JP2014030440A - Radiation imaging system and control method therefor - Google Patents

Abstract

In energy subtraction imaging, occurrence of artifacts due to body movement of a subject is reduced.
Imaging is performed by switching the energy of X-rays emitted from an X-ray source between low energy and high energy at each relative position when the second grid is relatively scanned with respect to the first grid. The X image detector detects a periodic pattern image of X-rays that has passed through the first and second grids, and generates first and second image data for low energy and high energy, respectively. Then, a difference image is calculated by subtracting the other from one of the image signals of the first and second phase differential images calculated based on the first and second image data.
[Selection] Figure 6

Description

  The present invention relates to a radiographic imaging system using radiation such as X-rays and a control method thereof.

  It is known that when X-rays are incident on an object, the intensity and phase change due to the interaction, and the phase change (angle change) is larger than the intensity change. Using this X-ray property, X-ray phase imaging is used to obtain a high-contrast phase image (hereinafter referred to as a phase contrast image) from a subject having a low X-ray absorption capacity based on the phase change of the X-ray by the subject. Research is actively conducted and various methods are being developed.

  For example, Patent Document 1 describes a diagnostic information generation system using an enlarged imaging system and a microfocus X-ray source. This system irradiates a subject with X-rays emitted from a microfocus X-ray light source, and the shift of the image due to the X-rays refracted by the subject is caused by blurring due to geometric instability due to the focal point size of the X-ray source. This is a method of detecting with a detector that is sufficiently separated from the subject so that it is spatially separated. By this method, it is possible to acquire a clearer and easier-to-see image by enhancing the outline of the conventional absorption contrast image.

  Further, Patent Document 1 discloses an energy for capturing a phase contrast image using X-rays in a low energy band and a phase contrast image using X-rays in a high energy band and performing image subtraction processing between both image data. Subtraction shooting is described.

  However, since the refraction angle of X-rays generated in the subject is extremely small, the above-described method using the magnified imaging system and the microfocus X-ray source has an X-ray image detector having high spatial resolution and an X-ray with a small focus. A radiation source is required. Compared with the above method, in the Talbot interferometry, the phase change of the wave front caused when X-rays pass through the subject, that is, the refraction angle of the X-rays refracted by the subject is determined from the intensity change of the pixel data. By calculating, it is possible to more clearly depict the structure of soft tissue that was difficult to depict. Furthermore, by using a source grating, a normal X-ray source can be used while satisfying the requirement for a fine focal spot size.

  Patent Document 2 describes an X-ray imaging apparatus using Talbot interferometry. In this X-ray imaging apparatus, first and second grids are arranged between a subject and an X-ray image detector, and X-ray refraction generated by a slight refractive index distribution inside the subject is detected. In this method, the intensity of the image of the first grid is converted into a change in intensity and detected by the X-ray image detector via the second grid arranged behind the first grid. Specifically, based on a method called a fringe scanning method, the phase differential image is calculated by calculating the phase shift amount of the intensity modulation signal obtained by shifting the relative position of the second grid with respect to the first grid for each pixel. And a phase contrast image of the subject is obtained by integrating the phase differential image. X-ray Talbot interference is also described in Non-Patent Document 1, and the fringe scanning method in the Talbot interferometer is also disclosed in Non-Patent Document 2.

JP 2008-018059 A Japanese Patent No. 4445397

C. David, et al., Applied Physics Letters, Vol. 81, No. 17, October 2002, p. 3287 Hector Canabal, et al., Applied Optics, Vol.37, No.26, September 1998, 6227

  In the X-ray phase imaging using the Talbot interferometry, the structure of the subject surface such as the unevenness of the skin and the pores can be drawn in addition to the tissue to be subjected to the image diagnosis such as the bone and soft tissue inside the subject. However, in some cases, these may become obstacle shadows and hinder image diagnosis, so it is desirable to remove them from the image appropriately. For example, in a conventional absorption contrast image, an energy subtraction technique that removes an obstruction shadow using an image having different energy is used.

  However, in the X-ray phase imaging using the Talbot interferometry, a series of imaging is performed while changing the relative positions of the first grid and the second grid. Therefore, a series of X-rays using a low energy band is used. When a positional shift due to body movement of the subject occurs between imaging and a series of imaging using X-rays in a high energy band, there is a problem that artifacts due to the positional shift occur in the energy subtraction image. .

