant, and the second absorption grating 32 is placed at each position of k = 0, 1, 2,..., M−1 (see FIG. 8). Thus, pre-shooting is performed a plurality of times (steps SS1 to SS6).

The control device 22 stores the dose detected by the dose detector 35 in each pre-imaging, and determines the pre-imaging in which the minimum dose is detected. Then, the position (k _min ) of the second absorption grating 32 in the pre-imaging in which the minimum dose is detected is set to the position of the second absorption grating 32 in the plurality of imaging performed in the state where the subject H is arranged. The initial position is set (step SS7). In pre-photographing in which the minimum dose is detected among a plurality of pre-photographings with a constant exposure time, it can be said that the light receiving part of the dose detector 35 overlaps with the dark part of the moire fringes or the overlap is large.

Next, the subject H is placed, the second absorption grating 32 is placed at the initial position ( _kmin ), and exposure is continued until the dose detected by the dose detector 35 reaches a preset threshold dose. Thus, the first shooting is performed (steps S1 to S3). Then, while moving the second absorption type grating 32 by a predetermined scanning pitch (p ₂ / M), the exposure is continued until the exposure time required for the first imaging elapses, and the second and subsequent imaging is performed. (Steps S4 to S6).

According to this modification, in the first imaging, the dose is detected in the state where the dark part of the moire fringes overlaps the light receiving unit of the dose detection unit 35. Therefore, including the second and subsequent imaging, Even in the pixel 40 of the FPD 30 in which the dark portion overlaps, the dose of X-rays incident on the pixel is ensured, and the S / N of the signal of the pixel can be prevented from decreasing. In the first imaging, more appropriate exposure control is possible as compared with the case where the threshold dose is always set higher under the assumption that the dark part of the moire fringes does not overlap the light receiving part of the dose detector 35. It becomes.

Note that the initial position of the second absorption grating 32 in a plurality of times of imaging performed with the subject H being arranged is preferably the first position in the pre-imaging in which the minimum dose is detected among the plurality of times of pre-imaging. However, the present invention is not limited to this. For example, the position of the second absorption type grating 32 in the pre-photographing in which a substantially intermediate dose is detected among the changes in the dose detected in a plurality of pre-photographing can be set as the initial position. Further, it is possible to reduce a decrease in S / N of the signal of the pixel 40 where the dark portions of the moire fringes overlap.

FIG. 13 shows an imaging flow in another modification of the radiation imaging system 10.

In this modification, pre-photographing is performed a plurality of times with a constant exposure time in a state where the subject H is not arranged. Then, the threshold dose with respect to the dose detected by the dose detector 35 is corrected based on the change in the dose detected by the dose detector 35 during a plurality of pre-imaging operations.

The control device 22 stores the dose detected by the dose detector 35 in each pre-imaging, obtains the difference Δ between the minimum dose and the dose detected in the first pre-imaging, and the difference Δ is large. The threshold dose is corrected so that the threshold dose relative to the dose detected by the dose detector 35 increases (step SS7). It can be said that the larger the difference Δ is, the larger the overlap between the light receiving part of the dose detector 35 and the bright part of the moire fringes in the first pre-imaging. The correction amount of the threshold dose according to the difference Δ is appropriately determined within a range in which the pixel 40 of the FPD 30 is not saturated.

Next, the subject H is placed, the second absorption grating 32 is placed at the initial position (k = 0), and exposure is performed until the dose detected by the dose detector 35 reaches a preset threshold dose. The first shooting is continuously performed (steps S1 to S3). Then, while moving the second absorption type grating 32 by a predetermined scanning pitch (p ₂ / M), the exposure is continued until the exposure time required for the first imaging elapses, and the second and subsequent imaging is performed. (Steps S4 to S6).

