JP2014138626A - Radiographic apparatus and method for controlling the same - Google Patents

Abstract

PROBLEM TO BE SOLVED: To make it possible to prevent deterioration of image quality of a subject part caused by change of the Talbot distance.SOLUTION: A system control unit 18 comprises a first energy spectrum determination unit, a subject transmissivity determination unit, a second energy spectrum determination unit, and a Talbot distance calculation unit. The first energy spectrum determining unit determines a first energy spectrum of an X-ray prior to incidence on a subject H on the basis of a preset value for the tube voltage of an X-ray source 11. The subject transmissivity determination unit determines the transmissivity to X-rays of the subject H on the basis of the type of the subject H. The second energy spectrum determination unit determines a second energy spectrum by multiplying the first energy spectrum by the transmissivity. The Talbot distance calculation unit determines the center wavelength corresponding to the center energy of the second energy spectrum to calculate the Talbot distance. The system control unit 18 controls first and second position adjustment units 24, 25, and sets the distance between first and second lattices 21, 22 to the Talbot distance.

Description

  The present invention relates to a radiation imaging apparatus that performs phase imaging using a grating and a control method thereof.

  Radiation, such as X-rays, has a characteristic that it attenuates depending on the atomic number of the elements constituting the substance, and the density and thickness of the substance, so it is used as a probe for seeing through the inside of the subject. Yes. X-ray imaging is widely used in fields such as medical diagnosis and non-destructive inspection.

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

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

  In recent years, the development of X-ray phase imaging that obtains a phase contrast image based on a phase change of X-rays due to a difference in the refractive index of the subject instead of a change in X-ray intensity due to a difference in the X-ray absorption ability of the subject has been advanced. It has been. In X-ray phase imaging based on this phase change, a high-contrast image can be acquired even for a weakly absorbing object with low X-ray absorption capability.

  As an X-ray imaging apparatus capable of such X-ray phase imaging, first and second gratings are arranged in parallel at a predetermined interval between an X-ray source and an X-ray image detector. A self-image is formed at the position of the second grating by the Talbot effect by the grating of, and the self-image is partially shielded by the second grating and detected by the X-ray image detector. An X-ray imaging apparatus that acquires a phase contrast image based on this has been proposed (see, for example, Patent Documents 1 and 2).

  The self-image is formed at a position separated by a predetermined distance downstream from the position of the first grating, and this distance is called a Talbot distance. In an X-ray imaging apparatus using the Talbot effect, it is necessary to set the distance between the first grating and the second grating to the Talbot distance. This Talbot distance depends on the wavelength of X-rays emitted from the X-ray source. If the distance between the first grating and the second grating deviates from the Talbot distance, the image quality deteriorates. Therefore, in Patent Document 2, the positions of the first and second gratings are adjusted according to the tube voltage of the X-ray source. Has been proposed to do.

JP 2008-200361 A No. 2008/102598

  However, since the Talbot distance is determined by the wavelength of X-rays incident on the first grating, when the subject is placed between the X-ray source and the first grating, the interaction (absorption) in the subject. The wavelength of the X-rays after passing through the subject changes due to the scattering of the subject, and the Talbot distance changes in the portion corresponding to the subject, so that there is a problem that the image quality of the important subject portion deteriorates.

  The present invention has been made in view of the above problems, and in a radiographic apparatus using the Talbot effect, a radiographic apparatus capable of preventing image quality deterioration of a subject portion due to a change in Talbot distance and control thereof It aims to provide a method.

  In order to achieve the above object, a radiation imaging apparatus of the present invention passes a radiation emitted from a radiation source and generates a first periodic pattern image by the Talbot effect, and the first grating Adjusting the position of at least one of the first and second gratings, a second grating disposed opposite to the grating and partially shielding the first periodic pattern image to generate a second periodic pattern image A position adjustment unit that detects the second periodic pattern image to generate image data, a phase differential image generation unit that generates a phase differential image based on the image data, and the radiation source And the first grating and the second grating by controlling the position adjustment unit based on a parameter related to the wavelength of radiation after passing through the subject arranged between the first grating and the first grating. The distance between Characterized in that it comprises a control unit for setting the Talbot distance caused by radiation after passing through the sample.

