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WO2011011014A1 - Imagerie par diffusion de rayons x - Google Patents

Imagerie par diffusion de rayons x Download PDF

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Publication number
WO2011011014A1
WO2011011014A1 PCT/US2009/051642 US2009051642W WO2011011014A1 WO 2011011014 A1 WO2011011014 A1 WO 2011011014A1 US 2009051642 W US2009051642 W US 2009051642W WO 2011011014 A1 WO2011011014 A1 WO 2011011014A1
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Prior art keywords
image
ray
grating
intensity
scattering
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Han Wen
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US Department of Health and Human Services
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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/40—Arrangements for generating radiation specially adapted for radiation diagnosis
    • A61B6/4035—Arrangements for generating radiation specially adapted for radiation diagnosis the source being combined with a filter or grating
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/42—Arrangements for detecting radiation specially adapted for radiation diagnosis
    • A61B6/4291—Arrangements for detecting radiation specially adapted for radiation diagnosis the detector being combined with a grid or grating
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/48—Diagnostic techniques
    • A61B6/484—Diagnostic techniques involving phase contrast X-ray imaging
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N23/00—Investigating or analysing materials by the use of wave or particle radiation, e.g. X-rays or neutrons, not covered by groups G01N3/00 – G01N17/00, G01N21/00 or G01N22/00
    • G01N23/02—Investigating or analysing materials by the use of wave or particle radiation, e.g. X-rays or neutrons, not covered by groups G01N3/00 – G01N17/00, G01N21/00 or G01N22/00 by transmitting the radiation through the material
    • G01N23/04—Investigating or analysing materials by the use of wave or particle radiation, e.g. X-rays or neutrons, not covered by groups G01N3/00 – G01N17/00, G01N21/00 or G01N22/00 by transmitting the radiation through the material and forming images of the material
    • G01N23/041—Phase-contrast imaging, e.g. using grating interferometers
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/02—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/027—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis characterised by the use of a particular data acquisition trajectory, e.g. helical or spiral

Definitions

  • the present application relates to X-rays, and, more particularly, to obtaining X-ray images using scattering.
  • X-rays can penetrate high density materials, making it the radiation of choice for non-invasive evaluation of bone.
  • X-ray techniques have advanced continuously from planar radiography to dual-energy radiography, CT densitometry and other approaches. These techniques use attenuation contrast where X-rays passing through the body are variably attenuated according to local material density and elemental composition. Because X-ray is a form of high-energy light, many contrast mechanisms that are standard in visible-light microscopy have also inspired parallel development in X-ray, such as dark-field, phase-contrast, fluorescence and confocal.
  • the dark-field, or scattering-based contrast arises from microscopic density fluctuations of X-rays that are deflected from a straight line (i.e., that scatter). While traditional X-rays produce a simple absorption contrast image, the dark-field X-ray image technology captures the scattering of the radiation within a material to give a clear, defined image that can show subtle inner changes in bone, soft tissue, or alloys. Small angle X-ray scattering is fundamentally different from absorption and refraction in that it is caused by spatial variation of electron density on the atomic to sub-micrometer scale, which leads to dispersion of the transmitted X-ray beam.
  • Imaging the distribution of X-ray scattering remains a difficult task due to requirements for specialized X-ray optical components and/or brilliant sources.
  • the majority of available techniques use raster or line scans of narrowly collimated beams, which are produced with single-crystal filters or pin-hole and slit apertures.
  • Non-raster simultaneous scattering imaging using polychromatic X-ray tubes has been proposed which takes advantage of scattering-induced blurring of sharp features as a result of the disruption of X-ray spatial coherence.
  • One such technique employs a high-density X-ray phase grating to phase-modulate a monochromatic component of the transmitted beam, which results in intensity oscillation at proper distances from the phase grating due to interference effects. This oscillation was resolved by placing a second high-density intensity grating over the detector surface, and taking a series of X-ray exposures for different positions of the second grating.