According to this modification, the threshold dose is corrected according to the degree of overlap between the light receiving unit of the dose detection unit 35 and the dark part of the moire fringes in the first imaging, and therefore includes the second and subsequent imaging. Even in the pixel 40 of the FPD 30 where dark portions of moire fringes overlap, the dose of X-rays incident on the pixel is ensured, and the S / N of the signal of the pixel can be prevented from decreasing. In the first imaging, more appropriate exposure control is possible as compared to the case where the threshold dose is always set higher under the assumption that the dark part of the moire fringes does not overlap the light receiving part of the dose detector 35. It becomes.

FIG. 14 shows an imaging method in another modification of the X-ray imaging system 10.

In all of the above-described modifications, dose detection is performed according to the degree of overlap between the light receiving portion of the dose detector 35 and the dark portion of the moire fringe in the first imaging, that is, the relative positional relationship between the dose detector 35 and the moire fringe. The X-ray dose incident on the light receiving unit of the detector 35 per unit time is changed, but the change of the X-ray dose incident on the light receiving unit of the dose detector 35 is suppressed. Also good.

In this modification, the period T of moire fringes is made smaller than the dimension W of the light receiving part of the dose detector 35 in the period direction. The period T of moiré fringes is changed by, for example, the relative rotation mechanism 50, the relative tilt mechanism 51, or the relative movement mechanism 52 described above. In this case, a plurality of bright parts and dark parts of moire fringes overlap the light receiving part of the dose detector 35, and these bright parts and dark parts are averaged to detect a dose. Therefore, the X-ray dose incident on the dose detector 35 per unit time is substantially constant regardless of the relative positional relationship between the dose detector 35 and the moire fringes. Accordingly, it is possible to determine an appropriate exposure time based on the dose detected by the dose detector 35 without performing multiple pre-photographing, and that the dose is insufficient in the pixels 40 where the dark portions of the moire fringes overlap. Can be avoided.

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

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

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

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

FIG. 16 shows a modification of the radiation imaging system of FIG.

16 differs from the above-described mammography apparatus 80 in that the first absorption grating 31 is disposed between the X-ray source 11 and the compression plate 84. The mammography apparatus 90 illustrated in FIG. The first absorption type lattice 31 is accommodated in a lattice accommodation portion 91 connected to the arm member 81. The imaging unit 92 includes an FPD 30, a second absorption grating 32, a scanning mechanism 33, and a dose detector 35. The dose detector 35 is disposed behind the FPD 30.

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

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

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

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

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

The multi-slit 103 is an absorption type grating (third absorption type grating) having a configuration similar to that of the first and second

absorption type gratings

31 and 32 provided in the imaging unit 12, and is in one direction (y direction). The extended X-ray shielding portions are periodically arranged in the same direction (x direction) as the

X-ray shielding portions

31b and 32b of the first and

second absorption gratings

31 and 32. The multi-slit 103 is intended to form a large number of small-focus light sources (dispersed light sources) arranged at a predetermined pitch in the x direction by partially shielding the radiation emitted from the X-ray focal point 18b. .

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

Expression (18) indicates that the projection image (G1 image) of the X-rays emitted from the small focus light sources dispersedly formed by the multi-slit 103 by the first absorption type grating 31 is the second absorption type grating 32. This is a geometric condition for matching (overlapping) in position.

In addition, since the position of the multi slit 103 is substantially the X-ray focal position, the grating pitch p2 of the _second absorption grating 32 is determined so as to satisfy the relationship of the following equation (19).

As described above, in the present X-ray imaging system 100, the G1 images based on the plurality of small focus light sources formed by the multi slit 103 are superimposed, thereby improving the image quality of the phase contrast image without decreasing the X-ray intensity. Can be improved. The multi slit 103 described above can be applied to any of the X-ray imaging systems described above.

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

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

Therefore, this X-ray imaging system uses an arithmetic processing unit 190 that can generate an absorption image and a small-angle scattered image from a plurality of images acquired for a phase contrast image. 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 I _k ( After fitting x, y) with a sine wave, an average value of the fitted sine wave may be obtained. The generation of the absorption image is not limited to the average value, and an addition value obtained by adding the pixel data I _k (x, y) with respect to k can be used as long as the amount corresponds to the average value.