  The control unit controls the position adjusting unit based on a Talbot distance determined by a first parameter related to characteristics of radiation emitted from the radiation source and a second parameter representing the type of subject. To do. The first parameter is a set value of the tube voltage of the radiation source.

  The control unit is configured to determine a first energy spectrum determining unit that determines a first energy spectrum of radiation before entering the subject based on the first parameter, and based on the second parameter. A subject transmittance determining unit that determines the transmittance of the sample, a second energy spectrum determining unit that determines a second energy spectrum by multiplying the first energy spectrum by the transmittance, and the second A Talbot distance calculating unit that calculates a center wavelength corresponding to the center energy of the energy spectrum and calculates a Talbot distance from the center wavelength.

  The subject transmittance determining unit determines the transmittance based on a linear attenuation coefficient corresponding to the subject and the thickness of the subject.

  The radiation emitted from the radiation source has a cone beam shape, and the position adjusting unit adjusts the position of the first grating and the first position adjusting part for adjusting the position of the first grating. The position adjustment unit, and the control unit controls the first and second position adjustment units to control the distance between the radiation source and the first grating, and the first position adjustment unit. A distance between the grating and the second grating is set.

  The radiation imaging apparatus of the present invention includes a phase contrast image generation unit that generates a phase contrast image by integrating the phase differential image along a direction orthogonal to the lattice lines of the first and second gratings. .

  Scanning means for changing the relative position of the second grating with respect to the first grating and sequentially setting a plurality of scanning positions, and the radiation image detector includes the second periodic pattern image at each scanning position. The phase differential image generation unit generates a phase differential image based on a plurality of image data generated by the radiation image detector.

  The radiation imaging apparatus of the present invention includes a multi-slit that partially blocks the radiation emitted from the radiation source and disperses the focal point.

  In the method for controlling a radiographic apparatus according to the present invention, a radiation emitted from a radiation source is allowed to pass, and a first grating that generates a first periodic pattern image by the Talbot effect is disposed opposite to the first grating. A second grating that partially shields the first periodic pattern image to generate a second periodic pattern image, and a position adjusting unit that adjusts a position of at least one of the first and second gratings. Control of a radiation imaging apparatus comprising: a radiation image detector that detects the second periodic pattern image and generates image data; and a phase differential image generation unit that generates a phase differential image based on the image data In the method, the position adjustment unit is controlled based on a parameter related to a wavelength of radiation after passing through an object disposed between the radiation source and the first grating, and the first grating and Above The distance between the second grating is set to Talbot distance caused by radiation after passage through the object.

  According to the present invention, the position adjustment unit is controlled based on the parameter related to the wavelength of the radiation after passing through the subject arranged between the radiation source and the first grating, and the first grating and the first grating are controlled. Since the distance between the two grids is set to the Talbot distance generated by the radiation after passing through the subject, image quality deterioration of the subject portion can be prevented.

It is a block diagram which shows the structure of an X-ray imaging apparatus. It is a schematic diagram which shows the structure of a X-ray image detector. It is a schematic side view which shows the structure of the 1st and 2nd grating | lattice. It is a figure explaining the translational movement of a 2nd grating | lattice. It is a graph which illustrates an intensity modulation signal. It is a block diagram which shows the structure of a system control part. It is a graph which illustrates the 1st and 2nd energy spectrum.

(First embodiment)
In FIG. 1, an X-ray imaging apparatus 10 includes an X-ray source 11, an imaging unit 12, a memory 13, an image processing unit 14, an image recording unit 15, an imaging control unit 16, a console 17, and a system control unit 18. . The X-ray source 11 includes, for example, a rotary anode type X-ray tube and a collimator that limits an X-ray irradiation field, and emits X-rays toward the subject H.