  • the X-ray gratings require unique fabrication and are expensive. Additionally, precise distances need to be established between the two gratings for proper imaging. Finally, multiple exposures need to occur, making the overall process time consuming, expensive.
  • a mask or multiple masks of periodically arranged opaque areas are placed in the x-ray path, such that periodic dark shadows are created on a recorder surface either by direct geometric shadowing or by wave-interference effects.
  • the shadow areas only receive x-ray which is scattered in the object.
  • the signals of these shadow areas are subtracted from the raw image to yield an image free of the effects of scattering.
  • the present invention overcomes the drawbacks of the prior art by allowing scattering images to be obtained using a single exposure and using commercially available gratings.
  • a scattering imaging method uses an intensity grating to modulate the intensity of a beam of an X-ray radiation source.
  • a detector captures a raw image from the modulated intensity pattern.
  • a scattering image can be automatically generated from the detected modulated intensity pattern.
  • both a scattering image and a phase-contrast image are obtained in a single exposure. Additionally, exact positioning of the grating is unnecessary, as the method works for any non-zero distance between the grating and the detector so long as the grating is far enough from the x-ray source to cause sufficient modulation of the x-ray intensity. Thus, the speed and ease of implementation makes it suitable for non-destructive testing, security screening, and medical diagnostic exams.
  • FIG. 1 shows a perspective view of an apparatus, according to one embodiment, for obtaining scattering images.
  • FIG. 2 is an embodiment of a method for generating a scattering image.
  • FIG. 3 shows further details of a method that can be used in conjunction with the method of FIG. 2 for generating a scattering image.
  • FIG. 4 is an embodiment of a method for generating a phase-contrast image.
  • FIG. 5 is an example showing absorption and scattering images of a tree branch.
  • FIG. 6 shows a view of a Fourier spectrum using both vertical and horizontal gratings.
  • FIG. 7 shows images obtained using the method according to one embodiment.
  • FIG. 8 shows an attenuation image and scattering images obtained using vertical and horizontal gratings.
  • FIG. 9 is an embodiment of an intensity grating using X-ray opaque and
  • FIG. 10 is an embodiment of an intensity grating having hexagonal X-ray transparent elements.
  • FIG. 11 is an illustration of reduction of artifacts through relative movement of a subject during imaging.
  • FIG. 12 is a flowchart of a method for reducing artifacts through movement.
  • FIG. 13 is another embodiment of an apparatus for removing scattering due to water.
  • FIG. 14 is a flowchart of a method for reduction of water scattering.
  • FIG. 15 is an illustration showing the reduction of water scattering.
  • a method is disclosed to form an X-ray scattering image from a single fundamental exposure.
  • An X-ray phase-contrast image can also be derived from the fundamental image.
  • Figure 1 shows an apparatus 10 having an X-ray source 12 and an X-ray detector 14.
  • a grating 16 having a periodic alternating pattern of X-ray transparent and X-ray opaque regions is interposed between the X-ray source 12 and the X-ray detector 14, either before or after an imaged object 18, such that the image at the detector surface contains a periodically modulated intensity pattern.
  • the grating is a low-density intensity grating used to create a pattern on the detector surface.
  • a distance is maintained between the detector surface and both the grating and the object, such that X-ray scattering in the object blurs the grating pattern and X-ray refraction in the object deforms the pattern. The distance need not be precise and is not fixed.
  • the X-ray image is recorded at a resolution smaller than one third of the period of the grating pattern.
  • the detected image is Fourier transformed and separate sub-images of complex values are formed by inverse Fourier transformation of separate regions surrounding the spectral peaks in the Fourier spectrum. These sub-images can be normalized to reference sub-images (which are acquired without any samples) to yield harmonic images.
  • the intensity ratio between a K ⁇ -order harmonic image and an N ⁇ -order harmonic image provides an image of X-ray scattering distribution.
  • the phase of a K ⁇ -order harmonic image represents gradients of the index of refraction in the direction perpendicular to the grating lines.
