WO2014077376A1 - Dispositif de diagnostic par rayons x - Google Patents

Dispositif de diagnostic par rayons x Download PDF

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Publication number
WO2014077376A1
WO2014077376A1 PCT/JP2013/080954 JP2013080954W WO2014077376A1 WO 2014077376 A1 WO2014077376 A1 WO 2014077376A1 JP 2013080954 W JP2013080954 W JP 2013080954W WO 2014077376 A1 WO2014077376 A1 WO 2014077376A1
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image
grid
image data
data
ray
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PCT/JP2013/080954
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English (en)
Japanese (ja)
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加藤　久典
靖宏 菅原
由昌 小林
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株式会社 東芝
東芝メディカルシステムズ株式会社
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Publication of WO2014077376A1 publication Critical patent/WO2014077376A1/fr

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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
    • 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

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  • Embodiments of the present invention relate to an X-ray diagnostic apparatus that generates image data representing an internal form of a subject by X-rays.
  • X-rays incident on the subject cause scattering in the subject.
  • Scattered X-rays (scattered rays) generated at this time reduce the contrast and sharpness of the image and adversely affect image diagnosis. Therefore, in order to remove turbulence, it is common to arrange a grid with alternating X-ray absorption intermediate material (for example, aluminum or fiber) and lead foil on the detection surface side of the X-ray detector. is there.
  • interference fringes between FPD pixels and a grid grating may occur in an image.
  • FPD Full Panel Detector
  • the maximum lattice density that can be stably manufactured with the current technology is approximately 80 LP / cm.
  • the intermediate material of the grid is limited to aluminum.
  • the intermediate material is still limited to aluminum.
  • aluminum which is a metal, has higher rigidity than a fiber such as paper, and can be processed with high accuracy.
  • Grids using aluminum as an intermediate material have a lower direct line transmittance than grids using fiber as an intermediate material.
  • JP-A-5-244508 There is JP-A-5-244508 as a related technology.
  • the problem to be solved by the present invention is to provide an X-ray diagnostic apparatus capable of obtaining an image with good image quality while reducing exposure of a subject.
  • the X-ray diagnostic apparatus generates an X-ray source, a grid, an X-ray detector, a first image generation unit, and a second image data that generates a plurality of second image data having different spatial frequency bands from the first image data.
  • a final image generation unit that generates a final image by combining the plurality of second images and the second image subjected to the scattered radiation correction process.
  • FIG. 1 is a block diagram showing a main configuration of an X-ray fluoroscopic apparatus according to an embodiment.
  • FIG. 2 is a block diagram showing a main configuration of the image correction unit according to the embodiment.
  • FIG. 3 is a diagram showing a flow of processing by the interference fringe removal processing unit according to the embodiment.
  • FIG. 4 is a diagram illustrating an example of frequency characteristics of the LPF processing according to the embodiment.
  • FIG. 5 is a diagram showing a flow of processing by the scattered radiation correction processing unit according to the embodiment.
  • FIG. 6 is a diagram for explaining the scattered radiation correction processing unit according to the modification.
  • FIG. 7 is a block diagram illustrating a main configuration of an image correction unit according to the second embodiment.
  • FIG. 8 is a diagram illustrating an example of a frequency characteristic graph of frequency band data and background data.
  • FIG. 9 is a diagram showing a graph in which only one frequency band data in FIG. 8 is extracted.
  • FIG. 10 is a diagram illustrating an example of the relationship between the background data and the Fourier transform of the point spread function.
  • an X-ray fluoroscopic imaging apparatus 1 that performs fluoroscopy and imaging of a subject is disclosed as an example of an X-ray diagnostic apparatus.
  • FIG. 1 is a block diagram showing a main configuration of the X-ray fluoroscopic apparatus 1.
  • the X-ray fluoroscopic apparatus 1 includes an X-ray high voltage unit 2, an X-ray tube 3 (X-ray source), an X-ray movable diaphragm 4, a top plate 5, a grid 6 (first grid), and an FPD 7 (X-ray detector).
  • An AD converter 8 image generation unit
  • a pixel value calculation unit 9 an image correction unit 10
  • an image processing unit 11 an image display unit 12, a system control unit 13, and an X-ray control unit 14.
  • the X-ray high voltage unit 2 generates a high voltage based on a power source supplied from a commercial AC power source or the like, and applies this high voltage to the X-ray tube 3.
