JP2011206162A - Radiographic system and method - Google Patents

Abstract

PROBLEM TO BE SOLVED: To prevent noise due to a marker indicating the position of a subject when capturing a phase contrast image using a fringe scanning method.SOLUTION: A subject H to which a marker 11 is attached is arranged between an X-ray source and an FPD 21. On the basis of a distance L4 in a z direction from an X-ray focus 12a to the edge part Ha of the subject H, a distance L5 in the z direction from the X-ray focus 12a to the edge part 11a of the marker 11, a distance L6 in the z direction from the edge part 11a of the marker 11 to the FPD 21, a distance T1 in an x direction from one end of the X-ray focus 12a to the edge part Ha of the subject H, and the width w of the X-ray focus 12a, a distance T for preventing the image of the subject H detected by the FPD 21 and the image of the marker 11 from overlapping is calculated, and the position of the marker 11 is adjusted such that an interval between the edge part Ha of the subject H and the edge part 11a of the marker 11 is equal to or longer than the distance T.

Description

  The present invention relates to a radiation imaging system and method for imaging a subject with radiation such as X-rays. More specifically, in order to correct displacement of a subject image due to body movement of the subject, a marker is attached to the subject. The present invention relates to a radiation imaging system and method for performing imaging.

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

  In a general X-ray imaging system, a subject is placed between an X-ray source that emits X-rays and an X-ray image detector that detects X-rays, and a transmission image of the subject is taken. In this case, 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 each pixel of the X-ray image detector. As a result, an X-ray absorption image of the subject is detected and imaged by the X-ray image detector. As an X-ray image detector, in addition to a combination of an X-ray intensifying screen and a film and a stimulable phosphor, a flat panel detector (FPD) using a semiconductor circuit is widely used.

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

  Against this background, in recent years, instead of X-ray intensity changes by the subject, an image based on the X-ray phase change (angle change) by the subject (hereinafter referred to as a phase contrast image) is obtained. Research on line phase imaging has been actively conducted. In general, when X-rays are incident on an object, the phase exhibits a higher interaction than the intensity of the X-rays. Therefore, X-ray phase imaging using a phase difference is a weakly absorbing object with low X-ray absorption capability. Can also obtain a high-contrast image. As a kind of such X-ray phase imaging, an X-ray imaging system using an X-ray Talbot interferometer composed of two transmission diffraction gratings and an X-ray image detector has been devised (for example, Patent Document 1). reference).

  The X-ray Talbot interferometer has a first diffraction grating disposed behind the subject, and a second diffraction grating downstream by a specific distance (Talbot interference distance) determined by the grating pitch of the first diffraction grating and the X-ray wavelength. Are arranged, and an X-ray image detector is arranged behind them. The Talbot interference distance is a distance at which X-rays that have passed through the first diffraction grating form a self-image (stripe image) due to the Talbot interference effect. This self-image is the same as the X-ray source and the first diffraction grating. Are modulated by the interaction (phase change) between the subject and the X-rays arranged between the two.

  In the X-ray Talbot interferometer, a fringe image whose intensity is modulated by superimposing the self-image of the first diffraction grating and the second diffraction grating is detected by a fringe scanning method, and a change in the fringe image (phase shift by the subject) is detected. ) To obtain a phase contrast image of the subject. In the fringe scanning method, as in the X-ray imaging system 60 shown in FIG. 10 (A), the second diffraction grating 62 is substantially placed on the surface of the first diffraction grating 61 with respect to the first diffraction grating 61. X-rays from the X-ray source 63 to the subject H are translated in parallel at a scanning pitch obtained by equally dividing the grating pitch in a direction substantially parallel to the grating direction (strip direction) of the first diffraction grating 61. A plurality of times of irradiation and imaging are performed and detected by the FPD 64, and a phase differential image (a subject to be detected) is obtained from a phase shift amount (phase shift amount with and without the subject H) of pixel data of each pixel 65. (Corresponding to the angular distribution of X-rays refracted by the specimen), and a phase contrast image of the subject is obtained by integrating this phase differential image along the above-described fringe scanning direction. Since the pixel data is a signal whose intensity is periodically modulated with respect to the scanning pitch, it may be hereinafter referred to as an “intensity modulated signal”.

  In the fringe scanning method, imaging is performed a plurality of times while moving the second diffraction grating 62. Therefore, if the subject H moves during imaging, the phase information cannot be calculated correctly. For example, as in the X-ray imaging system 60 shown in FIG. 10A, X-rays transmitted through the imaging region 67 of the subject H are detected by the pixels 65a of the FPD 64, as shown in FIG. In addition, if the subject H moves upward during imaging, a different imaging region 68 of the subject H is imaged at the pixel 65a, so that the phase differential image obtained from the same pixel 65a is different from that of the imaging region. turn into. As for the body movement of the human body, for example, when the body movement was measured with the knee fixed, it was found that body movement of about 100 μm occurred per second. Since the realistic pixel size of the FPD is 150 to 300 μm, in the fringe scanning method that takes several seconds to shoot, body movement for several pixels occurs.

  Conventionally, in an apparatus that takes a phase contrast image by performing a plurality of measurements while moving the radiation detector in the optical axis direction, the subject position is determined from each measurement image, and thus a subject is attached with a marker for imaging. (For example, refer to Patent Document 2).

JP 2008-200399 A JP 2002-336229 A

  In the fringe scanning method, a phase contrast image is obtained by integrating a phase differential image calculated from each pixel data of the FPD 64 in a direction substantially perpendicular to the lattice direction (strip direction). Therefore, if there is a marker that is a substance different from the subject in the integration range, the phase information of the marker becomes noise, so extra processing is required to exclude the marker phase information from the phase contrast image. Become. Further, since the marker exists outside the subject, the X-ray is refracted at the edge portion of the marker, which is the boundary between the marker and air. Since the refraction at the edge portion of the marker is larger than the refraction between the body fluid and the soft tissue in the body of the subject, a large noise occurs in the phase contrast image unless the phase information of the marker is properly excluded.