  The present invention has been made in view of the above problems, and provides a radiographic imaging system and a control method thereof that can reduce the occurrence of artifacts due to body movement of a subject in energy subtraction imaging. For the purpose.

  The radiographic imaging system of the present invention partially passes a first grid for generating a first periodic pattern image by passing radiation emitted from a radiation source and a periodic pattern image of the first image. A second grid for generating a second periodic pattern image; scanning means for changing the relative position of the second grid with respect to the first grid by a predetermined amount; and radiation from the radiation source at each relative position. Control means for switching the energy of the radiation to a plurality of types of energy and the second periodic pattern image corresponding to each energy at each relative position to detect image data corresponding to each energy. A radiation image detector; phase image generating means for generating a phase image corresponding to each energy based on image data corresponding to each energy; and the plurality of types Characterized by comprising a differential image generating means for generating a difference image by the phase images difference processing, the. Here, the difference processing is not limited to simple difference processing, and includes performing difference processing after weighting, filtering, and the like of the phase image.

  The control means converts the energy of the radiation radiated from the radiation source at each relative position into two types of energy: a first energy and a second energy that is lower than the first energy. It is preferable to switch.

  Moreover, it is preferable that the said control means switches the energy of a radiation in order of 1st energy and 2nd energy in each said relative position.

  The phase image is preferably a phase differential image. Instead of this, the phase image may be a phase contrast image obtained by integrating the phase differential image.

  Further, the subject is arranged by the correction data storage means for storing the phase image for each energy obtained in a state where the subject is not arranged as correction data, and the correction data stored in the correction data storage means. It is preferable to further include correction processing means for correcting the phase image for each energy obtained in the state.

  Preferably, the first grid is an absorption type grating, and the first periodic pattern image is formed as a projection image. Instead of this, the first grid may be a phase grating, and the first periodic pattern image may be formed as a self-image by the Talbot interference effect.

  Moreover, it is preferable to provide a radiation source having a radiation source grid that generates a large number of point light sources by selectively shielding radiation.

  Furthermore, the control method of the present invention partially passes the first grid for generating the first periodic pattern image by passing the radiation emitted from the radiation source, and the periodic pattern image of the first image. A second grid for generating a second periodic pattern image; scanning means for changing a relative position of the second grid with respect to the first grid by a predetermined amount; and the second periodic pattern at each relative position. A radiation imaging system control method comprising a radiation image detector that detects an image and generates image data, respectively, wherein the radiation energy emitted from the radiation source at each relative position is a plurality of types of energy. And detecting the second periodic pattern image corresponding to each energy at each relative position by the radiation image detector. The image data is generated which, based on said image data corresponding to each energy generates a differential image by performing the generating and difference processing phase image corresponding to each energy.

  According to the present invention, energy subtraction imaging is performed by switching the energy of radiation radiated from the radiation source at each relative position of the first and second grids to a plurality of types of energy, so the first and second grids Therefore, the relative scanning of the subject is only once, and the total imaging time is shortened. Therefore, the influence of the body movement of the subject is suppressed, and the occurrence of artifacts due to the body movement is reduced.

It is a schematic diagram which shows the structure of the X-ray image imaging system of 1st Embodiment. It is a block diagram which shows the structure of an image process part. 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 grid. It is explanatory drawing explaining a fringe scanning method. It is a flowchart explaining a fringe scanning method. It is a graph which illustrates an intensity modulation signal. It is a perspective view which shows the source grid used in other embodiment. It is a block diagram which shows the image process part used by other embodiment.

(First embodiment)
In FIG. 1, an X-ray imaging system 10 according to the first embodiment of the present invention 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 the X-ray irradiation field, and emits X-rays to the subject H.

  The X-ray imaging system 10 performs imaging using an X-ray in a low energy band (hereinafter referred to as low energy imaging) and imaging using an X-ray in a high energy band in order to perform image subtraction processing in the image processing unit 14. (Hereinafter referred to as high energy imaging). For example, low energy imaging and high energy imaging are performed by changing the tube voltage applied to the X-ray source 11.

  The imaging unit 12 includes an X-ray image detector 20 and first and second grids 21 and 22. The first grid 21 and the second grid 22 are absorption type grids, and are disposed to face the X-ray source 11 in the z direction that is the X-ray irradiation direction. A space is provided between the X-ray source 11 and the first grid 21 so that the subject H can be arranged. The X-ray image detector 20 is, for example, a flat panel detector (FPD) using a semiconductor circuit, and is arranged behind the second grid 22 so that the detection surface is orthogonal to the z direction. ing.