  • the grating can have different patterns of opaque and transparent regions, such as, linear, rectangular or hexagonal patterns. Other patterns can also be used whose Fourier transformation contains multiple peaks that are arranged at different angles around the origin.
  • a small relative movement of the sample and the imaging system in the direction that is perpendicular to the grating lines during the X-ray exposure can be used to remove artifacts at sharp interfaces in the sample. This can be further refined by modulating the X-ray intensity level during the movement.
  • One technique for causing such a relative movement is to use a stage 22 that moves the object during the exposure. Such a stage can be used in any of the embodiments described herein.
  • FIG. 2 is a flowchart of a method for obtaining a scattering image.
  • an intensity grating is positioned between an X-ray source and X-ray detector.
  • the X-ray source emits a beam of X-ray radiation towards a subject.
  • an X-ray detector captures (detects) a single image of the subject by detecting an X-ray beam that has passed through the intensity grating positioned before or after the subject.
  • the intensity grating modulates an amplitude of the X-ray beam with substantially no phase modulation. The result is that a modulated intensity pattern is generated and detected by the X-ray detector.
  • a scattering image is automatically generated using processing techniques carried out by the computer 20.
  • the scattering image can be derived from the detected image using only a single exposure and using an intensity grating, rather than a phase grating.
  • an intensity grating By using an intensity grating, commercially available gratings are readily available.
  • the distances between the grating and the detector need not be precise in order to obtain the scattering image.
  • FIG. 3 shows an embodiment of a method for generating the scattering image.
  • a raw image in the spatial domain is captured using an X-ray detector, such as a camera.
  • the raw image is converted into the spatial frequency domain image using well-known techniques, such as Fourier transforms, Laplace transforms, etc.
  • various peaks are visible, which corresponds to integer multiples of the basic spatial frequencies of the shadows of the grid projected on the detector surface.
  • an N*- order peak is selected from the image in the spatial frequency domain.
  • the N*-order peak is a zero-order peak, but any desired harmonic can be selected.
  • a spatial frequency domain filter such as a Hanning or Fermi filter
  • An inverse transform is performed to obtain an N ⁇ -order harmonic image in the space domain using the area around the Nth-order peak.
  • calibration is performed on the N ⁇ -order harmonic image by dividing that image with an N ⁇ -order harmonic image of a reference image that was taken without a subject. Such a reference image only needs to be taken once and stored on the computer 20 for later use.
  • the calibration can include taking a ratio of corresponding portions of the N ⁇ -order harmonic image and the reference image.
  • a K ⁇ -order harmonic is selected from the image in the spatial frequency domain and an area around it is multiplied with a filter and is inverse transformed in order to obtain a K ⁇ -order harmonic image in the space domain.
  • calibration is performed on the K ⁇ -order harmonic image in the space domain using the same calibration technique described above.
  • a scattering image is automatically generated by calculating a ratio between the N ⁇ -order harmonic image and the K ⁇ -order harmonic image.
  • FIG. 4 is a flowchart of a method for obtaining a phase-contrast image.
  • the reference image is provided in process block 400.
  • the reference image is taken without a subject, as already described.
  • a K ⁇ -order harmonic of the reference image is obtained.
  • Such a K ⁇ -order harmonic is obtained through a transformation of the reference image into the spatial frequency domain, selecting the K ⁇ -order peak, and filtering and inverse transforming an area around the Reorder peak to obtain a reference image of the K ⁇ -order harmonic in the space domain.
  • the phase of the K ⁇ -order harmonic of the reference image is subtracted from the phase of the K ⁇ -order harmonic of the image being processed in order to obtain the phase-contrast image.
  • FIG. 5 shows the advantage of obtaining the scattering image.
  • the top photo shows an absorption image of a cedar branch, while the bottom photo shows the scattering image.
  • the scattering image shows structures that are not visible in the absorption image. For example, the central core of the cedar branch running the length of the stem is clearly visible in the scattering image, but absent from the absorption image. This bright band running down the center of the scattering image corresponds to the pith of the stem. Additionally, the combination of absorption and scattering data can help distinguish different materials that appear similar in absorption images.