  An object of the present invention has been made in view of the above problems, and is to prevent noise from being generated by a marker when a phase contrast image is captured using a fringe scanning method.

  In order to achieve the above object, the radiation imaging system of the present invention includes a radiation source for irradiating radiation, a first grating, intensity modulation means, a marker, and a radiation image detector. The first grating generates a fringe image by passing radiation. The intensity modulation means applies intensity modulation to the fringe image at a plurality of relative positions having different phases with respect to the periodic pattern of the fringe image. The marker is attached to a subject arranged between the radiation source and the first grating, or between the first grating and the intensity modulating means. The radiation image detector detects a plurality of scanning images including the image of the subject and the image of the marker that have been intensity-modulated by the intensity modulation means at each relative position. In this radiation imaging system, the marker is disposed at a position away from the subject by a predetermined distance or more so that the marker image detected by the radiation image detector does not overlap the subject image.

The predetermined distance is a distance T between the edge portions of the marker and the subject that are close to each other in the arrangement direction of the marker with respect to the subject, and the distance from the radiation focus in the optical axis direction of the radiation source to the edge portion of the subject. L ₄ , the distance from the radiation focus to the edge of the marker is L ₅ , the distance from the edge of the marker to the radiation image detector is L _6, and the width of the radiation focus in the marker placement direction with respect to the subject is w, the distance from one end of the radiation focus to the edge portion of the subject in the case of a T _1, it is preferable to satisfy the following equation.
T ≧ {((L ₅ / L ₄ ) −1) × T ₁ } + {(L ₆ / (L ₅ + L ₆ )) × w}

  In addition, a correction processing unit that detects a marker image from each of the plurality of scanning images and corrects the position of each scanning image so that the marker images coincide with each other, and an intensity for each pixel based on the subject image of each scanning image. Phase differential image generating means for calculating the phase shift amount of the modulation signal and generating a phase differential image of the subject based on the phase differential value corresponding to the phase shift amount, and including a marker image based on the position of the marker image A phase contrast image generation unit that determines a non-integral range and integrates the phase differential image within the integration range generated by the phase differential image generation unit to generate a phase contrast image.

  The intensity modulation means may be composed of a second grating having a periodic pattern in the same direction as the fringe image, and scanning means for moving one of the first and second gratings at a predetermined pitch.

  The first and second gratings may be absorption gratings, and the first grating may project radiation from the radiation source onto the second grating as a fringe image.

  Alternatively, the first grating may be a phase grating, and the radiation from the radiation source may be projected as a fringe image onto the second grating by the Talbot interference effect of the first grating.

  The radiation image detector is a radiation image detector provided for each pixel with a conversion layer that converts radiation into charges and a charge collection electrode that collects charges converted in the conversion layer. It is preferable that a plurality of linear electrode groups having a periodic pattern in the same direction as the fringe image are arranged so that the phases thereof are different from each other, and the intensity modulation means is constituted by charge collecting electrodes.

  In the radiation imaging method of the present invention, a fringe image is generated by passing radiation through a first grating, and intensity modulation is performed on the fringe image by intensity modulation means at a plurality of relative positions whose phases are different from the periodic pattern of the fringe image. A plurality of scanned images including an image of a subject disposed between the radiation source and the first grating or between the first grating and the intensity modulating means and a marker attached to the subject; Are detected by the radiation image detector at each relative position. At that time, the marker is arranged at a position away from the subject by a predetermined distance or more so that the marker image detected by the radiation image detector and the subject image do not overlap.

  When the distance T between the adjacent edge portions of the marker and the subject in the arrangement direction of the marker with respect to the subject is obtained as the predetermined distance, it is preferable to use the above-described mathematical formula of the radiation imaging system.

  Further, a step of detecting a marker image from each of the plurality of scanning images and correcting the position of each scanning image so that the marker images coincide with each other, and an intensity modulation signal for each pixel based on the subject image of each scanning image. A phase shift amount of the subject, a step of generating a phase differential image of the subject based on the phase differential value corresponding to the phase shift amount, and an integration range not including the marker image based on the position of the marker image is determined. Integrating a phase differential image within the integration range to generate a phase contrast image.

  According to the present invention, since the marker is separated from the subject by a predetermined distance or more so that the subject image and the marker image do not overlap, the subject is not subjected to extra processing such as removing phase information of the marker. The specimen image and the marker image can be separated. In addition, since the integration range is determined so as not to include the marker image, a phase contrast image based only on the subject image can be obtained. Furthermore, since the distance between the subject and the marker can be calculated based on the distance from the radiation focus to the subject and the edge portion of the marker, the marker can be placed at an appropriate position.

It is a schematic diagram which shows the structure of the X-ray imaging system of this invention. It is a schematic diagram which shows the structure of a flat panel detector. It is a schematic side view which shows the structure of the 1st and 2nd absorption type grating | lattice. It is explanatory drawing for demonstrating a fringe scanning method. It is a graph which illustrates pixel data (intensity modulation signal) which changes with fringe scanning. It is the schematic diagram showing the distance which arrange | positions a marker away from a subject. It is a flowchart which shows the procedure of radiography. It is explanatory drawing which shows several image data obtained by fringe scanning imaging | photography. It is a schematic diagram which shows the structure of the X-ray detector used by 2nd Embodiment of this invention. It is a schematic diagram which shows the structure of the X-ray imaging system using the conventional fringe scanning imaging | photography.