  • the X-ray high voltage unit 2 may use any of a transformer type, an inverter type, and a capacitor type.
  • the X-ray tube 3 generates X-rays having a dose and quality according to the tube voltage and tube current applied from the X-ray high voltage unit 2.
  • the X-ray movable diaphragm 4 is a device for narrowing X-rays generated from the X-ray tube 3 with respect to the region of interest of the subject P.
  • the X-ray movable diaphragm 4 has four slidable diaphragm blades, and narrows the X-rays by sliding these diaphragm blades.
  • the top plate 5 is a bed on which the subject P is placed.
  • the FPD 7 is provided below the top plate 5.
  • the FPD 7 has a large number of detection elements that detect X-rays generated by the X-ray tube 3. These detection elements convert the X-rays that have passed through the subject P into charges and accumulate them.
  • the grid 6 is a focusing grid in which, for example, an intermediate substance and a lead foil with little X-ray absorption are alternately arranged, and each lead foil is inclined toward one point on the grid center line in a direction perpendicular to the grid surface. is there.
  • the grid 6 may be a parallel grid in which lead foils are arranged in parallel.
  • the grid 6 is provided between the top 5 and the FPD 7 and removes a part of the scattered radiation from the X-ray transmitted through the subject P.
  • the AD converter 8 reads out the electric charge accumulated in the FPD 7 in synchronization with the X-ray irradiation from the X-ray tube 3 to the subject P, and AD converts the read electric charge (analog signal) to X-ray image data. Is generated.
  • the AD converter 8 outputs the generated X-ray image data to the pixel value calculation unit 9.
  • the pixel value calculation unit 9 sets a calculation ROI (region of interest) for the X-ray image data input from the AD converter 8, and the average pixel value, maximum pixel value, minimum pixel value, and center pixel in this ROI Statistics such as the value and the most frequent pixel value are calculated and output to the system control unit 13 and the X-ray control unit 14. Further, the pixel value calculation unit 9 outputs the X-ray image data input from the AD converter 8 to the image correction unit 10. When the statistic regarding the pixel value is input from the pixel value calculation unit 9, the system control unit 13 transfers the statistic to the image correction unit 10.
  • the image correction unit 10 performs various corrections on the X-ray image data input from the pixel value calculation unit 9. Details of this correction will be described later.
  • the image correction unit 10 outputs the corrected X-ray image data to the image processing unit 11.
  • the image processing unit 11 performs display image processing (spatial filter processing, window conversion, gamma curve processing, etc.) on the X-ray image data input from the image correction unit 10.
  • the image processing unit 11 outputs the X-ray image data after the image processing to the image display unit 12.
  • the image display unit 12 displays an image based on the X-ray image data input from the image processing unit 11.
  • the system control unit 13 controls each unit of the X-ray fluoroscopic imaging apparatus 1 in accordance with a command input by a user from a console (not shown). For example, the system control unit 13 outputs X-ray movable diaphragm information that specifies the stop position of each diaphragm blade to the X-ray movable diaphragm 4. The X-ray movable diaphragm 4 operates each diaphragm blade based on the X-ray movable diaphragm information. Further, the system control unit 13 outputs X-ray conditions such as tube voltage and tube current target values to the X-ray control unit 14. In addition, the system control unit 13 controls a fluoroscopic table on which the X-ray tube 3, the X-ray movable diaphragm 4, the top plate 5, the grid 6 and the FPD 7 are provided, and controls the FPD 7.
  • the X-ray control unit 14 causes the X-ray high voltage unit 2 to apply the target tube voltage and tube current indicated by the X-ray condition to the X-ray tube 3. By changing the X-ray conditions, the system control unit 13 can adjust the quality and dose of X-rays generated by the X-ray tube 3.
  • the X-ray control unit 14 has a function of performing automatic brightness adjustment (ABC) based on the statistics input from the pixel value calculation unit 9. (grid)
  • ABSC automatic brightness adjustment
  • the grid 6 will be described. In this embodiment, in order to reduce the exposure of the subject P, for example, a grid 6 having the following specifications is employed.
  • Lattice density N The frequency f of the interference fringes generated in the X-ray image by the grid 6 is as large as possible in the range from 1/2 of the Nyquist frequency fa of the FPD 7 to the Nyquist frequency fa (fa / 2 ⁇ f ⁇ fa).
  • the lattice density N is preferably closer to 34 LP / cm within the range of 34 to 50 LP / cm.
  • Lattice ratio r, intermediate material Direct line transmittance is as high as possible.
  • the lattice ratio r is about 6: 1 to 10: 1 and the intermediate material is a fiber.