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

According to the present X-ray imaging system, an absorption image and a small angle scattered image are generated from a plurality of images acquired for the phase contrast image of the subject. There is no deviation, and it is possible to superimpose the phase contrast image with the absorption image and the small-angle scattered image, and the burden on the subject is reduced as compared with the case of separately shooting for the absorption image and the small-angle scattered image. be able to.

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

As described above, the present specification includes a period substantially matching the pattern period of the radiation image formed by the first grating and the radiation that has passed through the first grating, and the radiation. A second grating placed at a plurality of relative positions different from each other in phase with respect to the image, a radiation image detector for detecting the radiation image masked by the second grating, and the second grating. A dose detection unit that is positioned downstream of the second grating in the radiation traveling direction and detects an incident radiation dose; and a control unit that controls exposure, wherein the control unit includes the second grating In a plurality of photographings placed at the relative positions different from each other, in the first photographing, the exposure is controlled to continue until the dose detected by the dose detection unit reaches a preset threshold dose, and the second and subsequent times. of The shadow radiographic system for controlling to continue the exposure until the exposure time has elapsed taken in the first imaging is disclosed.

Further, the radiation imaging system disclosed in the present specification is configured such that the control unit performs the pre-imaging a plurality of times when the second grating is placed at different relative positions with a constant exposure time. A relative position of the second grating in the first imaging is determined based on a change in dose detected by the detection unit.

Further, the radiographic system disclosed in the present specification is configured so that the control unit detects a minimum dose among changes detected by the dose detection unit when the pre-imaging is performed a plurality of times. The relative position of the second grating in photographing is set as the relative position of the second grating in the first photographing.

In addition, the radiation imaging system disclosed in the present specification is configured such that the control unit performs the pre-imaging a plurality of times when the second grating is placed at different relative positions with a constant exposure time. The threshold dose is corrected based on a change in dose detected by the detector.

Further, in the radiographic system disclosed in this specification, the dose detected by the dose detection unit in the first pre-imaging when the plurality of pre-imaging is performed is smaller than the minimum dose in the change. The larger the value is, the larger the control unit increases the threshold dose.

Also, in the radiation imaging system disclosed in this specification, the pattern period of the radiation image masked by the second grating is shorter than the dimension of the light receiving unit of the dose detection unit with respect to the periodic direction.

Further, in the radiographic system disclosed in this specification, at least one of a relative posture and a relative position of the second grating with respect to the first grating is changed, and the masking is performed by the second grating. A change mechanism for changing the pattern period of the radiation image is further provided.

Further, in the radiation imaging system disclosed in this specification, the changing mechanism has at least one of the first grating and the second grating around an optical axis of radiation irradiated on the first grating. Rotate one.

Further, in the radiation imaging system disclosed in this specification, the changing mechanism is configured so that at least one of the first grating and the second grating with respect to an optical axis of radiation irradiated on the first grating. Tilt one of them.

Further, in the radiation imaging system disclosed in this specification, the changing mechanism includes at least one of the first grating and the second grating along an optical axis of radiation irradiated on the first grating. Move one.

Further, in the radiation imaging system disclosed in this specification, the dose detection unit is disposed between the second grating and the radiation image detector.

In the radiographic system disclosed in this specification, the dose detection unit is provided in the radiographic image detector.

Further, in the radiographic system disclosed in this specification, the dose detection unit is disposed behind the radiographic image detector.

The radiation imaging system disclosed in this specification calculates a distribution of refraction angles of radiation incident on the radiation image detector from a plurality of radiation images acquired by the radiation image detector. And a calculation unit for generating a phase contrast image of the subject based on the distribution of the subject.

The radiation imaging system disclosed in this specification further includes a radiation source that irradiates radiation toward the first grating.