(First embodiment)
1, an X-ray imaging system 10 according to the first embodiment of the present invention includes a marker 11 attached to a subject H, an X-ray source 12 that irradiates the subject H with X-rays, and an X-ray source 12. An imaging unit 13 that detects X-rays transmitted through the subject H from the X-ray source 12 and generates image data, a memory 14 that stores image data read from the imaging unit 13, and a memory 14, an image processing unit 15 that generates a phase contrast image by image processing the plurality of image data stored in the image processing unit 14, an image recording unit 16 that records the phase contrast image generated by the image processing unit 15, and an X-ray source 12 and the imaging control unit 17 that controls the imaging unit 13, a console 18 including an operation unit and a monitor, and an overall control of the X-ray imaging system 10 based on an operation signal input from the console 18. And a system controller 19 for.

  The marker 11 is imaged together with the subject H, and is used as a reference when correcting the body movement of the subject H in the image data read from the imaging unit 13. The marker 11 has, for example, a spherical shape with a diameter of several millimeters to several tens of millimeters, and is formed using a material having high X-ray absorption, such as lead. The marker 11 is held by an attachment member made of a material having high X-ray transparency made of, for example, transparent plastic, and the attachment member takes an image of a joint portion such as a knee, an elbow, and a finger of the subject H. It is fixed to the subject H by adhesion, adhesion, or affixing to the site. The attachment member preferably has a structure that can move the position of the marker 11 so that the position of the marker 11 with respect to the subject H can be adjusted.

  The X-ray source 12 includes a high voltage generator, an X-ray tube, a collimator (all not shown) and the like, and irradiates the subject H with X-rays based on the control of the imaging control unit 17. For example, the X-ray tube is of a rotating anode type, emits an electron beam from a filament in accordance with a voltage from a high voltage generator, and collides the electron beam with a rotating anode rotating at a predetermined speed, thereby causing an X-ray. Is generated. The rotating anode rotates in order to reduce deterioration due to the electron beam continuing to hit the fixed position, and the collision part of the electron beam becomes an X-ray focal point 12a (see FIG. 3) that emits X-rays. . The collimator limits the irradiation field so as to shield a portion of the X-ray emitted from the X-ray tube that does not contribute to the examination region of the subject H.

  The imaging unit 13 includes a flat panel detector (FPD) 21 made of a semiconductor circuit, a first absorption grating 22 for detecting phase change (angle change) of X-rays by the subject H, and performing phase imaging. Two absorption gratings 23 are provided. The FPD 21 is arranged so that the detection surface is orthogonal to a direction along the optical axis A of the X-rays irradiated from the X-ray source 12 (hereinafter referred to as z direction).

The first absorption type grating 22 has a plurality of X-ray shielding portions 22a extending in one direction in the plane orthogonal to the z direction (hereinafter referred to as the y direction), in a direction orthogonal to the z direction and the y direction (hereinafter referred to as the y direction). the x that direction) in which are arranged at a predetermined pitch p _1. Similarly, the second absorption type grating 23, a plurality of X-ray shielding portion 23a which extend in the y direction, in which are arranged at a predetermined pitch p ₂ in the x-direction. As a material of the X-ray shielding portions 22a and 23a, a metal excellent in X-ray absorption is preferable, and for example, gold, platinum, lead, tungsten and the like are preferable.

  The imaging unit 13 translates the second absorption type grating 23 in a direction (x direction) perpendicular to the grating direction, so that the relative position of the first absorption type grating 22 to the second absorption type grating 23 is changed. A scanning mechanism 25 is provided for changing the above. The scanning mechanism 25 is configured by an actuator such as a piezoelectric element, for example. The scanning mechanism 25 is driven based on the control of the imaging control unit 17 during the stripe scanning described later. As will be described in detail later, the memory 14 stores image data obtained by the photographing unit 13 at each scanning step of fringe scanning. The second absorption type grating 23 and the scanning mechanism 25 constitute the intensity modulation means described in the claims.

  The image processing unit 15 includes a correction processing unit 27, a phase differential image generation unit 29, and a phase contrast image generation unit 30. The correction processing unit 27 detects a marker image from a plurality of image data (scanned images) photographed by the photographing unit 13 and stored in the memory 14 at each scanning step of fringe scanning, and the positions of the detected marker images match. Thus, by correcting a plurality of image data, the positions of the subject images of the respective image data are matched.

  The phase differential image generation unit 29 generates a phase differential image based on the plurality of image data corrected by the correction processing unit 27. The phase contrast image generation unit 30 generates a phase contrast image by integrating the phase differential image along the x direction. The phase contrast image generated by the phase contrast image generation unit 30 is recorded in the image recording unit 16, and then output to the console 18 and displayed on a monitor (not shown).

  In addition to the monitor, the console 18 includes an input device (not shown) through which an operator inputs a shooting instruction and the content of the instruction. As this input device, for example, a switch, a touch panel, a mouse, a keyboard or the like is used, and an X-ray imaging condition such as a tube voltage of the X-ray tube or an X-ray irradiation time, an imaging timing, etc. are input by operation of the input device. The The monitor is composed of a liquid crystal display or a CRT display, and displays characters such as X-ray imaging conditions and the phase contrast image.

  In FIG. 2, the FPD 21 includes an image receiving unit 36 in which a plurality of pixels 35 that convert X-rays into electric charges and store them two-dimensionally on the active matrix substrate along the x and y directions, and the image receiving unit 36. The scanning circuit 37 controls the readout timing of the charges from the pixel, and the readout circuit 38 that reads out the charges accumulated in each pixel 35, converts the charges into image data, and stores them. The scanning circuit 37 and each pixel 35 are connected by a scanning line 39 for each row, and the reading circuit 38 and each pixel 35 are connected by a signal line 40 for each column. The arrangement pitch of the pixels 35 is about 100 μm in each of the x direction and the y direction.

  The pixel 35 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. It can be configured as a conversion type X-ray detection element. A TFT switch (not shown) is connected to each pixel 35, a gate electrode of the TFT switch is connected to the scanning line 39, a source electrode is connected to the capacitor, and a drain electrode is connected to the signal line 40. When the TFT switch is turned on by the drive pulse from the scanning circuit 37, the charge accumulated in the capacitor is read out to the signal line 40.