  The first grid 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 a 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, and form a striped grid. The second grid 22 includes a plurality of X-ray absorbing portions 22 a and X-ray transmitting portions 22 b that extend in the y direction and are alternately arranged along the x direction, like the first grid 21. 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.

  The memory 13 temporarily stores the image data read from the photographing unit 12. The image processing unit 14 generates a phase contrast image based on a plurality of image data stored in the memory 13. The image recording unit 15 records the phase contrast image generated by the image processing unit 14. The imaging control unit 16 controls the X-ray source 11 and the imaging unit 12. The console 17 includes an input device such as a keyboard and a display device such as a monitor, and enables input of shooting conditions and shooting instructions, and display of shooting information and images. The system control unit 18 comprehensively controls each unit in response to an input instruction from the input device of the console 17.

  The imaging unit 12 is provided with a scanning mechanism 23 that translates the second grid 22 in the x direction and changes the relative position of the second grid 22 with respect to the first grid 21. The scanning mechanism 23 is configured by an actuator such as a piezoelectric element, for example. 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 captured by the X-ray image detector 20 in each scanning step of fringe scanning. The image data is divided into image data obtained by low energy imaging (hereinafter referred to as first image data) and image data obtained by high energy imaging (hereinafter referred to as second image data).

  In FIG. 2, the image processing unit 14 includes a phase differential image generation unit 30, a first image storage unit 31, a second image storage unit 32, an image subtraction processing unit 33, and a phase contrast image generation unit 34. The phase differential image generation unit 30 generates a first phase differential image of the subject H based on the plurality of first image data stored in the memory 13, and the plurality of second differential images stored in the memory 13. A second phase differential image of the subject H is generated based on the image data. The first image storage unit 31 stores a first differential phase image, and the second image storage unit 32 stores a second differential phase image.

  The image subtraction processing unit 33 reads the first and second phase differential images from the first and second image storage units 31 and 32, respectively, and the image of the second phase differential image from the image signal of the first phase differential image. The difference image is obtained by subtracting the signal. Note that the image subtraction processing unit 33 is not limited to simple difference processing, and performs difference processing after applying intensity weighting, filter processing, and the like to at least one pixel signal of the first and second phase differential images. It may be.

  In general, the refraction angle of X-rays by the subject H is larger at low energy than at high energy, and the detection sensitivity of phase information is larger at low energy. For this reason, when the subject H has a very low phase contrast such as a cartilage portion, the first phase differential image obtained by low energy imaging with enhanced detection sensitivity of phase information is used as a diagnostic image. The second phase differential image obtained by high-energy imaging is used as an image for removing obstacle shadows such as bones, and the second phase differential image is obtained from the image signal of the first phase differential image as described above. It is preferable to subtract the signal. Thereby, the difference image becomes an image (contour image) in which the bone part is removed and the cartilage part is clarified. In order to increase the S / N ratio of the diagnostic image, it is preferable to make the X-ray dose at the time of low energy imaging larger than the X-ray dose at the time of high energy imaging.

  The phase contrast image generation unit 34 generates a phase contrast image by integrating the difference image obtained by the image subtraction processing unit 33 along a direction corresponding to the x direction. The phase contrast image generated by the phase contrast image generation unit 34 is recorded in the image recording unit 15 and then output to the console 17 and displayed on the monitor.

  In FIG. 3, an X-ray image detector 20 includes an image receiving unit 41 in which a plurality of pixels 40 that convert X-rays into electric charges and are stored are two-dimensionally arranged on an active matrix substrate along the x and y directions. And a scanning circuit 42 that controls the readout timing of the charge from the pixel 40, and a readout circuit 43 that reads out the charge from the pixel 40, converts the charge into image data, and outputs the image data. The scanning circuit 42 and each pixel 40 are connected by a scanning line 44 for each row, and the readout circuit 43 and each pixel 40 are connected by a signal line 45 for each column. The arrangement pitch of the pixels 40 is about 100 μm in each of the x direction and the y direction.