  • FIG. 6 illustrates the effects of using a vertical grating versus a horizontal grating. The result is that the alignment of the peaks in the spatial frequency domain is rotated ninety degrees.
  • X-ray scattering is the broadest in a plane perpendicular to fibers in the subject and results in different scattering image intensities between vertical and horizontal grating placement.
  • the X-ray radiation source was a fixed- anode tungsten-target tube operating at 50 kVp/0.6mA (SB-80-lk, Source-Ray Inc., Bohemia, NY, USA).
  • the tube has a beryllium window, and the grating mask effectively adds 1.7 mm of aluminum filtration. No other filters were used.
  • the half-value layer (HVL) of the X-ray cone -beam is 1.3 mm aluminum (ionization chamber dosimeter, PTW, Freiburg, Germany).
  • An example exposure lasted 10 seconds and delivered 0.19 mGy incident radiation at the sample, equivalent to a thoracic X-ray exam.
  • the X-ray camera used was a 16 bit CCD camera of matrix size of
  • a grating was used that had 200 lines-per-inch (lpi) parallel radiography anti-scatter grating of 10:1 grating ratio (MXE Inc., Los Angeles, CA, USA) and it was interposed between the source and the camera, with the sample placed immediately down-beam from it. Both the scattering length scale and the exposure time were factors when determining the geometry of the device.
  • the scattering length scale, d is the maximal length scale of the electron density fluctuation in the sample that still gives rise to appreciable scattering signal. For example, particles of radius d' scatters X-rays into a cone. For a given device geometry, the scattered cone from a single X-ray appears as a dispersed pattern on the X-ray camera.
  • the size of the dispersion pattern on the camera is inversely related to the size of the particle and the X-ray wavelength.
  • the size of the structure whose scattering can be detected by this means is limited to an upper threshold which is dependent on the X-ray wavelength ⁇ , the grating period Po, the distances between the source and the grating D 1 , between the grating and the sample D 2 , between the sample and the camera D 3 (FIG. 1), and the order of the harmonic peak n, as
  • the upper threshold is defined as the scattering length scale of the nth-order harmonic image of this specific device setup: d n - ⁇ D ⁇ .
  • the larger this length scale the higher the scattering signal. Since the sample-to-grating distance is small, D 2 ⁇ 0. Then, for a given total length of the layout, the scattering length scale is maximized when the grating and sample are placed midway between the X-ray tube and the camera. Additionally, the scattering length scale increases with the camera-to-source distance, but at the cost of exposure time. With a camera-to-source spacing of 1 meter, it is desirable to have 10 seconds of exposure. The distance between the grating and the camera was then 0.5 m.
  • the scattering length scale is not a single value, but a distribution of values corresponding to the distribution of X-ray energies. Based on the peak of the energy spectrum, the first-order scattering length scale observed was approximately 100 nanometers.
  • the exposure time should be two orders of magnitude shorter than the current 10 seconds. This is in line with current chest X-ray exams.
  • the resolution of the camera or flat panel detector should be sufficient to sample the first-order modulation of the grating.
  • the detector pixel should not be larger than one-third the period of the grating shadow.
  • FIG. 7 shows actual images taken in conjunction with the method for obtaining the scattering and phase-contrast images, from a single exposure.
  • the raw image from the camera is shown in FIG. 7a.
  • the raw image is Fourier transformed to reveal several distinct peaks (FIG. 7b), including the zero ⁇ -order peak at the center containing radiation un-modulated by the grating mask, and the first and higher-order peaks from the grating shadows. Similar to the use of physical apertures in dark-field visible-light or electron microscopy, these peaks are individually selected with mask filters in the Fourier space. The areas around the selected peaks are then inverse Fourier transformed to yield the zero*, first and higher-order images separately (FIGs. 7c and 7d).