  • X-rays generated in the X-ray tube 3 so that the direct dose detected by the FPD 7 when using the grid 6 having the above specifications is as close as possible to the direct dose when using the conventional general grid.
  • the exposure dose of the subject P can be reduced, but the scattered radiation is increased, and the SN ratio and the contrast are reduced. Furthermore, it is expected that interference fringes appear in the X-ray image data as in the conventional case.
  • the image correction unit 10 removes interference fringes from the X-ray image data, reduces noise generated in the X-ray image data due to a decrease in the SN ratio, and improves the contrast of the X-ray image data.
  • FIG. 2 is a block diagram illustrating a main configuration of the image correction unit 10.
  • the image correction unit 10 includes an interference fringe removal processing unit 100, a noise reduction processing unit 101, a scattered radiation correction processing unit 102, and a system information processing unit 103.
  • Each of the processing units 100 to 103 may be a circuit including components such as independent processors, or may be a function realized by software by one processor.
  • the interference fringe removal processing unit 100 executes processing for removing the interference fringes by the grid 6 from the X-ray image data input from the pixel value calculation unit 9.
  • the flow of processing executed by the interference fringe removal processing unit 100 is shown in FIG.
  • the X-ray image data input from the pixel value calculation unit 9 to the image correction unit 10 is referred to as original image data I0, and the column direction of pixels included in the original image data I0 is defined as the x direction and the row direction is defined as the y direction. It is assumed that the interference fringes due to the grid 6 are generated in parallel with the y direction. In this case, the arrangement direction of the interference fringes coincides with the x direction.
  • the interference fringe removal processing unit 100 performs a one-dimensional spatial filter (LPF: Low Pass Filter) process S11 through which a low frequency component passes in the x direction of the original image data I0.
  • An example of the frequency characteristic of the LPF process S11 is shown in FIG.
  • the horizontal axis is the spatial frequency (LP / cm), and the vertical axis is the LPF gain.
  • the frequency f of the interference fringes is in the range from 1 ⁇ 2 of the Nyquist frequency fa to the Nyquist frequency fa (fa / 2 ⁇ f ⁇ fa).
  • the gain is set in the range of 0 or more and 1 or less.
  • the frequency band component with a gain of 0 is removed, and the frequency band component with a gain of 1 remains without being affected by the LPF process. This gain is suddenly reduced from 1 to 0 immediately before the frequency f of the interference fringes.
  • the frequency band of an image necessary for diagnosis (such as a portion representing the internal form of the subject P) is sufficiently smaller than 1/2 of the Nyquist frequency fa. Therefore, even if the gain is reduced to 0 immediately before the frequency f of the interference fringes as shown in FIG. 4, the influence on the image necessary for diagnosis hardly occurs.
  • X-ray image data was collected by a 2 ⁇ 2 or 3 ⁇ 3 detection element in addition to a mode in which image data is configured with an output from one detection element included in the FPD 7 as one pixel. There is also a mode in which the charge is averaged to one pixel. The frequency f of the interference fringe varies depending on these modes. Further, interference fringes do not occur depending on the mode.
  • the system control unit 13 notifies the system information processing unit 103 of the types of these modes.
  • the system information processing unit 103 stores in advance an optimal kernel size and gain for each mode, and notifies the interference fringe removal processing unit 100 of the kernel size and gain corresponding to the mode type notified from the system control unit 13. To do.
  • the interference fringe removal processing unit 100 executes the LPF process S11 using the kernel size and gain notified from the system information processing unit 103. Note that, in a mode in which no interference fringes are generated, the interference fringe removal processing unit 100 may skip the LPF processing S11.
  • the interference fringe removal processing unit 100 executes a difference process S12 for obtaining a difference between the original image data I0 and the LPF processed image data I1. Thereby, interference fringe image data I2 mainly composed of interference fringe components is obtained.
  • the interference fringe removal processing unit 100 After the difference processing S12, the interference fringe removal processing unit 100 performs the LPF processing S13 in the y direction on the interference fringe image data I2.
  • the interference fringe removal processing unit 100 executes a difference process S14 for obtaining a difference between the original image data I0 and the interference fringe image data I3. Thereby, X-ray image data I4 from which interference fringes are removed from the original image data I0 is obtained.