Further, the present specification uses a first grating and a second grating having a period that substantially matches a pattern period of a radiation image formed by radiation that has passed through the first grating. A radiation imaging method in which the second grating is placed at relative positions different from each other in phase with respect to the radiation image, and the radiation image masked by the second grating is detected for each imaging. In the first imaging, the exposure is continued until the dose detected downstream of the second grating in the traveling direction of the radiation passing through the second grating reaches a preset threshold dose, In the second and subsequent imaging, a radiation imaging method is disclosed in which exposure is continued until the exposure time required for the first imaging has elapsed.

Further, the radiographic method disclosed in the present specification performs a plurality of pre-photographs by placing the second grating at relative positions different from each other with respect to the radiographic image with a constant exposure time. The relative position of the second grating in the first imaging is determined based on the change in the dose detected downstream of the second grating.

Further, the radiography method disclosed in the present specification provides the second radiography in the pre-imaging in which a minimum dose is detected among changes in the dose detected when the pre-imaging is performed a plurality of times. Let the relative position of the grating be the relative position of the second grating in the first imaging.

Further, the radiographic method disclosed in this specification performs a plurality of pre-photographing operations by placing the second grating at relative positions different from each other in phase with respect to the radiographic image with a constant exposure time. The threshold dose is corrected based on a change in dose detected downstream of the second grating.

Further, in the radiographic method disclosed in the present specification, as the dose detected in the first pre-imaging at the time of performing the plurality of pre-imaging is larger than the minimum dose in the change, Increase the threshold dose.

Further, in the radiographic method disclosed in the present specification, the pattern period of the radiation image masked by the second grating is detected by a light receiving unit of a dose detector that detects a dose downstream of the second grating. The imaging is performed a plurality of times with a dimension shorter than the dimension in the periodic direction.

Further, in the radiographic method disclosed in this specification, at least one of a relative posture and a relative position of the second grating with respect to the first grating is changed, and the masking by the second grating is performed. Change the pattern period of the radiation image.

Although the present invention has been described in detail and with reference to specific embodiments, it will be apparent to those skilled in the art that various changes and modifications can be made without departing from the spirit and scope of the invention.
This application is based on a Japanese patent application (Japanese Patent Application No. 2010-241098) filed on Oct. 27, 2010, the contents of which are incorporated herein by reference.

DESCRIPTION OF SYMBOLS 10 X-ray imaging system 11 X-ray source 12 Imaging part 13 Console 20 Control apparatus 30 FPD
31 First Absorption Type Grating 32 Second Absorption Type Grating 33 Scanning Mechanism 35 Dose Detector 40 Pixel