  • the scattering image (FIG. 7f) was obtained by dividing the magnitude of the first harmonic image by the magnitude of the attenuation image.
  • the phase contrast image (FIG. 7e) is obtained by using the phase of the first-order harmonic image and subtracting of phase of the corresponding no-sample reference image.
  • the harmonic images can be normalized (calibrated) with reference images acquired without samples (not shown here).
  • FIG. 8 illustrates X-ray attenuation and scattering images of the hind limb of a rat. Differences in scattering intensity are visible between perpendicular and parallel alignment of the grating and the bone axes. In the dense superficial cortical bone of the tibia, the scattering signal is brighter when the bone axis is parallel to the grating.
  • FIG.8a shows regular attenuation.
  • FIG. 8b shows a scattering image of the same limb with vertical placement of the grating.
  • FIG. 8c shows a scattering image with horizontal placement of the grating.
  • FIG. 8d shows a graph of scattering versus attenuation. Pixels are divided into four bins according to their attenuation values. Mean values of scattering of the bins, with corresponding standard deviations, are plotted. For all bins, scattering is greater when the grid is parallel to the tibia than when it is perpendicular.
  • the gratings do not affect phase, although some phase modulation is acceptable.
  • FIG. 9 shows a horizontal grating 900.
  • the grating has a periodic alternating pattern of X-ray opaque lines 902 and X-ray transparent lines 904.
  • Example materials include lead for the X-ray opaque lines and aluminum or plastics for the X-ray transparent lines. Other materials can be used that are well-known in the art.
  • the grating is structured so that the period P of grating shadows on the surface of the X-ray detector is equal or less than P/3 of the detector pixel spacing.
  • FIG. 10 is an illustration of a hexagonal intensity grating 1000 that generates a Fourier spectrum shown at 1002.
  • the hexagonal grating can be made of similar material that was described above. Other shapes can be used, such as square, rectangular, circular, etc.
  • the gratings modify amplitude and only minimally impact phase of an X-ray beam passing through the grating.
  • Gratings of square or hexagonal cells enable imaging of multiple scattering directions in a single exposure, while gratings of narrow passing slits more evenly distribute the transmitted X-ray energy among the Fourier harmonic peaks and allow scattering imaging at two or more length scales simultaneously.
  • a further consideration is that very sharp edges in the sample can interfere with the fine grating shadows, or equivalently in the Fourier spectrum, adjacent peaks may overlap. This can be prevented by slightly moving the sample in the direction perpendicular to the grating lines over a distance equal to the grating period during the X-ray exposure. The result is that sharp edges in the sample itself can be blurred to widths that are similar to the grating period, while the grating shadows are not affected. The motion-blurring effectively limits the extent of each peak in the Fourier spectrum to less than half the distance between the peaks, and thus avoiding any chance of interference between the peaks.
  • the stage 22 can be moved at a constant speed, during the exposure time T, over a distance ⁇ in the y direction (for a vertical grating) or the x direction (for a horizontal grating), or in a circular motion (for a hexagonal grating).
  • the X-ray tube current during the exposure time can vary and is denoted as I s (t).
  • the resulting sample image F(x d ,y d ) on the detector surface is related to the image without motion F 0 (XdJd) by the relationship
  • M(k y ) is the sample motion projected onto the detector. If the tube current I s remains constant over the exposure time, then M(k y ) is a sine function:
  • the band-limiting function M(k y ) can be improved by modulating the tube current I s over the time of exposure.
  • a Gaussian modulation of the tube current is more desired than a constant tube current.
  • FIG. 11 shows two images.
  • a first static image has artifacts around the edges, while the sample that was moved has reduced artifacts.
  • FIG. 12 shows an embodiment of a method for reducing artifacts.
  • an X-ray beam is emitted for a period of time (the time can vary depending on the application), such as between 1 and 15 seconds.
  • the subject of the X-ray is moved relative to the apparatus.
  • the stage FIG. 1
  • the movement can be linear, circular, or other motions.
  • the apparatus can be moved while the subject remains static.