Note that the pixel 35 temporarily converts X-rays into visible light with a scintillator (not shown) made of gadolinium oxide (Gd ₂ O ₃ ), cesium iodide (CsI), or the like, and converts the converted visible light into a photodiode ( It is also possible to configure as an indirect conversion type X-ray detection element that converts into electric charge and accumulates it (not shown). In this embodiment, an FPD based on a TFT panel is used as a radiation image detector. However, the present invention is not limited to this, and various types of radiation image detectors based on a solid-state imaging device such as a CCD sensor or a CMOS sensor. It is also possible to use.

  The read circuit 38 includes an integration amplifier, a correction circuit, an A / D converter (none of which are shown), and the like. The integrating amplifier integrates the charges output from each pixel 35 via the signal line 40 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 offset correction, gain correction, linearity correction, and the like on the image data, and inputs the corrected image data to the memory 14.

In FIG. 3, X-ray shielding portions 22a of the first absorption grating 22, at a predetermined pitch p ₁ in the x direction, and are arranged at a predetermined interval d ₁ from each other. Similarly, X-rays shielding portions 23a of the second absorption-type grating 23, at a predetermined pitch p ₂ in the x direction, and are arranged at a predetermined interval d ₂ from each other. The first and second absorption gratings 22 and 23 do not give a phase difference to incident X-rays but give an intensity difference, and are also called amplitude gratings. Incidentally, (a region of the interval d _1, d ₂₎ slit portion may not be a void, X-rays low absorption material such as a polymer or light metal may be filled.

The first and second absorption type gratings 22 and 23 are configured to linearly project the X-rays that have passed through the slit portion regardless of the presence or absence of the Talbot interference effect. Specifically, by setting the distances d ₁ and d ₂ to a value sufficiently larger than the peak wavelength of X-rays emitted from the X-ray source 12, most of the X-rays included in the irradiated X-rays are slit at the slit portion. It is configured to pass through without being diffracted while maintaining straightness. For example, when tungsten is used as the rotating anode of the aforementioned X-ray tube and the tube voltage is 50 kV, the peak wavelength of the X-ray is about 0.4 mm. In this case, if the distances d ₁ and d ₂ are about ₁ to 10 μm, most of the X-rays are projected linearly without being diffracted by the slit portion. In this case, the lattice pitches p ₁ and p ₂ are about ₂ to 20 μm.

Since the X-rays emitted from the X-ray source 12 are not parallel beams but cone beams having an X-ray focal point as a light emission point, a projected image projected through the first absorption type grating 22 (hereinafter, referred to as a projection image) This projection image is referred to as a G1 image or a fringe image) and is enlarged in proportion to the distance from the X-ray focal point 12a. The grating pitch p ₂ and the interval d ₂ of the second absorption type grating 23 are determined so that the slit portions thereof substantially coincide with the periodic pattern of the bright part of the G1 image at the position of the second absorption type grating 23. Yes. That is, when the distance from the X-ray focal point 12a to the first absorption-type grating 22 is L ₁ and the distance from the first absorption-type grating 22 to the second absorption-type grating 23 is L ₂ , the grating pitch p ₂ and the distance d ₂ are determined so as to satisfy the relationship of the following expressions (1) and (2).

In the case of a Talbot interferometer, the distance L ₂ from the first absorption type grating 22 to the second absorption type grating 23 is limited to the Talbot interference distance determined by the grating pitch of the first diffraction grating and the X-ray wavelength. However, in the imaging unit 13 of the present embodiment, the first absorption type grating 22 is configured to project incident X-rays without diffracting, and the G1 image of the first absorption type grating 22 is the first one. because at every position of the rear absorption gratings 22 of similarly obtained, the distance L _2, can be set independently of the Talbot distance.

As described above, the imaging unit 13 of the present embodiment does not constitute a Talbot interferometer, but it is assumed that X-ray diffraction occurs in the first absorption grating 22 and a Talbot interference effect is generated. The Talbot interference distance Z is expressed by the following equation (3) using the grating pitch p _{1 of} the first absorption grating 22, the X-ray wavelength (peak wavelength) λ, and a positive integer m.

In the present embodiment, as described above, the distance L ₂ can be set regardless of the Talbot interference distance. Therefore, the objective is to reduce the thickness of the photographing unit 13 and the minimum Talbot when the distance L ₂ is m = 1. A value shorter than the interference distance Z is set. That is, the distance L ₂ is set to a value in the range satisfying the following equation (4).

  The X-ray shielding portions 22a and 23a preferably shield (absorb) X-rays completely in order to generate a periodic pattern image with high contrast. However, the materials (gold, platinum) having excellent X-ray absorption properties described above are preferred. Even if lead, tungsten, etc.) are used, there are not a few X-rays that are transmitted without being absorbed. For this reason, in order to improve the shielding property of X-rays, it is preferable to make each of the X-ray shielding portions 22a and 23a (thickness in the z direction) as thick as possible (that is, increase the aspect ratio). For example, when the tube voltage of the X-ray tube is 50 kV, it is preferable to shield 90% or more of the irradiated X-rays. In this case, the thicknesses of the X-ray shielding portions 22a and 23a are converted to gold (Au). It is preferable that it is 30 micrometers or more.