  • the interference fringe removal processing unit 100 outputs the X-ray image data I4 to the noise reduction processing unit 101.
  • the noise reduction processing unit 101 has a plurality of noise reduction processing parts, and the S / N ratio associated with the adoption of the grid 6 according to the present embodiment from the X-ray image data I4 input from the interference fringe removal processing unit 100. Processing for reducing noise caused by the reduction is executed.
  • a “coherent filter” disclosed in Japanese Patent No. 4170767 may be employed as a process for reducing noise.
  • the coherent filter can effectively reduce noise while maintaining the resolution.
  • the local pixels such as 3 ⁇ 3 in the vicinity are weighted and averaged, and the weighted average value is used as the value of the local central pixel. It changes according to the similarity between.
  • the similarity referred to here is an index indicating the degree of possibility that the tissues are close to anatomical, specifically, the brain tissues (capillaries) under the control of the same cerebral artery,
  • processing for reducing noise in the time direction may be performed on the X-ray image data I4.
  • processing for example, a technique disclosed in Japanese Patent Application No. 2011-250066 can be employed.
  • the X-ray image data after the noise is reduced by the noise reduction processing unit 101 is referred to as X-ray image data I5.
  • the noise reduction processing unit 101 outputs the X-ray image data I5 to the scattered radiation correction processing unit 102.
  • the scattered radiation correction processing unit 102 reduces components based on scattered radiation included in the X-ray image data I5.
  • the scattered radiation correction processing unit 102 does not set the component based on the scattered radiation included in the X-ray image data I5 to zero, but selects a target grid (second grid) having higher scattered radiation removal performance than the grid 6. It corrects to the component corresponding to the scattered radiation generated when it is used.
  • p (x, y) is X-ray image data of a direct line generated in the X-ray tube 3 and transmitted through the subject P or the grid 6 and incident on the FPD 7.
  • SPR1 is a value obtained by dividing the scattered dose incident on the FPD 7 when the grid 6 is used by the direct dose incident on the FPD 7 when the grid 6 is used.
  • SPR0 is a value obtained by dividing the scattered dose incident on the FPD 7 when the target grid is used by the direct dose incident on the FPD 7 when the target grid is used.
  • psf1 (x, y) is a point spread function in which the integral value is normalized to 1 with respect to the scattered radiation that passes through the grid 6 and enters the FPD 7.
  • psf0 (x, y) is a point spread function obtained by normalizing the integral value to 1 with respect to the scattered radiation that passes through the target grid and enters the FPD 7.
  • ⁇ is a delta function. Further, “*” represents convolution and “ ⁇ ” represents a product.
  • the target grid is a grid with higher scattered ray removal performance than the grid 6, the relationship 0 ⁇ SPR0 ⁇ SPR1 holds. In other words, the target grid may be selected based on this relationship.
  • SPR0 and SPR1 vary according to tube voltage, irradiation field area, and subject thickness. Therefore, SPR0 and SPR1 are obtained in advance by using a phantom for various tube voltages, irradiation field areas, and subject thickness conditions. Further, the subject thickness can be estimated by an empirical formula using pixel value statistics such as tube voltage, time product of tube current, X-ray focus-X-ray detector distance, set dose, and average pixel value. Therefore, such an empirical formula is determined in advance. With regard to psf0 (x, y) and psf1 (x, y), if at least one is prepared, sufficient accuracy can be corrected. However, a different point spread function may be prepared for each condition similar to SPR0 and SPR1 in order to further improve the correction accuracy.
  • the X-ray image data q0 (x, y) that is the target image (final image) is referred to as a second image.
  • the second image has a scattered radiation correction effect of a certain degree or more.
  • the system information processing unit 103 includes a database composed of SPR0 and SPR1 measured in advance for each tube voltage, irradiation field area, and subject thickness, a time product of tube voltage, tube current, X-ray focus-X-ray.
  • An empirical formula for estimating the object thickness from the inter-detector distance, the set dose, and the statistic of the pixel value, and predetermined PSF0 (u, v) and PSF1 (u, v) are stored. .
  • the system control unit 13 calculates the statistics of the pixel values such as the tube voltage, the time product of the tube current, the distance between the X-ray focus and the X-ray detector, the set dose, and the average pixel value.
  • the image correction unit 10 is notified as system information.
  • the system information processing unit 103 includes a tube voltage, a time product of tube current, an X-ray focal point-X-ray detector distance, a set dose, and a pixel value statistic included in the system information notified from the system control unit 13.