  • Another optional feature is to have the X-ray beam vary in intensity (process block 1204) over the period while the relative movement is occurring.
  • the artifacts are removed using a ratio of harmonic images, such as outlined in FIG. 3.
  • FIG. 13 shows how the apparatus can be modified in order to reduce scattering caused by water.
  • a first grating 1302 is placed before or after the subject 1304.
  • a second grating 1306 is placed near the X-ray detector (e.g., camera).
  • the first g rating has a different orientation that the second grating so as to obtain adequate separation in the Fourier spectrum.
  • the resultant Fourier spectrum shows a peak 1308 associated with grating 1302 and a peak 1310 associated with grating 1306. These different peaks can be used to cancel the scattering effects of water.
  • reducing water scattering provides a clearer image of the more interesting aspects of the tissue being analyzed.
  • FIG. 14 shows a flowchart of a method for reducing scattering caused by water.
  • a first intensity grating is placed before or after the subject.
  • a second intensity grating is placed adjacent to the X-ray detector.
  • a raw image is then taken and a Fourier transform is performed to place the raw image in the Fourier spectrum.
  • a N ⁇ -order peak and K ⁇ -order peak are selected such that one peak is associated with one grating and the other peak with the other grating.
  • the areas around the peak are determined and the same filter can be applied to both.
  • the inverse transforms are performed and a ratio is taken (process block 1408) in a manner similar to as previously described in order to obtain a scatter image with the scattering of water reduced or removed.
  • FIG. 15 shows the results of the water-suppressed scatter image.
  • the zero ⁇ -order image provides the distribution of the attenuation of transmitted X-rays through the sample:
  • D 0 (x, y) - In[Z 0 (x, y) I I Og (x, y)] , (6)
  • / 0 is the magnitude of the zero ⁇ -order image
  • / Og is the magnitude of a zero*- order image without any samples
  • (x, y) are the coordinates in the image
  • the natural log yields values that are linearly related to the sample thickness.
  • This is a regular radiography attenuation image.
  • the higher-order images are frequency- modulated versions of the zero ⁇ -order image and are therefore more severely attenuated by the blurring effect of scattering in the sample.
  • the magnitude ratio of a high-order image to the zero ⁇ -order image measures the degree of blurring of the grating shadows.
  • the expression for an image that is solely dependent on X-ray scattering in the sample is as follows:
  • I n is the magnitude of the nth-order harmonic image
  • / ng is the magnitude of a nth-order image without any samples.
  • the normalization (calibration) relative to the no-sample reference images remove features of the grating itself, and the natural log makes the values linearly related to the thickness of the sample. Scattering images of different orders reflect structures of different length scales. In practice, the first- order peak usually has the highest intensity and provides the best signal-to-noise ratio.
  • phase of the high-order images is influenced by the slight bending of X-rays from refraction. X-ray refraction in the direction perpendicular to the grating shifts the grating shadows, resulting in phase shifts in the high-order images. Phase-contrast images can be obtained concurrent with the scattering images as
  • the current model of compact bone is a lamellar formation of bundles of mineralized collagen fibrils with crystalline material between them. The bundles are aligned in each layer, and the size of a fibril is about 80 nm, on the same order as the length scale observed by the scattering image.
  • the angular distribution of scattering is the broadest in the equatorial plane.
  • the accumulated effect through multiple layers is broader dispersion in the direction perpendicular to the layers, or the periosteal surface. Therefore, when the stripes of the grating mask are parallel to the bone surface (bone axis), their shadows are maximally blurred by X-ray dispersion, resulting in the highest scattering image intensity.
  • the notion that the anisotropy of the scattering image reflects the ordered structure of compact bone is in agreement with previous small-angle X-ray scattering measurements of bone sections. These measurements show that the azimuthal scattering distribution of a single fibril bundle has an ellipsoidal shape with the short axis in the direction of the bundle, and the scattering distribution of a plate-shaped mineral particle is an ellipsoid with the long axis perpendicular to the plate.