  The pixel 40 directly converts X-rays into charges by a conversion layer (not shown) such as amorphous selenium, and directly stores the converted charges in a capacitor (not shown) connected to an electrode below the conversion layer. This is a conversion type X-ray detection element. Each pixel 40 is provided with a TFT switch (not shown). The gate electrode of the TFT switch is connected to the scanning line 44, the source electrode is connected to the capacitor, and the drain electrode is connected to the signal line 45. 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 45.

Note that the pixel 40 temporarily converts X-rays into visible light using a scintillator (not shown) made of gadolinium oxide (Gd ₂ O ₃ ), cesium iodide (CsI), or the like, and converts the converted visible light into a photodiode ( It is also possible to use an indirect conversion type X-ray detection element that converts the charge into a charge and stores it. The X-ray image detector 20 is not limited to an FPD based on a TFT panel, and a radiation image detector 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, 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 40 via the signal line 45 and converts them into a voltage signal (image signal). The A / D converter converts the image signal converted by the integrating amplifier into digital image data. The correction circuit performs dark current correction, gain correction, linearity correction, and the like on each pixel data constituting the image data, and inputs the corrected image data to the memory 13.

In FIG. 4, X-rays emitted from the X-ray source 11 are cone beams having the X-ray focal point 11 a as a light emission point, and a first pattern of X-rays generated by passing through the first grid 21. The image (hereinafter referred to as G1 image) is enlarged in proportion to the distance from the X-ray focal point 11a. The arrangement pitch p ₂ and the width d _{2 in} the x direction of the X-ray absorber 22 a of the second grid 22 are the distance L ₁ between the X-ray focal point 11 a and the first grid 21, the first grid 21 and the first grid 21. The distance L ₂ between the two grids 22 and the arrangement pitch p ₁ and the width d ₁ of the X-ray absorber 21 a of the first grid 21 are determined as shown in the following equations (1) and (2). Is done.

For example, the arrangement pitch p ₂ is 5 μm and the width d ₂ is half that of 2.5 μm. The thickness of the X-ray absorber 21a in the z direction is, for example, about 100 μm in consideration of corneal X-ray vignetting emitted from the X-ray source 11.

The first and second grids 21 and 22 are configured to linearly project the X-rays that have passed through the X-ray transmission parts 21b and 22b. Specifically, by setting the width of the X-ray transmission parts 21b and 22b in the x direction (same as the widths d ₁ and d ₁ ) to a value sufficiently larger than the peak wavelength of the X-rays emitted from the X-ray source 11. The first and second grids 21 and 22 allow most of the X-rays to pass through without being diffracted while maintaining straightness. For example, when tungsten is used as the rotating anode of the X-ray tube of the X-ray source 11 and the tube voltage is 50 kV, the peak wavelength of the X-ray is about 0.4 mm. In this case, a range of about 1 to 10 μm is allowed as the width of the X-ray transmission parts 21b and 22b.

In the case of the Talbot interferometer, the distance L ₂ is limited to the Talbot interference distance. However, in the present embodiment, the first and second grids 21 and 22 project X-rays linearly, and thus the distance L _2. Can be set independently of the Talbot interference distance.

As described above, the imaging unit 12 of the present embodiment does not constitute a Talbot interferometer, but the Talbot when it is assumed that X-ray diffraction occurs in the first grid 21 and a Talbot interference effect occurs. The interference distance Z is expressed by the following expression (3) using the arrangement pitches p ₁ and p ₂ , the X-ray wavelength (peak wavelength) λ, and a positive integer m.

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

In the present embodiment, it is possible to set the distance L ₂ as described above irrespective of the Talbot distance, for the purpose of thinning in the z-direction of the imaging unit 12, the distance L _2, the case of m = 1 Set to a value shorter than the minimum Talbot interference distance. That is, the distance L ₂ is set to a range satisfying the following equation (4).

In the above configuration, the G1 image generated by the first grid 21 is intensity-modulated by overlaying with the second grid 22 to generate a second periodic pattern image (hereinafter referred to as G2 image). This G2 image is picked up by the X-ray image detector 20. If there is a slight difference between the pattern period of the G1 image at the position of the second grid 22 and the arrangement pitch p2 of the _second grid 22 due to an arrangement error or the like, moire fringes occur in the G2 image. Even if this occurs, there is no particular problem in obtaining an intensity modulation signal, which will be described later, if the period of moire fringes in the x direction is different from the size of the X-ray light receiving region of the pixel 40.