  The imaging unit 12 includes an X-ray image detector 20 and first and second gratings 21 and 22. The first and second gratings 21 and 22 are absorption gratings, and are disposed to face the X-ray source 11 with respect to the z direction that is the X-ray irradiation direction. A space is provided between the X-ray source 11 and the first grating 21 so that the subject H can be arranged. The X-ray image detector 20 is a flat panel detector using a semiconductor circuit, and is disposed behind the second grating 22 so that the detection surface 20a is orthogonal to the z direction.

  The first grating 21 includes a plurality of X-ray absorption parts 21a and X-ray transmission parts 21b that are extended in the y direction, which is one direction in the grating plane orthogonal to the z direction. The X-ray absorption part 21a and the X-ray transmission part 21b are alternately arranged along the x direction orthogonal to the z direction and the y direction, forming a striped pattern. Similar to the first grating 21, the second grating 22 includes a plurality of X-ray absorbing parts 22a and X-ray transmitting parts 22b that are extended in the y direction and are alternately arranged along the x direction. The X-ray absorption parts 21a and 22a are made of a metal having X-ray absorption such as gold (Au) or platinum (Pt). The X-ray transmission portions 21b and 22b are made of a material having X-ray transmission properties such as silicon (Si) or resin.

  X-rays radiated from the X-ray source 11 have a spatial coherence sufficient to generate a Talbot effect on the X-rays passing through the first grating 21. The first grating 21 partially transmits X-rays emitted from the X-ray source 11 and generates a self-image (first periodic pattern image) at a position separated by the Talbot distance due to the Talbot effect. The second grating 22 partially transmits the first periodic pattern image (hereinafter referred to as G1 image) generated by the first grating 21 and transmits the second periodic pattern image (hereinafter referred to as G2 image). Generate. The G1 image substantially matches the lattice pattern of the second lattice 22.

  The X-ray image detector 20 detects the G2 image and generates image data. The memory 13 temporarily stores the image data read from the X-ray image detector 20. The image processing unit 14 includes a phase differential image generation unit 14a and a phase contrast image generation unit 14b. The phase differential image generation unit 14 a generates a phase differential image based on the image data stored in the memory 13. The phase contrast image generation unit 14b generates a phase contrast image based on the phase differential image. The image recording unit 15 records a phase differential image and a phase contrast image. The imaging control unit 16 controls the X-ray source 11 and the imaging unit 12.

  The imaging unit 12 is provided with a scanning mechanism 23 that translates the second grating 22 in the x direction and changes the relative position of the second grating 22 with respect to the first grating 21. The scanning mechanism 23 is configured by an actuator such as a piezoelectric element. The scanning mechanism 23 is driven based on the control of the imaging control unit 16 at the time of stripe scanning described later. As will be described in detail later, the memory 13 stores image data photographed by the X-ray image detector 20 at each scanning position of fringe scanning.

  The imaging unit 12 includes a first position adjusting unit 24 that adjusts the position of the first grating 21 in the z direction, and a second position adjusting unit 25 that adjusts the position of the second grating 22 in the z direction. And are provided.

  The console 17 displays an imaging unit, a phase differential image, a phase contrast image, and the like, and an operation unit 17a that enables operations such as setting of imaging conditions (tube voltage, type of subject H, etc.) and an imaging execution instruction. A monitor 17b is provided. The system control unit 18 comprehensively controls each unit in accordance with a signal input from the operation unit 17a.

  In FIG. 2, an X-ray image detector 20 collects charges generated in a semiconductor film such as amorphous selenium (a-Se) by incident X-rays, and reads out the charges collected by the pixel electrodes 31. A plurality of pixels 30 each having a thin film transistor (TFT) 32 are two-dimensionally arranged. The X-ray image detector 20 includes a gate scanning line 33 provided for each row of the pixels 30, a scanning circuit 34 that applies a scanning signal for turning on and off the TFT 32 to each gate scanning line 33, A signal line 35 is provided for each column, and a readout circuit 36 that reads out charges from the pixels 30 via the signal lines 35, converts the charges into image data, and outputs the image data. The detailed layer configuration of each pixel 30 is the same as the layer configuration described in JP-A-2002-26300.