  • the object thickness is estimated using the above empirical formula. Further, the system information processing unit 103 extracts the SPR0 and SPR1 corresponding to the estimated object thickness and the tube voltage and the area of the irradiation field included in the system information notified from the system control unit 13 from the above database. To do.
  • the scattered radiation correction processing unit 102 executes FT processing S21 for obtaining Q1 (u, v) by performing Fourier transform on the X-ray image data q1 (x, y).
  • the scattered radiation correction processing unit 102 obtains SPR0 and SPR1 extracted from the database by the system information processing unit 103, and PSF0 (u, v) and PSF1 (u, u) stored by the system information processing unit 103.
  • An acquisition process S23 for acquiring v) is executed.
  • the scattered radiation correction processing unit 102 uses the coefficients SPR0, PSF0 (u, v) using SPR0, SPR1, PSF0 (u, v), PSF1 (u, v) acquired in the acquisition processes S22 and S23.
  • An arithmetic processing S24 for calculating (+1) / (SPR1 ⁇ PSF1 (u, v) +1) is executed.
  • the scattered radiation correction processing unit 102 executes a calculation process S25 for obtaining a product of Q1 (u, v) obtained in the FT process S21 and a coefficient that is a calculation result of the calculation process S24.
  • the scattered radiation correction processing unit 102 obtains X-ray image data q0 (x, y) by executing an IFT process S26 that performs inverse Fourier transform on the calculation result of the calculation process S25.
  • the X-ray image data q0 (x, y) generated through such correction processing is substantially the same as data obtained when a target grid having a higher scattered ray removal performance than the grid 6 is used. The contrast is improved as compared with the X-ray image data q1 (x, y).
  • the scattered radiation correction processing unit 102 outputs the X-ray image data q 0 (x, y) to the image processing unit 11.
  • the image processing unit 11 performs image processing for display on the X-ray image data q0 (x, y) input from the image correction unit 10, and the processed X-ray image data q0 (x , y) is output to the image display unit 12.
  • the image display unit 12 displays an image based on the X-ray image data q 0 (x, y) input from the image processing unit 11.
  • the image correction unit 10 executes the correction as described above, the interference fringes due to the grid 6 appearing in the X-ray image data are removed, noise appearing in the X-ray image data is reduced, and the contrast of the X-ray image data is improved. can do.
  • processing using wavelet transform disclosed in Japanese Patent Application Laid-Open No. 2011-10829 may be employed.
  • the processing executed by the scattered radiation correction processing unit 102 can be realized by the circuit shown in FIG.
  • This circuit is a modification of FIG. 1 disclosed in Japanese Patent No. 2509181.
  • the two-dimensional memory 201, the scattered radiation response function storage memory 202, the filter coefficient calculation circuit 203, the inverse Fourier transformer 204, the filter calculation circuit 205, the subtractor 206, and the X-ray mount 207 are the two-dimensional memory in FIG. 1 corresponds to a scattered radiation response function storage memory 2, a filter coefficient calculation circuit 3, an inverse Fourier transformer 4, a filter calculation circuit 5, a subtractor 6, and an X-ray mount 7.
  • a scattered radiation elimination ratio calculation circuit 208 and a multiplier 209 are further added.
  • the scattered radiation removal ratio calculation circuit 208 receives conditions such as the tube voltage, the irradiation field area, and the subject thickness from the system information processing unit 103, and calculates the scattered radiation removal ratio corresponding to these conditions.
  • the scattered radiation removal ratio is a coefficient representing how much of the scattered dose when using the grid 6 can be corrected to the scattered dose of the target grid.
  • the scattered radiation removal ratio varies depending on conditions such as tube voltage, irradiation field area, and subject thickness.
  • the arithmetic expression may be set based on experimental results and the like in advance.
  • the scattered radiation removal ratio for each condition such as tube voltage, irradiation field area, and subject thickness is experimentally determined in advance, and the result is stored in the memory of the scattered radiation removal ratio calculation circuit 208 or the like.
  • the line removal ratio calculation circuit 208 may select the scattered radiation removal ratio from this memory.
  • the multiplier 209 corrects the filter coefficient by multiplying the filter coefficient calculated by the filter coefficient calculation circuit 203 by the scattered radiation removal ratio calculated by the scattered radiation removal ratio calculation circuit 208.
  • the inverse Fourier transformer 204, the filter arithmetic circuit 5, and the subtractor 6 perform processing using the filter coefficients after being corrected in this way.