  When the subject H is disposed between the X-ray source 11 and the first grid 21, the captured image of the X-ray image detector 20 is modulated by the subject H. This modulation amount reflects the angle of the X-ray deflected by the refraction effect by the subject H.

  Next, the fringe scanning method will be described. In the figure, one path of X-rays refracted according to the phase shift distribution Φ (x) in the x direction of the subject H is illustrated. Reference numeral 50 indicates a path along which the X-ray goes straight when the subject H does not exist. X-rays traveling along the path 50 pass through the first and second grids 21 and 22 and enter the X-ray image detector 20. Reference numeral 51 denotes an X-ray path refracted by the subject H when the subject H exists. X-rays traveling along the path 51 pass through the first grid 21 and are then absorbed by the X-ray absorbing portion 22 a of the second grid 22.

  The phase shift distribution Φ (x) of the subject H is expressed by the following equation (5), where n (x, z) is the refractive index distribution of the subject H. Here, the y-coordinate is omitted for simplification of description.

  The G1 image projected on the position of the second grid 22 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. This displacement amount Δx is approximately expressed by the following equation (6) based on the fact that the X-ray refraction angle φ is very small.

  Here, the refraction angle φ is expressed by the following equation (7) using the X-ray wavelength λ and the phase shift distribution Φ (x).

  Thus, the displacement amount Δx is related to the phase shift distribution Φ (x) of the subject H. Further, the displacement amount Δx and the refraction angle φ are related to the phase shift amount ψ of the intensity modulation signal of each pixel 40 detected by the X-ray image detector 20 as shown in the following equation (8). Here, the intensity modulation signal is a waveform signal representing an intensity change of pixel data with respect to a relative position of the first and second grids 21 and 22.

  Therefore, by obtaining the phase shift amount ψ of the intensity modulation signal of each pixel 40, the refraction angle φ is obtained from the equation (8), and the differential image of the phase shift distribution Φ (x) is obtained using the equation (7).

In the fringe scanning method, each time a predetermined amount of translation is performed while translating one of the first and second grids 21 and 22 relative to the other in the x direction in order to obtain the intensity modulation signal. Take a picture. In the present embodiment, the first grid 21 is fixed, and the second grid 22 is moved in the x direction by the scanning mechanism 23. The moire fringes move with the movement of the second grid 22 and return to the original position when the translational movement distance reaches the lattice period (arrangement pitch p ₂ ) of the second grid 22.

FIG. 5 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 a scanning pitch, and the second grid 22 is translated in units of the scanning pitch. Show. The scanning mechanism 23 sequentially moves the second grid 22 to M relative positions of k = 0, 1, 2,..., M−1.

  At the position of k = 0, X-rays that are not refracted by the subject H mainly pass through the second grid 22. When the second grid 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 grid 22. The component refracted by the subject H increases. In particular, at the position of k = M / 2, the X-rays passing through the second grid 22 are almost only components refracted by the subject H. When the position exceeds k = M / 2, the component of the X-ray passing through the second grid 22 refracted by the subject H decreases, while the component not refracted by the subject H increases.

  When imaging is performed by the X-ray image detector 20 at each relative position of k = 0, 1, 2,..., M−1, M pixel data for each pixel 40 is obtained. This set of M pixel data constitutes an intensity modulation signal.

In the present embodiment, as shown in the flowchart of FIG. 6, first, the second grid 22 is set at the initial position of k = 0 (step S10), and high energy imaging (step S11) and low energy imaging (step S12). ) In order. Next, it is determined whether or not k = M−1 (step S13). If k = M−1 (step S13; NO), the second grid 22 is moved by the scanning pitch (p ₂ / M) (step S14), and 1 is added to k (step S15). Energy imaging (step S11) and low energy imaging (step S12) are sequentially performed. If the above operation is repeated and k = M−1 (step S13; YES), the second grid 22 is returned to the initial position and the operation is terminated. In this way, the first and second image data are acquired by one scan, and an intensity modulation signal by high energy imaging and an intensity modulation signal by low energy imaging are obtained for each pixel 40.

Hereinafter, a method of calculating the phase shift amount ψ of each intensity modulation signal will be described. The pixel data I _k (x) at the relative position k is generally expressed by the following equation (9).

Here, A ₀ is the intensity of the incident X-ray, and _An is a value related to the amplitude value of the intensity modulation signal. N is a positive integer and i is an imaginary number. Further, φ (x) represents the refraction angle φ as a function of x.