    The readout circuit 36 includes an integration amplifier, an A / D converter, a correction circuit (none of which is shown), and the like. The integrating amplifier integrates the charges output from each pixel 30 through the signal line 35 to generate an image signal. The A / D converter converts the image signal generated by the integrating amplifier into digital image data. The correction circuit performs dark current correction, gain correction, linearity correction, and the like on the image data, and inputs the corrected image data to the memory 13.

  The X-ray image detector 20 is not limited to a direct conversion type that directly converts incident X-rays into electric charges, but converts incident X-rays into visible light with a scintillator such as cesium iodide (CsI) or gadolinium oxysulfide (GOS). Alternatively, an indirect conversion type in which visible light is converted into electric charge by a photodiode may be used. Further, a CMOS sensor or the like can be used instead of the TFT.

In FIG. 3, the X-rays emitted from the X-ray source 11 are cone beams having the X-ray focal point 11a as a light emission point. G1 image by the first grating 21, resulting from the first grating 21 in a position apart Talbot distance Z _m downstream. The grating pitch _{p 1} of the first grating 21, the grating pitch _{p 2} of the second grating 22, the X-ray wavelength λ incident on the first grating 21, with the positive integer m, the Talbot distance _{Z m,} It is represented by the following formula (1).

The G1 image is enlarged in proportion to the distance from the X-ray focal point 11a. In order for the G1 image to substantially match the grating pattern of the second grating 22, the distance between the X-ray focal point 11 a and the first grating 21 is set to L ₁ , and the distance between the first grating 21 and the second grating 22 is the distance between the _{L 2,} the grating pitch _{p 1} of the first grating 21, the grating pitch _{p 2} of the second grating 22, it is necessary to satisfy the following expression (2) approximately.

In the present embodiment, m = 1 is set, the distance L ₂ is set so as to satisfy L ₂ = Z ₁ based on the above formula (1), and the distance L ₁ is set based on the above formula (2). The distances L ₁ and L ₂ are set by the first and second position adjusting units 24 and 25 described above.

  The X-ray irradiated from the X-ray source 11 undergoes a phase shift due to the subject H when passing through the subject H, so that the wavefront of the X-ray incident on the first grating 21 is distorted, and G1 is accordingly generated. The image is deformed. Moire fringes occur in the G2 image, but the moire fringes are modulated by the subject H. This modulation amount is proportional to the refraction angle of the X-ray refracted by the subject H. Therefore, the phase information of the subject H can be acquired by detecting the G2 image with the X-ray image detector 20 and analyzing the image data generated by the X-ray image detector 20.

In the present embodiment, the fringe scanning method is used to acquire the phase information of the subject H. In the fringe scanning method, a G2 image is taken at a plurality of predetermined scanning positions while one of the first and second gratings 21 and 22 is translated (scanned) in the x direction relative to the other. In the present embodiment, the first grating 21 is fixed, and the second grating 22 is moved in the x direction by the scanning mechanism 23. Moire fringes generated in the G2 image move as the second grating 22 moves, and return to the original position when the moving distance in the x direction reaches the grating period (grating pitch p ₂ ) of the second grating 22.

FIG. 4 schematically shows a state in which the value (p ₂ / M) obtained by dividing the arrangement pitch p ₂ into M (integers of 2 or more) is used as the scanning pitch, and the second grating 22 is translated in each scanning pitch. It shows. The scanning mechanism 23 moves the second grating 22 in order to M scanning positions of k = 0, 1, 2,..., M−1.

  At the position of k = 0, X-rays that are emitted from the X-ray source 11 and are not refracted by the subject H mainly pass through the second grating 22. When the second grating 22 is moved in order of k = 1, 2,..., the components that are not refracted by the subject H are reduced in the X-rays passing through the second grating 22. The component refracted by the subject H increases. In particular, at the position of k = M / 2, the X-ray passing through the second grating 22 is almost only the component refracted by the subject H. When the position exceeds k = M / 2, the component of the X-ray passing through the second grating 22 is refracted by the subject H while the component not refracted by the subject H is increased.