  • the actually collected X-ray image data can be corrected to X-ray image data obtained when a target grid having a non-zero scattered dose is used.
  • the main configuration of the X-ray fluoroscopic apparatus 1 is the same as that shown in FIG. In the present embodiment, the configuration of the image correction unit 10 is different from that of the first embodiment.
  • FIG. 7 is a block diagram illustrating a main configuration of the image correction unit 10 according to the present embodiment.
  • the image correction unit 10 in the present embodiment includes an interference fringe removal processing unit 100, a noise reduction processing unit 101, a scattered radiation correction processing unit 102, and a system information processing unit 103, as well as a decomposition section 110 and a synthesis unit.
  • a section 120 is provided.
  • it referred to X-ray image data inputted from the pixel value calculation unit 9 to the image correction unit 10 and the original image data g 0.
  • the decomposition section 110 includes a front-stage low-pass filter 111-1, a down-sampling processing unit 111-2 for reducing the resolution by down-sampling, an up-sampling processing unit 112-1 for returning the resolution to the original resolution by up-sampling, and a rear-stage low-pass filter 112- 2 and the circuit constituted by the adder 113 are connected from the first stage (the uppermost stage in FIG. 7) to the sixth stage (the lowermost stage in FIG. 7).
  • Filter 111-1 in the first stage performs LPF processing on the original image data g 0.
  • Down-sampling processing unit 111-2 by downsampling the original image data g 0 after LPF processing, to generate a low-resolution image data g 1.
  • Downsampling is performed, for example, by extracting pixels in every other column in every other row from the original image data g 0 after the LPF processing. That is, the low-resolution image data g 1 is 1/4 of the size of the original image data g 0.
  • the first stage of the up-sampling processing section 112-1 complements the "0" in the column and for each row-by-row of the pixels constituting the low-resolution image data g 1.
  • Filter 112-2 performs LPF processing four times each element of the filter 111-1 to the low resolution image data g 1 after the completion.
  • This complementary, low-resolution image data g 1 becomes the original image data g 0 and the same size.
  • the first-stage adder 113 subtracts the low-resolution image data g 1 that has passed through the upsampling processing unit 112-1 and the subsequent-stage low-pass filter 112-2 from the original image data g 0 for each pixel, so that the frequency band data b 0 Is generated.
  • a Gaussian filter of about 5 ⁇ 5 can be employed.
  • the second stage up-sampling processing unit 112-1 and the subsequent low-pass filter 112-2 performs completion and LPF treatment by setting a low-resolution image data g 2.
  • the adder 113 in the second stage generates the frequency band data b 1 by subtracting the low resolution image data g 2 passed through the complementer 112 from the low resolution image data g 1 for each pixel.
  • the low-pass filter 111-1, the down-sampling processing unit 111-2, the up-sampling processing unit 112-1, the subsequent-stage low-pass filter 112-2, and the adder 113 in the third and subsequent stages perform the same processing.
  • low-resolution image data g 3 to g 6 are generated by the low-pass filter 111-1 and the downsampling processing unit 111-2 at each stage, and frequency band data b 2 to b 5 are generated by the adder 113 at each stage. Is done.
  • it referred to the low-resolution image data g 6 of the sixth stage of the low-pass filter 111-1 and the down-sampling processing section 111-2 is generated background data and g 6.
  • the frequency band data b 0 to b 5 indicate data of a plurality of images having different spatial frequency bands.
  • FIG. 9 shows a graph obtained by extracting only the frequency band data b 0 in FIG.
  • the frequency f of the interference fringes is in the range from 1 ⁇ 2 of the Nyquist frequency fa to the Nyquist frequency fa (fa / 2 ⁇ f ⁇ fa). Therefore, in this example, most of the components corresponding to the interference fringes between the grid 6 and the FPD 7 are included in the frequency band data b0. Therefore, the first-stage adder 113 outputs the generated frequency band data b 0 to the interference fringe removal processing unit 100. Further, the interference fringe removal processing unit 100 in this embodiment, the process for removing the interference fringes is performed on the frequency band data b 0. This process may be the process described in the first embodiment, or may be a process using the wavelet transform described in the modification.
  • background data g 6 shown in FIG. 8 shows the point spread function psf1 (x, y) of an example of the relationship between PSF1 obtained by Fourier transform (u, v) in FIG. 10.
  • the sixth-stage reduction filter 111 outputs the generated background data g 6 to the scattered radiation correction processing unit 102.