When the scan pitch is constant and equally divided arrangement pitch p _2, the following equation (10) holds.

  When equation (10) is applied to equation (9), the refraction angle φ (x) is expressed by the following equation (11).

Here, arg [] means extraction of the declination, and corresponds to the phase shift amount ψ (x) of the intensity modulation signal I _k (x) as shown in the following equation (12).

  Although the above description does not consider the y coordinate, a two-dimensional distribution ψ (x, y) of the phase shift amount can be obtained by considering the y coordinate. This two-dimensional distribution ψ (x, y) is a phase differential image, and is represented by the following equation (13) using a trigonometric function.

FIG. 7 illustrates the intensity modulation signal I _k (x, y) obtained at each pixel 40 in high energy imaging or low energy imaging. A broken line is an intensity modulation signal assumed when the subject H is not arranged. A solid line illustrates an intensity modulation signal obtained by high energy imaging or low energy imaging when the subject H is arranged. The phase shift amount ψ (x, y) is calculated by the above equation (13).

  Next, the operation of the X-ray imaging system 10 will be described. When the subject H is disposed between the X-ray source 11 and the first grid 21 and an imaging instruction is input from the input device of the console 17, the scanning mechanism 23 causes the second grid 22 to move to the relative positions k. The high energy imaging and the low energy imaging are sequentially performed at each relative position k, and the X-ray image detector 20 generates the first and second image data. The first and second image data are sequentially stored in the memory 13.

  When the movement of the second grid 22 is finished up to the relative position of k = M−1, the second grid 22 is returned to the initial position of k = 0, and the first and the first from the memory 13 by the image processing unit 14. Each of the two pieces of image data is read out to the image processing unit 14. In the image processing unit 14, the phase differential image generation unit 30 generates a first phase differential image from the plurality of first image data based on the above equation (13), and the first phase differential image is generated from the plurality of second image data. Two phase differential images are generated and stored in the first image storage unit 31 and the second image storage unit 32, respectively.

  And the 1st and 2nd phase differential image memorize | stored in the 1st and 2nd image memory | storage parts 31 and 32 is read by the image subtraction process part 33, and it is 2nd from the image signal of a 1st phase differential image. The image signal of the phase differential image is subtracted to generate a difference image. This difference image is input to the phase contrast image generation unit 34, and an integration process along the x direction is performed to generate a phase contrast image. The phase contrast image is recorded in the image recording unit 15 and then displayed on the monitor in the console 17.

  As described above, in the energy subtraction imaging according to the present embodiment, the relative scanning of the second grid 22 with respect to the first grid 21 is only once, and the total imaging time is shortened. It can be suppressed. Thereby, a favorable phase contrast image in which artifacts due to body movement are reduced is obtained. Further, in this embodiment, every time the relative position is changed, high energy imaging and low energy imaging are performed at each relative position, so that the relative position displacement between low energy imaging and high energy imaging is prevented. There is also.

  In this embodiment, high-energy imaging and low-energy imaging are performed in this order at each relative position. This is because X-ray images are obtained by low-energy imaging in order to increase the S / N of the diagnostic image. This is because the amount of X-rays reaching the detector 20 is large and the amount of afterimage generation is large. By performing high energy imaging with a small afterimage generation amount, the influence of afterimage can be reduced.

(Other embodiments)
In the following, other embodiments of the present invention will be described. In the following embodiments, the same reference numerals are used for the same configurations as those already described, and detailed description thereof is omitted.

  In the first embodiment, shooting is performed in the order of high energy shooting and low energy shooting at each relative position, but conversely, shooting may be performed in the order of low energy shooting and high energy shooting. In order to suppress the exposure of the subject H, the second phase differential image obtained by high energy imaging is used as a diagnostic image (high X-ray dose), and the first phase differential image obtained by low energy imaging is obstructed. This is suitable for a shadow (low X-ray dose).

  In the first embodiment, in order to reduce the influence of afterimages between high energy imaging and low energy imaging at each relative position, obstacle shadows (low X-ray dose), diagnostic images ( In order to prevent the effects of afterimages between adjacent relative positions, the diagnostic images (high X dose) and obstructions are detected at each relative position, as opposed to the above. You may image | photograph in order of a shadow (low X dose). By preventing the influence of afterimages between adjacent relative positions, the amplitude of the intensity modulation signal is increased, and the calculation accuracy of the phase differential image is improved.