  When the G2 image is captured by the X-ray image detector 20 at each scanning position of k = 0, 1, 2,..., M−1, M images generated by the X-ray image detector 20 are captured. Data is stored in the memory 13.

  The M pixel values obtained for each pixel 30 periodically change with respect to the scanning position k as illustrated in FIG. This is called an intensity modulation signal. The broken line in the figure illustrates the intensity modulation signal obtained without the subject H being placed. On the other hand, the solid line illustrates the intensity modulation signal in which the phase shift amount ψ (x, y) is generated due to the influence of the subject H in a state where the subject H is arranged. Here, x and y indicate the coordinates of the pixel 30.

Next, the calculation principle of the phase shift amount ψ (x, y) will be described. The intensity modulation signal representing the change of the pixel value P _k (x, y) with respect to the scanning position k is generally represented by the following expression (3).

Here, A ₀ corresponds to the intensity of the incident X-ray, A _n is a value corresponding to the amplitude of the intensity-modulated signal. n is a positive integer and i is an imaginary unit.

The grating pitch p ₂ is equally divided, in the case where the scanning pitch is constant, the following equation (4) is satisfied. When this relational expression is applied, the phase shift amount ψ (x, y) is expressed by the following expression (5).

  Further, the phase shift amount ψ (x, y) is expressed by the following equation (6) using a trigonometric function. Hereinafter, the phase shift amount ψ (x, y) is referred to as a phase differential value ψ (x, y).

  The phase differential image generation unit 14a generates a phase differential image by performing an operation using the above equation (9) based on the M pieces of image data acquired by fringe scanning and stored in the memory 13. The phase contrast image generation unit 14b generates a phase contrast image by integrating the phase differential image generated by the phase differential image generation unit 14a along the x direction.

  The system control unit 18 includes a CPU, a RAM, a ROM, and the like. As shown in FIG. 6, the first energy spectrum determination unit 40, the subject transmittance determination unit 41, the second energy spectrum determination unit 42, and the Talbot distance. The calculation unit 43 is configured.

The first energy spectrum determination unit 40 radiates from the X-ray source 11 and enters the subject H based on the set value (first parameter) of the tube voltage input by the operation unit 17a. The first energy spectrum I ₁ (E) is determined as shown in FIG. The determination of the energy spectrum is performed by selecting the one closest to the set value of the tube voltage from a plurality of energy spectra stored in advance, or by calculating based on a theoretical formula of the energy spectrum. FIG. 7 shows the energy spectrum of continuous X-rays emitted from the X-ray source 11.

The subject transmittance determination unit 41 uses the following equation (7) based on the type (second parameter) of the subject H input by the operation unit 17a, and transmits the transmittance T _t ( E) is determined.

Here, μ (E) is a linear attenuation coefficient that represents the attenuation due to the interaction (absorption, scattering) of X-rays with the subject H, and depends on the energy E of the X-rays. t represents the thickness of the subject H with respect to the X-ray transmission direction (z direction). The subject transmittance determination unit 41 holds a linear attenuation coefficient μ (E) and a thickness t corresponding to the type of the subject H, and applies these to the above equation (7) to obtain the transmittance T _t ( E).

As shown in the following formula (8), the second energy spectrum determination unit 42 multiplies the first energy spectrum I ₁ (E) by the transmittance T _t (E), thereby obtaining the second energy spectrum I. ₂ Generate (E).

Since the X-ray has a characteristic that the smaller the energy E, the easier it is absorbed by the subject H (the linear attenuation coefficient μ (E) is larger as the energy E is smaller). Therefore, as shown in FIG. The center energy E _c2 of ₂ (E) is shifted to a higher energy side than the center energy E _{c1 of} the first energy spectrum I ₁ (E).