In the first and second absorption type gratings 22 and 23 configured as described above, intensity modulation is performed by superimposing the G1 image (stripe image) of the first absorption type grating 22 and the second absorption type grating 23. The striped image is captured by the FPD 21. There is a slight difference between the pattern period of the G1 image at the position of the second absorption type grating 23 and the grating pitch p _{2 of} the second absorption type grating 23 due to manufacturing errors and arrangement errors. Thus, moire fringes are generated in the intensity-modulated fringe image. Further, when an error occurs in the lattice arrangement direction of the first and second absorption type gratings 21 and 22 and the arrangement directions are not the same, so-called rotational moire occurs. However, even when such moire fringes occur in the fringe image, there is no particular problem as long as the period of the moire fringes in the x direction or y direction is larger than the arrangement pitch of the pixels 40. Ideally, it is preferable not to generate moire fringes, but the moire fringes can be used to confirm the scanning amount of the fringe scanning (translation distance of the second absorption grating 22), as will be described later. .

  When the subject H is disposed between the X-ray source 12 and the first absorption type grating 22, the fringe image detected by the FPD 21 is modulated by the subject H. This modulation amount is proportional to the angle of the X-ray deflected by the refraction effect by the subject H. Therefore, a phase contrast image of the subject H can be generated by analyzing the fringe image detected by the FPD 21.

  Next, a fringe image analysis method will be described. FIG. 3 illustrates one X-ray refracted according to the phase shift distribution Φ (x) in the x direction of the subject H. Reference numeral 45 indicates an X-ray path that travels straight when the subject H does not exist. The X-ray that travels along the path 45 passes through the first and second absorption gratings 22 and 23 to the FPD 21. Incident. Reference numeral 46 indicates an X-ray path that is refracted and deflected by the subject H when the subject H exists. X-rays traveling along the path 46 pass through the first absorption type grating 22 and are then shielded by the X-ray shielding part 23 a of the second absorption type grating 23.

  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, and z is the direction in which the X-rays travel.

  The G1 image projected from the first absorption grating 22 to the position of the second absorption grating 23 is displaced in the x direction by an amount corresponding to the refraction angle φ due to refraction of X-rays at the subject H. It will be. This 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) of the subject H.

  Thus, the displacement amount Δx of the G1 image due to X-ray refraction at the subject H is related to the phase shift distribution Φ (x) of the subject H. This displacement amount Δx is the amount of phase shift ψ of the intensity modulation signal of each pixel 35 detected by the FPD 21 (the amount of phase shift of the intensity modulation signal of each pixel 35 with and without the subject H). ) Is related to the following equation (8).

  Accordingly, by obtaining the phase shift amount ψ of the intensity modulation signal of each pixel 35, the refraction angle φ is obtained from the equation (8), and the differential amount of the phase shift distribution Φ (x) is obtained using the equation (7). By integrating this with respect to x, the phase shift distribution Φ (x) of the subject H, that is, the phase contrast image of the subject H can be generated. In the present embodiment, the phase shift amount ψ is calculated using the fringe scanning method shown below.

In the fringe scanning method, imaging is performed while one of the first and second absorption type gratings 22 and 23 is translated in the x direction relative to the other (that is, while changing the phase of both grating periods). Take a picture). In the present embodiment, the second absorption type grating 23 is moved by the scanning mechanism 25 described above. As the second absorption type grating 23 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 23 (grating pitch p ₂ ). (Ie, when the phase change reaches 2π), the moire fringes return to their original positions. In this way, while moving the second absorption type grating 23 by an integer of the grating pitch p ₂ , a stripe image is taken with the FPD 21, and intensity modulation signals of each pixel are acquired from the taken plurality of stripe images. The phase differential image generation unit 29 in the image processing unit 15 performs arithmetic processing to obtain the phase shift amount ψ of the intensity modulation signal of each pixel. This two-dimensional distribution of the phase shift amount ψ corresponds to a phase differential image.

FIG. 4 schematically shows a state in which the second absorption type grating 23 is moved by a scanning pitch (p ₂ / M) obtained by dividing the grating pitch p ₂ into M (integers of 2 or more). The scanning mechanism 25 translates the second absorption type grating 23 in order to M scanning positions k = 0, 1, 2,..., M−1. In the figure, the initial position of the second absorption type grating 23 is set such that the dark part of the G1 image at the position of the second absorption type grating 23 when the subject H is not present is almost at the X-ray shielding part 23a. Although the matching position (k = 0) is assumed, this initial position may be any position among k = 0, 1, 2,..., M−1.

  First, at the position of k = 0, mainly X-rays that are not refracted by the subject H pass through the second absorption type grating 23. Next, when the second absorption type grating 23 is moved in order of k = 1, 2,..., The X-rays passing through the second absorption type grating 23 were not refracted by the subject H. While the X-ray component decreases, the X-ray component refracted by the subject H increases. In particular, at the position of k = M / 2, mainly only the X-rays refracted by the subject H pass through the second absorption type grating 23. When the position exceeds k = M / 2, the X-ray component passing through the second absorption type grating 23 is refracted by the subject H while the X-ray component refracted by the subject H decreases. The X-ray component that has not been increased.

  When photographing is performed by the FPD 21 at each position of k = 0, 1, 2,..., M−1, M pieces of image data are obtained. Since each image data includes pixel data of each pixel 35, M pixel data can be obtained for each pixel 35. M pieces of image data are stored in the memory 14.

  The marker 11 is imaged by the FPD 21 together with the subject H at each position of k = 0, 1, 2,..., M−1. The correction processing unit 27 reads M pieces of image data from the memory 14, and sets, for example, the first image data taken, that is, the image data taken at the position of k = 0 as the reference image data. The marker image and its position are detected.

  Next, the correction processing unit 27 detects the marker image and its position also from the image data photographed at each position of k = 1, 2,..., M−1, and the position of these marker images is the reference. Image data photographed at each position of k = 1, 2,..., M−1 is corrected so as to coincide with the position of the marker image of the image data. Even if the position of the subject H in each image data is different due to the movement of the subject H during the fringe scanning imaging, the position of the image of the subject H is corrected by correcting each image data based on the position of the marker image. Can be matched. Thereby, the phase shift amount ψ of the intensity modulation signal of each pixel 35 can be appropriately calculated from the M pixel data.