  In the first embodiment, the X-ray energy is changed by the tube voltage of the X-ray source 11 in order to perform high-energy imaging and low-energy imaging, but a plurality of additions are added to the collimator of the X-ray source 11. The energy of the X-ray may be changed by providing a filter that can be changed and switching the type of the additional filter.

  Further, in the first embodiment, the first and second grids 21 and 22 are provided between the X-ray source 11 and the X-ray image detector 20, but as shown in FIG. A radiation source grid 60 may be provided on the emission side of the source 11. The radiation source grid 60 includes X-ray absorption units 60a and X-ray transmission units 60b arranged alternately along the x direction, and reduces the effective focal spot size in the x direction and a large number in the x direction. A point light source (dispersed light source) is formed. Therefore, the position of the source grid in the z direction becomes the X-ray focal position. The X-ray source 11 may include the source grid 60.

  In the first embodiment, the first and second grids 21 and 22 are configured to linearly project the X-rays that have passed through the X-ray transmission portions 21b and 22b. It is not limited to this configuration, but may be a configuration described in Japanese Patent No. 4445397 in which the Talbot interference effect is generated by diffracting X-rays at the X-ray transmission part. In this case, the distance between the first and second grids is limited to the Talbot interference distance, but a phase type can be used as the first grid instead of the absorption type. In order to make the first grid a phase type, the thickness and the material are set so that a phase difference of “π” or “π / 2” is generated in the X-ray between the X-ray transmitting portion and the X-ray transmitting portion. You only have to set it. The phase-type first grid forms a self-image generated by the Talbot interference effect at the position of the second grid.

  In the first embodiment, the subject H is disposed between the X-ray source 11 and the first grid 21. However, the subject H is disposed between the first grid 21 and the second grid 22. You may arrange in. In this case as well, a phase contrast image can be similarly generated.

  In the first embodiment, the phase contrast image formed by integrating the difference image is stored in the image recording unit 15 and displayed on the monitor, but only the difference image or the difference image and the phase contrast image are displayed. May be stored in the image recording unit 15 and displayed on a monitor.

  In the first embodiment, the acquisition of correction data in a state where the subject H is not placed is not mentioned, but the subject H is not placed between the X-ray source 11 and the X-ray image detector 20. In this state, the same imaging as described above (hereinafter, imaging in a state where the subject H is not disposed is referred to as pre-imaging), and correction data may be acquired. Hereinafter, an embodiment of the present invention that enables pre-photographing will be described.

  In the present embodiment, an image processing unit 70 shown in FIG. 9 is used. The image processing unit 70 has a first correction data storage unit 71, a second correction data storage unit 72, a first correction processing unit 73, and a second correction processing unit 74 added to the image processing unit 14 of the first embodiment. Is.

  In the present embodiment, when a pre-imaging instruction is input from the input device of the console 17 in a state where the subject H is not disposed between the X-ray source 11 and the first grid 21, the flowchart of FIG. Accordingly, high energy imaging and low energy imaging are performed at each relative position of the second grid 22 with respect to the first grid 21, and a plurality of first and second image data are transferred to the image processing unit 70 via the memory 13. Entered. The phase differential image generation unit 30 generates first and second phase differential images based on the first and second image data, and stores the first correction data as the first correction data as the first phase differential image. The second phase differential image is input to the second correction data storage unit 72 as second correction data. In the case of pre-shooting, the operation ends here.

  Next, when the main imaging instruction is input from the input device of the console 17 in a state where the subject H is disposed between the X-ray source 11 and the first grid 21, the first and second image data are similarly obtained. Is generated and input to the image processing unit 70 via the memory 13. In this case, the phase differential image generation unit 30 generates first and second phase differential images based on the first and second image data, and inputs them to the first and second image storage units 31 and 32, respectively. .

  Then, the first correction processing unit 73 converts the image signal of the first correction data stored in the first correction data storage unit 71 from the image signal of the first phase differential image stored in the first image storage unit 31. An offset process to be subtracted is performed, and the second correction processing unit 74 uses the second phase differential image image signal stored in the second image storage unit 32 to store the second correction data stored in the second correction data storage unit 72. Offset processing for subtracting the image signal of the correction data is performed. The first and second phase differential images subjected to the offset process are subjected to a subtraction process by the image subtraction processing unit 33 to generate a difference image, as in the first embodiment. This difference image is converted into a phase contrast image by the phase contrast image generation unit 34.