Claims

A first lattice;
A second grating having a period substantially coincident with a pattern period of a radiation image formed by radiation passing through the first grating and being placed at a plurality of relative positions different from each other in phase with respect to the radiation image; When,
A radiation image detector for detecting the radiation image masked by the second grating;
A dose detection unit that is located downstream of the second grating in the traveling direction of the radiation passing through the second grating and detects the amount of incident radiation;
A control unit for controlling exposure;
With
In the plurality of imaging operations in which the second grating is placed at the relative positions different from each other, the control unit is configured to perform a first imaging until the dose detected by the dose detection unit reaches a preset threshold dose. A radiation imaging system that controls to continue exposure and controls to continue exposure until an exposure time required for the first imaging has elapsed in the second and subsequent imaging.
The radiation imaging system according to claim 1,
The control unit, based on a change in dose detected by the dose detection unit when performing a plurality of pre-imaging in which the second grating is placed at the relative position different from each other with a constant exposure time, A radiation imaging system for determining a relative position of the second grating in the first imaging.
The radiographic system according to claim 2,
The control unit is configured to determine a relative position of the second grating in the pre-imaging in which a minimum dose is detected among changes detected by the dose detection unit when the plurality of pre-imaging is performed. A radiation imaging system having a relative position of the second grating in the first imaging.
The radiation imaging system according to claim 1,
The control unit, based on a change in dose detected by the dose detection unit when performing a plurality of pre-imaging in which the second grating is placed at the relative position different from each other with a constant exposure time, A radiography system that corrects the threshold dose.
The radiation imaging system according to claim 4,
The control unit increases the threshold dose as the dose detected by the dose detection unit in the first pre-shooting when the plurality of pre-shoots is performed is larger than the minimum dose in the change. Radiography system.
The radiation imaging system according to claim 1,
A radiation imaging system in which a pattern period of the radiation image masked by the second grating is shorter than a dimension of a light receiving unit of the dose detection unit with respect to the periodic direction.
The radiographic system according to claim 6,
Radiation further comprising a changing mechanism for changing a pattern period of the radiation image masked by the second grating by changing at least one of a relative posture and a relative position of the second grating with respect to the first grating. Shooting system.
The radiation imaging system according to claim 7,
The change mechanism is a radiation imaging system in which at least one of the first grating and the second grating is rotated around an optical axis of radiation applied to the first grating.
The radiation imaging system according to claim 7,
The change mechanism is a radiation imaging system in which at least one of the first grating and the second grating is inclined with respect to an optical axis of radiation applied to the first grating.
The radiation imaging system according to claim 7,
The change mechanism is a radiation imaging system that moves at least one of the first grating and the second grating along an optical axis of radiation applied to the first grating.
The radiation imaging system according to claim 1,
The dose detection unit is a radiation imaging system arranged between the second grating and the radiation image detector.
The radiation imaging system according to claim 1,
The dose detection unit is a radiation imaging system provided in the radiation image detector.
The radiation imaging system according to claim 1,
The dose detection unit is a radiation imaging system arranged behind the radiation image detector.
The radiographic system according to any one of claims 1 to 13,
A distribution of refraction angles of radiation incident on the radiation image detector is calculated from a plurality of radiation images acquired by the radiation image detector, and a phase contrast image of a subject is generated based on the distribution of refraction angles. A radiation imaging system further comprising a calculation unit.
The radiation imaging system according to any one of claims 1 to 14, further comprising a radiation source that irradiates radiation toward the first grating.
A first grating and a second grating having a period that substantially matches a pattern period of a radiation image formed by radiation that has passed through the first grating, and the second grating is the radiation. A radiographic method for performing imaging a plurality of times at relative positions different from each other in phase with respect to an image, and detecting the radiographic image masked by the second grating for each imaging,
In the first imaging, the exposure is continued until the dose detected downstream of the second grating in the traveling direction of the radiation passing through the second grating reaches a preset threshold dose,
In the second and subsequent imaging, the radiation imaging method in which exposure is continued until the exposure time required in the first imaging has elapsed.
A radiography method according to claim 16, comprising:
The exposure time is constant, the second grating is placed at relative positions different from each other in phase with respect to the radiation image, and a plurality of pre-images are taken, and the dose detected downstream of the second grating at that time A radiation imaging method for determining a relative position of the second grating in the first imaging based on a change.
The radiography method according to claim 17, comprising:
The relative position of the second grating in the pre-imaging in which the minimum dose is detected among the changes in the dose detected when the plurality of pre-imaging is performed is the first position in the first imaging. A radiography method in which the relative position of two grids is set.
A radiography method according to claim 16, comprising:
The exposure time is constant, the second grating is placed at relative positions different from each other in phase with respect to the radiation image, and a plurality of pre-images are taken, and the dose detected downstream of the second grating at that time A radiography method for correcting the threshold dose based on a change.
The radiographic method according to claim 19, comprising:
A radiation imaging method in which the threshold dose is increased as the dose detected in the first pre-imaging when the plurality of pre-imaging is performed is larger than a minimum dose in the change.
A radiography method according to claim 16, comprising:
The patterning period of the radiation image masked by the second grating is made shorter than the dimension in the period direction of the light receiving unit of the dose detector that detects the dose downstream of the second grating, and the plurality of times of imaging. Radiography method to do.
The radiation imaging method according to claim 21,
A radiation imaging method for changing a pattern period of the radiation image masked by the second grating by changing at least one of a relative posture and a relative position of the second grating with respect to the first grating.