  • the scattered radiation correction processing unit 102 performs a process for reducing the component based on the scattered radiation relative to the background data g 6. This process may be the process described in the first embodiment or a process using the circuit described in the modification.
  • the background data g 6 after being processed by the scattered radiation correction processing unit 102 is referred to as background data g 6 ′.
  • the second to sixth stage adders 113 output the generated frequency band data b 1 to b 5 to the noise reduction processing unit 101.
  • the interference fringe removal processing unit 100 outputs the frequency band data b 0 after performing the processing for removing the interference fringes to the noise reduction processing unit 101.
  • the noise reduction processing unit 101 performs processing for reducing noise on each of the input frequency band data b 0 to b 5 .
  • a process using the coherent filter described in the first embodiment can be employed.
  • the frequency band data b 0 , b 1 , b 2 , b 3 , b 4 , b 5 after being processed by the noise reduction processing unit 101 are respectively converted into the frequency band data b 0 ′, b 1 ′, b 2 ′. , B 3 ′, b 4 ′, b 5 ′.
  • the synthesizing section 120 includes a circuit composed of the upsampling processing unit 121-1, the low-pass filter 121-1, and the adder 122 from the first stage (the lowest stage in FIG. 7) to the sixth stage (the highest stage in FIG. 7). It is connected over to.
  • the first-stage upsampling processing unit 121-1 and the low-pass filter 121-1 complement “0” for each column and row of the pixels constituting the background data g 6 ′, and the background data after this complementation Apply LPF processing to g 6 ′.
  • the background data g 6 ′ has the same size as the frequency band data b 5 ′.
  • the first-stage adder 122 adds the background data g 6 ′ and the frequency band data b 5 ′ that have passed through the upsampling processing unit 121-1 and the low-pass filter 121-1, for each pixel, thereby adding data g 5 'Occurs.
  • the up-sampling processing unit 121-1 and the low-pass filter 121-1 in the second stage complement “0” for each column and row of the pixels constituting the addition data g 5 ′, and the addition data after this complementation subjected to the LPF processing in g 5 '.
  • the addition data g 5 ′ has the same size as the frequency band data b 4 ′.
  • the adder 122 in the second stage adds the addition data g 5 ′ and the frequency band data b 4 ′ that have passed through the upsampling processing unit 121-1 and the low-pass filter 121-1, for each pixel, thereby adding the addition data g 4 'Occurs.
  • the up-sampling processing unit 121-1, the low-pass filter 121-1, and the adder 122 in the third and subsequent stages perform similar processing.
  • the addition data g 3 ′ to g 0 ′ is sequentially generated by the adders 122 of each stage.
  • X-ray image data g 0' addition data g 0 and.
  • the sixth stage adder 122 outputs the generated X-ray image data g 0 ′ to the image processing unit 11.
  • the image processing unit 11 performs image processing for display on the X-ray image data g 0 ′ input from the image correction unit 10, and the processed X-ray image Data g 0 ′ is output to the image display unit 12.
  • the image display unit 12 displays an image based on the X-ray image data g 0 ′ input from the image processing unit 11.
  • the interference fringes due to the grid 6 appearing in the X-ray image data are removed, noise appearing in the X-ray image data is reduced, and the X-ray image is obtained.
  • Data contrast can be improved.
  • noise can be removed at each resolution level, which leads to a significant improvement in the SN ratio.
  • the number of frequency band data obtained may be more than six layers or less than six layers as long as the purpose of image correction can be achieved.
  • any one of the addition data g 1 ′ to g 5 ′ is a point spread function psf1 ( If the frequency band of x, y) is included, the added data may be subjected to processing for reducing components based on scattered radiation.
  • the interference fringe removal processing unit 100 and the scattered radiation correction processing unit 102 may not be incorporated in the multiresolution analysis. That is, the original image data g 0 from the pixel value calculation unit 9 is first input to the interference fringe removal processing unit 100, and the data after the processing for removing the interference fringes is output to the decomposition section 110 to be combined. The addition data g 0 ′ from the section 120 may be finally input to the scattered radiation correction processing unit 102, and processing for reducing components based on scattered radiation may be performed.
  • SYMBOLS 1 ... X-ray fluoroscopic imaging apparatus, 3 ... X-ray tube, 6 ... Grid, 7 ... FPD, 8 ... AD converter, 9 ... Pixel value calculating part, 10 ... Image correction part, 100 ... Interference fringe removal process part, 101 ... Noise reduction processing section, 102 ... Scattered ray correction processing section, 103 ... System information processing section, S11, S13 ... LPF processing, S12, S14 ... Difference processing, S21 ... FT processing, S22, S23 ... Acquisition processing, S24, S25 ... calculation processing, S26 ... IFT processing.