  In the energy subtraction, the characteristics of the first and second grids 21 and 22 and the characteristics of the X-ray image detector 20 differ depending on the energy of the X-rays. Therefore, as described above, It is preferable to individually acquire the second correction data for low energy imaging.

  In the first embodiment, the difference between the first and second phase differential images obtained in the high energy imaging and the low energy imaging is taken, and a phase contrast image is generated by integrating the difference image. However, the first and second phase contrast images are respectively integrated by generating the first and second phase contrast images without taking the difference at the phase differential image stage, and the first and second phase contrast images are generated. A phase contrast image may be generated by taking a difference between contrast images. The difference processing in this case is not limited to simple difference processing, and the difference processing is performed after applying intensity weighting, filter processing, etc. to at least one pixel signal of the first and second phase contrast images. May be.

  In the first embodiment, the present invention has been described by taking an example of performing energy subtraction imaging with two types of energy, high energy and low energy. However, the present invention uses three or more types of energy. The present invention is also applicable when performing subtraction photography.

  The embodiment described above can be applied not only to a radiographic imaging system for medical diagnosis but also to other radiographic systems such as industrial use and nondestructive inspection. Furthermore, in the present invention, it is possible to use gamma rays or the like in addition to X-rays as radiation.

DESCRIPTION OF SYMBOLS 10 X-ray imaging system 11 X-ray source 21 1st grid 21a X-ray absorption part 21b X-ray transmission collection part 22 2nd grid 22a X-ray absorption part 22b X-ray transmission collection part

Claims (10)

A first grid for generating a first periodic pattern image by passing radiation emitted from a radiation source;
A second grid that partially passes the periodic pattern image of the first image to generate a second periodic pattern image;
Scanning means for changing a relative position of the second grid with respect to the first grid by a predetermined amount;
Control means for switching the energy of the radiation emitted from the radiation source at each relative position to a plurality of types of energy;
A radiation image detector for detecting the second periodic pattern image corresponding to each energy at each relative position and generating image data corresponding to each energy;
Phase image generating means for generating a phase image corresponding to each energy based on the image data corresponding to each energy;
Differential image generation means for generating a differential image by performing differential processing on the plurality of types of phase images;
A radiation imaging system comprising:   The control means switches the energy of the radiation radiated from the radiation source at each relative position to two types of energy: a first energy and a second energy having a lower energy than the first energy. The radiation imaging system according to claim 1.   The radiographic system according to claim 2, wherein the control means switches radiation energy in the order of first energy and second energy at each relative position.   The radiation imaging system according to claim 1, wherein the phase image is a phase differential image.   The radiation imaging system according to claim 1, wherein the phase image is a phase contrast image obtained by integrating a phase differential image. Correction data storage means for storing, as correction data, a phase image for each energy obtained in a state in which no subject is arranged;
Correction processing means for correcting a phase image for each energy obtained in a state in which the subject is arranged by correction data stored in the correction data storage means;
The radiation imaging system according to claim 1, further comprising:   The radiation imaging system according to claim 1, wherein the first grid is an absorption type grating, and the first periodic pattern image is formed as a projection image.   7. The radiation imaging system according to claim 1, wherein the first grid is a phase-type grating, and forms the first periodic pattern image as a self-image by a Talbot interference effect.   The radiation imaging system according to any one of claims 1 to 8, further comprising a radiation source having a radiation source grid that generates a large number of point light sources by selectively shielding radiation. A first grid that generates a first periodic pattern image by passing radiation emitted from a radiation source, and a second periodic pattern image that partially passes the periodic pattern image of the first image. A second grid, scanning means for changing the relative position of the second grid with respect to the first grid by a predetermined amount, the second periodic pattern image is detected at each relative position, and image data is respectively A radiography system control method comprising a radiological image detector to be generated, comprising:
The energy of the radiation radiated from the radiation source at each relative position is switched to a plurality of types of energy, and the second periodic pattern image corresponding to the energy at each relative position is changed by the radiation image detector. By detecting, the image data corresponding to each energy is generated,
A control method comprising: generating a difference image by generating a phase image corresponding to each energy and performing a difference process based on the image data corresponding to each energy.

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