The Talbot distance calculation unit 43 calculates the center energy E _c2 from the second energy spectrum I ₂ (E), calculates the center wavelength λ ₀ corresponding to the center energy E _c2 from the relationship E = hc / λ, and this wavelength By substituting λ into the above equation (1), the Talbot distance Z ₁ is calculated. Here, h is the Planck constant and c is the speed of light.

The system control unit 18 uses the Talbot distance Z ₁ calculated by the Talbot distance calculation unit 43 to satisfy L ₂ = Z _{1 and} to satisfy the above equation (2). The distances L ₁ and L ₂ are set by controlling 24 and 25.

Next, the operation of the X-ray imaging apparatus 10 configured as described above will be described. First, the technician inputs the tube voltage of the X-ray source 11 and the type of the subject H from the operation unit 17a. In this embodiment, the body part (finger or knee) and the classification of adult and child are used as the type of the subject H. When the tube voltage is input, the first energy spectrum determination unit 40 determines the first energy spectrum I ₁ (E) corresponding to the input tube voltage. When the type of the subject H is input, the subject transmittance determining unit 41 determines the line attenuation coefficient μ (E) and the thickness t corresponding to the input type of the subject H, and the above formula (7) is used to determine the transmittance T _t (E) of the X-ray subject H.

When the first energy spectrum I ₁ (E) and the transmittance T _t (E) are determined, the second energy spectrum is then determined by the second energy spectrum determination unit 42 using the above equation (8). I ₂ (E) is determined. Then, the Talbot distance calculation unit 43 _obtains the center wavelength λ ₀ corresponding to the center energy E _c2 of the second energy spectrum I ₂ (E), and calculates the Talbot distance Z ₁ using the above equation (1). Is done.

The center energy E _c2 and the center wavelength λ ₀ corresponding to the tube voltage and the type of the subject H are acquired as shown in Table 1 below, for example. Table 1 below assumes that the component of the subject H is water, the child's finger thickness is 15 mm, the adult finger thickness is 20 mm, the child's knee thickness is 60 mm, and the adult knee thickness is 100 mm. The value calculated using the linear attenuation coefficient μ (E) is shown.

When the Talbot distance Z ₁ is calculated, the system control unit 18 controls the first and second position adjustment units 24 and 25 so that L ₂ = Z ₁ and satisfy the above equation (2). The positions of the first and second gratings 21 and 22 are adjusted. When this adjustment is completed, the system control unit 18 displays on the monitor 17b that the preparation for photographing has been completed.

Next, when an imaging instruction is input from the operation unit 17a, the X-ray source 11 is scanned at each scanning position k while the second grating 22 is translated by a predetermined scanning pitch (p ₂ / M) by the scanning mechanism 23. X-ray irradiation and detection of the G2 image by the X-ray image detector 20 are performed. As a result, M pieces of image data are generated and stored in the memory 13.

  Thereafter, the M image data stored in the memory 13 is read by the image processing unit 14. In the image processing unit 14, a phase differential image is generated by the phase differential image generation unit 14a, and a phase contrast image is generated by the phase contrast image generation unit 14b. The phase contrast image and the phase differential image are recorded in the image recording unit 15, then input to the console 17, and displayed on the monitor 17b.

In the first embodiment, after passing through the subject H by the first energy spectrum determining unit 40, the subject transmittance determining unit 41, the second energy spectrum determining unit 42, and the Talbot distance calculating unit 43. However, instead of this, a lookup table (LUT) as shown in Table 1 is prepared in advance, and by referring to this LUT, the center wavelength λ ₀ is obtained. ₀ may be obtained.

In the first embodiment, the tube wavelength of the X-ray source 11 is the first parameter, the type of the subject H is the second parameter, and the X-ray central wavelength λ ₀ after passing through the subject H is obtained. Although determined, the first parameter may represent the characteristic of X-rays emitted from the X-ray source 11 and is not limited to the tube voltage. As the first parameter, tube current, anode target type, or a combination thereof may be used.

In the first embodiment, the positions of the first and second gratings 21 and 22 are adjusted in order to set the distances L ₁ and L ₂ . The position of one of the two gratings 21 and 22 and the X-ray source 11 may be adjusted. Further, the X-ray image detector 20 may be moved together with the second grating 22.