  • the values of the phase-contrast images are determined by spatial gradients of material density at the scale of the image resolution.
  • the calcified compartment is denser than the soft-tissue compartment, leading to local density gradients in random directions at the scale of the pore size of 100 ⁇ m.
  • a solution is to decrease the distance between the grating and the detector. For example, a grating period and final resolution of 0.2 mm is feasible if the grating is placed at 25 cm from the detector and 75 cm from the X-ray tube.
  • the tradeoff is that the patient may not fit into the 25 cm space and need to be positioned in-front of the grating. Without the screening of the grating the radiation exposure of the patient may increase by approximately a factor of two. Given all of the above considerations, the optimal device parameters and layout for human imaging need to be determined experimentally.
  • the thickness of the sample affects Fourier scattering imaging and phase-contrast imaging in different ways.
  • the transmitted X-rays become more dispersed, and the scattering signal increases proportionally with the sample thickness in the same fashion as regular X-ray attenuation.
  • the phase signal originates from the refractive bending of the X-rays at tissue interfaces, and depends on the number of interfaces the X-rays pass through and the direction of the density gradients at these interfaces.
  • the phase signals from different segments along an X-ray path do not add constructively, and it becomes weaker when the X-ray becomes more dispersed. For these reasons, scattering imaging is more suited for thick samples.
  • the technique of Fourier X-ray scattering radiography is able to acquire attenuation and scattering images in bone in a single exposure, which indicates material density and fine structure respectively.
  • the coordinates (x, y) in the sample are within planes perpendicular to the central axis of the X-ray beam, and the coordinate z is along the beam axis.
  • Positions in the sample are labeled r 0 , and positions on the detector plane r ⁇
  • the period of the grating shadows on the detector surface is P, and the X-ray wave length is ⁇ .
  • the projection image of the grating onto the detector surface is a periodic pattern of dark and bright bands. Its two-dimensional Fourier transformation, or spectrum, contains a number of discrete peaks corresponding to the harmonics of the basic grating frequency.
  • the projection image of the grating is denoted G, its spectrum as g.
  • the projection image of the sample alone without the grating is denoted F, its spectrum as /.
  • the image of the sample with the grating present is denoted as F g , its spectrum f g .
  • F g is the raw data obtained in the experiment.
  • the nth-order harmonic image is the two dimensional inverse Fourier transformation
  • the sample is viewed as constructed by adding thin layers that are perpendicular to the beam axis Z, incrementing from the grating side to the detector side.
  • the layer thickness dz is sufficiently small so that X-ray photons either go through without any collisions or be scattered once.
  • the image intensity at position r d on the detector surface is
  • ⁇ s (r) ⁇ ⁇ ( ⁇ , r)d ⁇ . (19)
  • the first integral in eq.(18) is the X-rays that strike r d on the detector without experiencing any events in the layer, and the second integral is the X-rays that were scattered from various locations within the layer before striking r d . Again under the assumption that scattering is limited to small angles, the X-rays that strike r d on the detector surface all exited the layer dz in a small area about a point r 0 , within which the absorption and scattering coefficients are uniform. Then eq.(18) is simplified to
  • Equation (22) means that the angular distribution of the scattered X-rays at any location in the layer is proportional to the three-dimensional Fourier transformation of an auto-correlation function of the electron charge density distribution at that location. Substituting eq.(22) into equation (21) yields
  • R ⁇ (x, y; r 0 ) is defined in terms of the auto-correlation in eq.(19) in the vicinity of r o :
  • R 2 (x, y) -R(x,y,w)]dw.
  • Equation (24) relates the scattering effect to electron charge density variation and thereby the microscopic structure of the sample.
  • n 0, dz (26) Therefore, the zeroth-order image is simply a conventional attenuation image. Summing all the layers that constitute the sample by integrating eq.(24) over the thickness of the sample yields
  • Equation (27) means that the sample harmonic images should be normalized with the no-sample reference harmonic images in order to remove any features associated with the device itself.