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  • Health & Medical Sciences (AREA)
  • Life Sciences & Earth Sciences (AREA)
  • Medical Informatics (AREA)
  • Engineering & Computer Science (AREA)
  • Radiology & Medical Imaging (AREA)
  • Molecular Biology (AREA)
  • Biophysics (AREA)
  • Nuclear Medicine, Radiotherapy & Molecular Imaging (AREA)
  • Optics & Photonics (AREA)
  • Pathology (AREA)
  • Physics & Mathematics (AREA)
  • Biomedical Technology (AREA)
  • Heart & Thoracic Surgery (AREA)
  • High Energy & Nuclear Physics (AREA)
  • Surgery (AREA)
  • Animal Behavior & Ethology (AREA)
  • General Health & Medical Sciences (AREA)
  • Public Health (AREA)
  • Veterinary Medicine (AREA)
  • Apparatus For Radiation Diagnosis (AREA)
  • Measurement Of Radiation (AREA)
  • Image Processing (AREA)
  • Image Analysis (AREA)

Abstract

La présente invention porte sur un dispositif de diagnostic par rayons X qui comporte : une source (3) de rayons X ; une grille (6) ; un détecteur (7) de rayons X ; un premier générateur (9) d'image ; un second générateur (110) d'image pour génération de données pour une pluralité de secondes images ayant différentes bandes de fréquence spatiale à partir de données pour une première image ; une pluralité de processeurs (101) de réduction de bruit pour soumettre individuellement la pluralité de secondes images à un traitement de réduction de bruit ; une unité (102) de correction pour soumettre l'une des secondes images à un traitement de correction de faisceau diffusé ; et un générateur (120) d'image finale pour générer une image finale par synthèse des secondes images soumises au traitement de réduction de bruit et de la seconde image soumise au traitement de correction de faisceau diffusé.
PCT/JP2013/080954 2012-11-15 2013-11-15 Dispositif de diagnostic par rayons x WO2014077376A1 (fr)

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JP2012251031 2012-11-15
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JP2013237401A JP6215011B2 (ja) 2012-11-15 2013-11-15 Ｘ線診断装置

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Cited By (1)

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WO2016063452A1 (fr) * 2014-10-21 2016-04-28 Canon Kabushiki Kaisha Appareil de traitement d'image, appareil de capture d'image, système de traitement d'image et processus de traitement d'image

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JP6296553B2 (ja) 2014-09-30 2018-03-20 富士フイルム株式会社 放射線画像撮影装置および放射線画像撮影装置の作動方法
JP6582510B2 (ja) * 2015-04-15 2019-10-02 コニカミノルタ株式会社 放射線画像撮影システム
JP6525772B2 (ja) * 2015-06-30 2019-06-05 キヤノン株式会社 画像処理装置、画像処理方法、放射線撮影システムおよび画像処理プログラム
JP6929343B2 (ja) * 2015-06-30 2021-09-01 キヤノン株式会社 画像処理装置および画像処理方法、画像処理プログラム
EP3403583A1 (fr) * 2017-05-19 2018-11-21 Koninklijke Philips N.V. Mesures géométriques améliorées dans une image à rayons x

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JP2003233818A (ja) * 2002-02-06 2003-08-22 Fuji Photo Film Co Ltd 周期的パターン判定方法および装置並びにプログラム
JP2004261514A (ja) * 2003-03-04 2004-09-24 Fuji Photo Film Co Ltd 画像処理装置、画像処理方法およびプログラム
JP2011010829A (ja) * 2009-07-01 2011-01-20 Toshiba Corp Ｘ線画像撮影装置及び画像処理装置

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JP2003233818A (ja) * 2002-02-06 2003-08-22 Fuji Photo Film Co Ltd 周期的パターン判定方法および装置並びにプログラム
JP2004261514A (ja) * 2003-03-04 2004-09-24 Fuji Photo Film Co Ltd 画像処理装置、画像処理方法およびプログラム
JP2011010829A (ja) * 2009-07-01 2011-01-20 Toshiba Corp Ｘ線画像撮影装置及び画像処理装置

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* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
WO2016063452A1 (fr) * 2014-10-21 2016-04-28 Canon Kabushiki Kaisha Appareil de traitement d'image, appareil de capture d'image, système de traitement d'image et processus de traitement d'image

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