  In the first embodiment, the phase shift value of the intensity modulation signal is used as the phase differential value. However, a value obtained by multiplying this by a constant may be used as the phase differential value.

  In the first embodiment, the phase differential image is generated. However, in addition to this, an absorption image and a small angle scattered image may be generated. The absorption image is generated by obtaining an average value of the intensity modulation signals exemplified in FIG. The small angle scattered image is generated by obtaining the amplitude of the intensity modulation signal.

In the first embodiment, the first grating 21 is an absorption type grating. However, the first grating 21 may be a phase type grating that applies phase modulation to X-rays. When the first grating 21 is a phase-type grating that gives a phase modulation of π / 2, the Talbot distance Z _m is expressed by the following expression (9). Here, m is 0 or a positive integer.

Further, when the first grating 21 is a phase type grating that gives a phase modulation of π, the Talbot distance Z _m is expressed by the following expression (10). Here, m is 0 or a positive integer.

(Second Embodiment)
Next, a second embodiment of the present invention will be described. In the first embodiment, the X-ray emitted from the X-ray source 11 is a cone beam. In the X-ray imaging apparatus of the second embodiment, an X-ray source that emits a parallel beam is used. Instead of (2), the first and second gratings 21 and 22 are configured to substantially satisfy p ₂ = p ₁ . As described above, when an X-ray source that emits a parallel beam is used, the relationship between the distance L ₁ and the distance L ₂ is not restricted by the above equation (2), and thus the first and second position adjusting units 24 and 25. Only one of these may be provided.

When the X-rays emitted from the X-ray source 11 are parallel beams and the first grating 21 is an absorption type grating, the Talbot distance Z _m is expressed by the following expression (11). Here, m is a positive integer.

Further, when the X-rays emitted from the X-ray source 11 are parallel beams and the first grating 21 is a phase-type grating giving a phase modulation of π / 2, the Talbot distance Z _m can be expressed by the following formula ( 12). Here, m is 0 or a positive integer.

When the X-ray emitted from the X-ray source 11 is a parallel beam and the first grating 21 is a phase-type grating that gives a phase modulation of π, the Talbot distance Z _m is expressed by the following equation (13). It is represented by Here, m is 0 or a positive integer.

(Third embodiment)
Next, a third embodiment of the present invention will be described. In the first embodiment, the X-ray source 11 has a single focal point. However, in the X-ray imaging apparatus of the second embodiment, immediately after the emission side of the X-ray source 11, it is described in WO2006 / 131235. A multi-slit (source grid) is provided to disperse the focal point. As a result, the amount of light of the G2 image is improved, and the calculation accuracy of the phase differential image can be improved and the photographing time can be shortened. The multi-slit pitch p ₀ needs to satisfy the following expression (14). In the present embodiment, the distance L ₁ is a distance from the multi slit to the first grating 21. Other configurations and operations are the same as those in the first embodiment.

(Fourth embodiment)
Next, a fourth embodiment of the present invention will be described. In the first embodiment, the phase differential image is generated based on a plurality of image data obtained by performing imaging while relatively moving the first and second gratings 21 and 22. It is also possible to generate a phase differential image based on single image data obtained by one imaging while the second gratings 21 and 22 are fixed.

  The X-ray imaging apparatus of this embodiment generates a phase differential image by the Fourier transform method described in WO2010 / 050484. Specifically, the phase differential image generation unit acquires a Fourier spectrum by performing Fourier transform on a single image data obtained by the X-ray image detector 20, and corresponds to the carrier frequency from this Fourier spectrum. The obtained spectrum is separated and further subjected to inverse Fourier transform to generate a phase differential image. In the present embodiment, the scanning mechanism 23 is not necessary. Other configurations and operations are the same as those in the first embodiment.

  The present invention is not limited to a radiographic apparatus for medical diagnosis, but can be applied to other radiographic apparatuses for industrial use. Further, the radiation is not limited to X-rays, and gamma rays can also be used.