  • Equations (27) and (28) are the basis for obtaining attenuation and scattering images from the harmonic images.
  • the normalization reference images G n are acquired for a given device setup beforehand and are used for all samples if the setup is not altered. Scattering length scale
  • R n jjdxdy JdZ 1 jdz 2 p(x, y, y + d ⁇ ,z 2 )l
  • P is the period of the grating shadows on the detector surface, it is related to the period of the grating itself by a geometric magnification:
  • the scattering signal depends on the difference between the local electron charge density distribution and the same distribution shifted by the distance d n in the direction perpendicular to the grating lines. If the sample is made of structures that are much larger than this scale, then the charge density changes little over this distance and the difference between the two copies of the density distribution is small, leading to low scattering signal; if on the other hand the structures are smaller than the length scale, than the shift results in appreciable changes of the density distribution and significant scattering signal.
  • the detector pixel size should be less than /73 in order to sufficiently sample the first-order harmonic peaks:

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Abstract

L'invention porte sur un procédé et un appareil (10) pour obtenir une image de diffusion de rayons X. Dans un mode de réalisation, une grille d'intensité (16) est utilisée pour moduler l'intensité d'un faisceau d'une source de rayonnement de rayons X (12). Un détecteur (14) capture une image brute provenant du motif à intensité modulée. Une image de diffusion peut être automatiquement générée à partir du motif à intensité modulée détecté. Dans un autre mode de réalisation, à la fois une image de diffusion et une image à contraste de phase sont obtenues en une seule exposition. De plus, un positionnement exact de la grille est inutile, étant donné que le procédé fonctionne pour toute distance non nulle entre la grille et le détecteur. Enfin, des grilles du commerce peuvent être utilisées. Ainsi, la vitesse et la facilité de mise en œuvre permettent des applications d'imagerie par rayons X.
PCT/US2009/051642 2009-07-24 2009-07-24 Imagerie par diffusion de rayons x Ceased WO2011011014A1 (fr)

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JP2013138836A (ja) * 2011-12-05 2013-07-18 Fujifilm Corp 放射線撮影装置
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WO2013084658A1 (fr) * 2011-12-05 2013-06-13 富士フイルム株式会社 Appareil de radiographie
US8989347B2 (en) 2012-12-19 2015-03-24 General Electric Company Image reconstruction method for differential phase contrast X-ray imaging
WO2014103269A1 (fr) * 2012-12-27 2014-07-03 Canon Kabushiki Kaisha Appareil de calcul, programme, et système d'imagerie par rayons x
US9014333B2 (en) 2012-12-31 2015-04-21 General Electric Company Image reconstruction methods for differential phase contrast X-ray imaging
US9412481B1 (en) * 2013-01-22 2016-08-09 Michael Keith Fuller Method and device for producing and using localized periodic intensity-modulated patterns with x-radiation and other wavelengths
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US9700275B2 (en) 2013-05-10 2017-07-11 Paul Scherrer Institut Quantitative X-ray radiology using the absorption and scattering information
WO2014180683A1 (fr) * 2013-05-10 2014-11-13 Paul Scherrer Institut Radiologie quantitative par rayons x faisant intervenir les informations d'absorption et de diffusion
CN104807841A (zh) * 2013-10-15 2015-07-29 财团法人工业技术研究院 增加穿透式小角度x光散射的散射强度的装置
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WO2016008956A1 (fr) * 2014-07-17 2016-01-21 Koninklijke Philips N.V. Procédé de reconstitution itérative pour imagerie spectrale à contraste de phase
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WO2017032512A1 (fr) 2015-08-25 2017-03-02 Paul Scherrer Institut Sensibilité de diffusion omnidirectionnelle et de phase bidirectionnelle en interférométrie à réseau de diffraction à cycle unique
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US10514342B2 (en) 2015-08-25 2019-12-24 Paul Scherrer Institut Omnidirectional scattering- and bidirectional phase-sensitivity with single shot grating interferometry
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