Hereinafter, a method of calculating the phase shift amount ψ of the intensity modulation signal of each pixel 35 from the M pixel data will be described. When the pixel data of each pixel 35 at the position k of the second absorption type grating 23 is denoted as I _k (x), I _k (x) is expressed by the following equation (9).

Here, x is a coordinate in the x direction of the pixel, A ₀ is the intensity of the incident X-ray, and _An is a value corresponding to the contrast of the intensity modulation signal (where n is a positive integer). is there). Φ (x) represents the refraction angle φ as a function of the coordinate x of the pixel 35.

  Next, using the relational expression of the following expression (10), the refraction angle φ (x) is expressed as the expression (11).

  Here, arg [] means extraction of the declination and corresponds to the phase shift amount ψ. Therefore, by calculating the phase shift amount ψ from the M pieces of pixel data (intensity modulation signal) obtained from each pixel 35 based on the equation (11), the refraction angle φ (x) is obtained, and the phase shift distribution is obtained. The differential amount of Φ (x) is obtained.

Specifically, as shown in FIG. 5, the M pixel data obtained in each pixel 35 is periodically with a period of the grating pitch p _{2 with} respect to the position k of the second absorption type grating 23. Change. The broken line in the figure shows the change of the pixel data when the subject H does not exist, and the solid line in the figure shows the change of the pixel data when the subject H exists. The phase difference between the two waveforms corresponds to the phase shift amount ψ.

  In the above description, the y coordinate in the y direction of the pixel 35 is not taken into consideration. However, by performing the same calculation for each y coordinate, the two-dimensional phase shift distribution ψ (x, y) is obtained. This phase shift distribution ψ (x, y) corresponds to the phase differential image. Since the refraction angle φ and the phase shift amount ψ are in a proportional relationship as shown in the above formula (7), both are physical quantities corresponding to the differential amount of the phase shift distribution Φ (x).

  The phase differential image is input to the phase contrast image generation unit 30. The phase contrast image generation unit 30 integrates the input phase differential image along the x axis to generate a phase shift distribution Φ (x, y) of the subject H, and outputs this as a phase contrast image. . The phase shift distribution Φ (x, y) of the subject H is expressed as in Expression (12).

  When the phase contrast image generator 30 integrates the phase differential image along the x-axis, if the phase differential image of the marker 11 is included in the integration range, the phase differential image of the marker 11 becomes noise. . Therefore, in this embodiment, as shown in FIG. 6, in order to prevent the image of the edge portion 11 a of the marker 11 detected by the FPD 21 from overlapping the image of the edge portion Ha of the subject H, the marker 11 is placed on the subject. It is arranged at a position away from H by a predetermined distance T or more. Hereinafter, a method for obtaining the distance T will be described.

In the rotary anode type X-ray tube used for the X-ray source 12, since a wide range on the anode target is an X-ray emission point, the X-ray focal point 12a has a length of width w. Therefore, since the subject H and the marker 11 are irradiated with X-rays from the entire region of the X-ray focal point 12a, the image of the edge portion Ha of the subject H and the edge portion 11a of the marker 11 detected by the FPD 21 includes The blurs with the widths V ₁ and V ₂ occur. Therefore, in the present embodiment, the distance T is set such that the blur V ₁ of the edge portion Ha of the subject H and the blur V ₂ of the edge portion 11 a of the marker 11 do not overlap.

In order to prevent the blur V _{1 at} the edge portion Ha of the subject H from overlapping the blur V _{2 at} the edge portion 11 a of the marker 11, the point on the marker 11 side of the blur V ₁ is the subject of A and blur V ₂ . When the point on the H side is B and the x coordinates of both points A and B are x _A and x _B , the following equation (13) may be satisfied.

Also, the distance in the z direction from the X-ray focal point 12a to the edge portion Ha of the subject H is L ₄ , the distance in the z direction from the X-ray focal point 12a to the edge portion 11a of the marker 11 is L ₅ , and the edge portion of the marker 11 When the distance in the z direction from 11 a to the FPD 21 is L ₆ and the distance in the x direction from one end of the X-ray focal point 12 a to the edge portion Ha of the subject H is T ₁ , the x coordinates x _A of the points A and B , _{x B} is the following formula (14) obtained respectively from (15).

Therefore, the condition “x _B −x _A ” for preventing the blur V ₁ and the blur V ₂ from overlapping each other is obtained by the following equation (16). Further, when the following equation (16) is converted with respect to the distance T from the subject H marker 11, the following equation (17) is obtained. Therefore, an appropriate lower limit of the distance T between the subject H and the marker 11 can be obtained based on the distances L _{4 to} L ₆ , w, and T ₁ .

  Further, the phase contrast image generation unit 30 determines the integration range F based on the detection position of the marker 11 so that the phase differential image based on the marker 11 is not included in the integration range. Thereby, the phase contrast image generation unit 30 can calculate a phase contrast image based only on the phase differential image of the subject H. Note that when the arrangement direction of the marker 11 with respect to the subject H is the y direction, that is, the lattice direction of the first absorption type grating 22 or the like, the y of the FPD 21 used for integration by the phase contrast image generation unit 30 is the same as the x direction. A directional pixel range is determined.

Next, the operation of the X-ray imaging system 10 configured as described above will be described with reference to the flowchart of FIG. A subject H to which the marker 11 is attached is disposed between the X-ray source 12 and the FPD 21. Then, the width w of the X-ray focal 12a, distance T is calculated based on each distance _L 4 _~L 6, _{T 1} from the X-ray focal point 12a to the edge portion Ha and 11a, a edge portion Ha of the subject H The distance between the marker 11 and the edge portion 11a is adjusted to be equal to or greater than the distance T (S1).