DESCRIPTION OF SYMBOLS 10 X-ray imaging apparatus 20 X-ray image detector 21 1st grating | lattice 21a X-ray absorption part 21b X-ray transmission part 22 2nd grating | lattice 22a X-ray absorption part 22b X-ray transmission part 30 Pixel 31 Pixel electrode 33 Gate scanning line 35 signal lines

Claims (10)

A first grating that transmits radiation emitted from a radiation source and generates a first periodic pattern image by the Talbot effect;
A second grating disposed opposite to the first grating and partially shielding the first periodic pattern image to generate a second periodic pattern image;
A position adjusting unit that adjusts the position of at least one of the first and second gratings;
A radiation image detector for detecting the second periodic pattern image and generating image data;
A phase differential image generation unit that generates a phase differential image based on the image data;
The position adjustment unit is controlled based on a parameter related to the wavelength of the radiation after passing through the subject arranged between the radiation source and the first grating, and the first grating and the second grating are controlled. A control unit for setting the distance between the grid and the Talbot distance generated by the radiation after passing through the subject;
A radiation imaging apparatus comprising:   The control unit controls the position adjusting unit based on a Talbot distance determined by a first parameter related to characteristics of radiation emitted from the radiation source and a second parameter representing the type of subject. The radiation imaging apparatus according to claim 1, wherein:   The radiation imaging apparatus according to claim 2, wherein the first parameter is a set value of a tube voltage of the radiation source. The controller is configured to determine a first energy spectrum of radiation before entering the subject based on the first parameter;
A subject transmittance determining unit that determines the transmittance of the subject of radiation based on the second parameter;
A second energy spectrum determination unit that determines a second energy spectrum by multiplying the first energy spectrum by the transmittance;
Calculating a center wavelength corresponding to the center energy of the second energy spectrum, and calculating a Talbot distance from the center wavelength;
The radiation imaging apparatus according to claim 2, further comprising:   The radiation imaging apparatus according to claim 4, wherein the subject transmittance determining unit determines the transmittance based on a linear attenuation coefficient corresponding to the subject and a thickness of the subject. The radiation emitted from the radiation source is cone-shaped,
The position adjustment unit includes a first position adjustment unit that adjusts the position of the first grating, and a second position adjustment unit that adjusts the position of the second grating,
The control unit controls the first and second position adjusting units to thereby determine a distance between the radiation source and the first grating, and the first grating and the second grating. The radiation imaging apparatus according to claim 1, wherein a distance between the two is set.   The phase contrast image generation part which produces | generates a phase contrast image by integrating the said phase differential image along the direction orthogonal to the lattice line of the said 1st and 2nd grating | lattice is provided. 6. The radiographic apparatus according to any one of items 6. Scanning means for changing a relative position of the second grating with respect to the first grating and sequentially setting a plurality of scanning positions;
The radiation image detector detects the second periodic pattern image at each scanning position to generate image data;
The radiographic apparatus according to claim 1, wherein the phase differential image generation unit generates a phase differential image based on a plurality of image data generated by the radiological image detector. .   The radiation imaging apparatus according to claim 1, further comprising a multi-slit that partially blocks the radiation emitted from the radiation source and disperses the focal point. A first grating that transmits radiation emitted from a radiation source and generates a first periodic pattern image by the Talbot effect;
A second grating disposed opposite to the first grating and partially shielding the first periodic pattern image to generate a second periodic pattern image;
A position adjusting unit that adjusts the position of at least one of the first and second gratings;
A radiation image detector for detecting the second periodic pattern image and generating image data;
A phase differential image generation unit that generates a phase differential image based on the image data;
In a method for controlling a radiographic apparatus comprising:
The position adjustment unit is controlled based on a parameter related to the wavelength of the radiation after passing through the subject arranged between the radiation source and the first grating, and the first grating and the second grating are controlled. A control method characterized in that the distance between the first and second gratings is set to a Talbot distance generated by radiation after passing through the subject.

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