  When an imaging operation is performed from the console 18 by the operator, the respective units of the X-ray imaging system 10 operate in cooperation to perform the above-described fringe scanning imaging, and a fringe image and a marker image are simultaneously photographed (S2). A plurality of image data created by the fringe scanning photographing is stored in the memory 14.

  The correction processing unit 27 reads M image data from the memory 14. FIGS. 8A to 8D are image data 50 to 50M−1 photographed at positions k = 0, 1, 2, and M−1, for example. For example, the correction processing unit 27 uses, as reference image data, image data 50 that is first captured, that is, image data 50 that is captured at a position of k = 0, and the position of the marker image 50a on the xy coordinates from the reference image data 50. Is detected. Next, the correction processing unit 27 detects the positions of the marker images 51a, 52a to 50M-1a from the image data 51 to 50M-1 photographed at the respective positions k = 1, 2,. (S3).

  The correction processing unit 27 corrects the image data 50 to 50M-1 so that the positions of the marker images 51a and 52a to 50M-1a coincide with the position of the marker image 50a of the reference image data 50 (S4). . Thereby, the position of the stripe image of the image data 51 to 50M-1 also coincides with the position of the stripe image of the image data 50.

  The phase differential image generation unit 29 generates a phase differential image from the fringe images of the plurality of corrected image data (S5). Since the positions of the fringe images of the respective image data 50 to 50M-1 coincide with each other, the phase differential image can be appropriately calculated.

  Based on the detection position of the marker 11, the phase contrast image generation unit 30 sets the integration range F so that the phase differential image based on the marker 11 is not included. Then, the phase differential image in the integration range F is integrated in the x-axis direction to generate a phase contrast image (S6). Thereby, a phase contrast image based only on the phase differential image of the subject H is calculated. The phase contrast image is output to the console 18 and displayed on the monitor.

  In the above embodiment, when the distance from the X-ray source 12 to the FPD 21 is increased, the blur of the G1 image due to the focus size of the X-ray focal point 12a (generally about 0.1 mm to 1 mm) is affected, and the phase contrast image Therefore, a multi-slit (source grid) may be arranged immediately after the X-ray focal point 12a.

  The multi-slit is an absorption type grating having a configuration similar to that of the first and second absorption type gratings 22 and 23, and a plurality of X-ray shielding portions extending in one direction (in this embodiment, y direction) The first and second absorption gratings 22 and 23 are periodically arranged in the same direction (in the present embodiment, the x direction) as the X-ray shielding portions 22a and 23a. The multi-slit partially shields X-rays from the X-ray source 12 to reduce the effective focal size in the x direction, and forms a large number of point light sources (dispersed light sources) in the x direction. Suppresses the blur of the G1 image.

In the above-described embodiment, the first and second absorption gratings 22 and 23 are configured to linearly project X-rays that have passed through the slit portion, but the present invention is limited to this configuration. Instead of this, it is also possible to adopt a configuration in which a so-called Talbot interference effect is generated by diffracting X-rays at the slit portion (configuration described in International Publication WO 2004/058070). However, in this case, it is necessary to set the distance L ₂ between the first and second absorption gratings 22, 23 Talbot distance. In this case, a phase type grating (phase type diffraction grating) can be used instead of the first absorption type grating 22, and the phase type grating used instead of the first absorption type grating 22 can be used. Projects the fringe image (self-image) generated by the Talbot interference effect onto the second absorption grating 23.

  Further, in the above embodiment, the subject H is disposed between the X-ray source 12 and the first absorption type grating 22, but the subject H is arranged with the first absorption type grating 22 and the second absorption type. A phase contrast image can be generated in the same manner even when it is arranged between the grating 23.

(Second Embodiment)
In each of the above embodiments, the second absorption type grating 23 is provided independently of the FPD 21, but by using an X-ray image detector having a configuration disclosed in Japanese Patent Laid-Open No. 2009-133823. The second absorption type grating 23 can be eliminated. This X-ray image detector is a direct conversion type X-ray image detector including a conversion layer that converts X-rays into electric charges and a charge collection electrode that collects electric charges converted in the conversion layer. The charge collection electrode is configured by arranging a plurality of linear electrode groups formed by electrically connecting linear electrodes arranged at a constant period so that the phases are different from each other. Constitutes the intensity modulating means described in the claims.

  FIG. 9 illustrates the configuration of the X-ray image detector (FPD) of this embodiment. The pixels 70 are two-dimensionally arranged at a constant pitch along the x and y directions, and each pixel 70 has a charge collection for collecting the charges converted by the conversion layer that converts X-rays into charges. An electrode 71 is formed. The charge collection electrode 71 is composed of first to sixth linear electrode groups 72 to 77, and the phase of the arrangement period of the linear electrodes of each linear electrode group is shifted by π / 3. Specifically, when the phase of the first linear electrode group 72 is 0, the phase of the second linear electrode group 73 is π / 3, the phase of the third linear electrode group 74 is 2π / 3, The phase of the fourth linear electrode group 75 is π, the phase of the fifth linear electrode group 76 is 4π / 3, and the phase of the sixth linear electrode group 77 is 5π / 3. Charges in the y direction of the pixels 70 are stored through the respective linear electrode groups.

  Further, each pixel 70 is provided with a switch group 78 for reading out the charges collected by the charge collecting electrode 71. The switch group 78 includes TFT switches provided in each of the first to sixth linear electrode groups 72 to 77. By collecting the charges collected by the first to sixth linear electrode groups 72 to 77 individually by controlling the switch group 78, six types of fringe images having different phases can be obtained by one imaging. A phase contrast image can be generated based on these six types of fringe images. In the marker detection portion, if the average value of the intensity of the six types of fringe images is calculated, a picture equivalent to the transmission image can be obtained, so that the position of the marker image can be detected.

  By using the X-ray image detector having the above configuration in place of the FPD 21, the second absorption type grating 23 is not required from the imaging unit 13, so that the cost can be reduced and the thickness can be further reduced. Further, in the present embodiment, it is possible to acquire a plurality of fringe images that have been intensity-modulated at different phases by one shooting, so that physical scanning for fringe scanning becomes unnecessary, and the above scanning is performed. The mechanism 25 can be eliminated. Instead of the charge collection electrode 71, it is possible to use a charge collection electrode having another configuration described in Japanese Patent Laid-Open No. 2009-133823.

  Furthermore, as another embodiment in the case where the second absorption type grating 23 is not disposed, the fringe image (G1 image) obtained by the X-ray image detector is periodically sampled while changing the phase by signal processing. Thus, it is also possible to apply intensity modulation to the fringe image.

  Each embodiment described above can be applied to other radiation imaging systems for industrial use, in addition to the radiation imaging system for medical diagnosis.

10 X-ray imaging system 11 Marker 12 X-ray source (radiation source)
12a X-ray focus 13 Imaging unit 15 Image processing unit 17 Imaging control unit 21 Flat panel detector (FPD)
22 First Absorption Grating 22c, 23c X-ray Transmission Unit 23 Second Absorption Grating 25 Scanning Mechanism 27 Correction Processing Unit 29 Phase Differential Image Generation Unit 30 Phase Contrast Image Generation Unit

Claims (10)

A radiation source that emits radiation; and
A first grating that passes the radiation to generate a fringe image;
Intensity modulation means for applying intensity modulation to the fringe image at a plurality of relative positions different in phase with respect to the periodic pattern of the stripe image;
A marker attached to a subject disposed between the radiation source and the first grating, or between the first grating and the intensity modulating means;
A radiation imaging system comprising: a radiation image detector that detects a plurality of scanning images including the image of the subject and the image of the marker whose intensity is modulated by the intensity modulation means at each relative position;
The radiographic system characterized in that the marker is arranged at a position separated from the subject by a predetermined distance or more so that the marker image detected by the radiological image detector and the subject image do not overlap. . The predetermined distance is a distance T between both edge portions of the marker and the subject in the arrangement direction of the marker with respect to the subject, and the subject from a radiation focus in the optical axis direction of the radiation source. L _{4 is} a distance from the edge of the marker to L ₅ , a distance from the edge of the marker to the edge of the marker is L ₅ , and a distance from the edge of the marker to the radiation image detector is L _6. The following formula is satisfied, where w is the width of the radiation focus in the marker arrangement direction and T ₁ is the distance from one end of the radiation focus to the edge of the subject. Radiography system.
T ≧ {((L ₅ / L ₄ ) −1) × T ₁ } + {(L ₆ / (L ₅ + L ₆ )) × w} Correction processing means for detecting the marker image from each of a plurality of the scanned images and correcting the position of each scanned image so that the marker images coincide with each other;
A phase differential image that calculates a phase shift amount of the intensity modulation signal for each pixel based on the object image of each scanned image and generates a phase differential image of the object based on a phase differential value corresponding to the phase shift amount. Generating means;
Phase contrast for determining an integration range that does not include the marker image based on the position of the marker image, and integrating a phase differential image within the integration range generated by the phase differential image generation means to generate a phase contrast image Image generating means;
The radiation imaging system according to claim 1 or 2, further comprising:   The intensity modulation means includes a second grating having a periodic pattern in the same direction as the fringe image, and a scanning means for moving one of the first and second gratings at a predetermined pitch. The radiation imaging system according to claim 1.   5. The first and second gratings are absorption gratings, and the first grating projects radiation from the radiation source onto the second grating as a fringe image. Radiography system.   5. The first grating is a phase grating, and the first grating projects radiation from the radiation source as a fringe image onto the second grating by a Talbot interference effect. The radiation imaging system described. The radiological image detector is a radiological image detector comprising a conversion layer for converting radiation into electric charge and a charge collecting electrode for collecting electric charge converted in the conversion layer for each pixel,
The charge collection electrodes are arranged such that a plurality of linear electrode groups having a periodic pattern in the same direction as the fringe image are arranged in different phases.
The radiation imaging system according to claim 1, wherein the intensity modulation unit is configured by the charge collection electrode. Creating a fringe image by passing radiation through the first grating;
Applying intensity modulation to the fringe image by intensity modulation means at a plurality of relative positions having different phases with respect to the periodic pattern of the stripe image,
A plurality of scans including images of an object disposed between the radiation source and the first grating or between the first grating and the intensity modulation means and a marker attached to the object. A radiography method for detecting an image by a radiological image detector at each relative position,
The radiographic imaging method, wherein the marker is arranged at a position away from the subject by a predetermined distance or more so that the marker image detected by the radiological image detector and the subject image do not overlap. . The predetermined distance is a distance T between both edge portions of the marker and the subject in the arrangement direction of the marker with respect to the subject, and the subject from a radiation focus in the optical axis direction of the radiation source. L _{4 is} a distance from the edge of the marker to L ₅ , a distance from the edge of the marker to the edge of the marker is L ₅ , and a distance from the edge of the marker to the radiation image detector is L _6. The following formula is satisfied, where w is a width of the radiation focus in the marker arrangement direction, and T ₁ is a distance from one end of the radiation focus to the edge portion of the subject. Radiography method.
T ≧ {((L ₅ / L ₄ ) −1) × T ₁ } + {(L ₆ / (L ₅ + L ₆ )) × w} Detecting the marker image from each of a plurality of the scanned images, and correcting the position of each scanned image so that the marker images coincide with each other;
Calculating a phase shift amount of the intensity modulation signal for each pixel based on the subject image of each scanned image, and generating a phase differential image of the subject based on a phase differential value corresponding to the phase shift amount;
Determining an integration range that does not include the marker image based on the position of the marker image, and integrating a phase differential image within the integration range to generate a phase contrast image;
The radiation imaging method according to claim 8 or 9, further comprising:

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