JP6215011B2 - X-ray diagnostic equipment - Google Patents

Description

  Embodiments described herein relate generally 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 within 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.

  When an FPD (Flat Panel Detector) is used as an X-ray detector, interference fringes between FPD pixels and a grid grating may occur in an image. In order to prevent the generation of the interference fringes, for example, it is necessary to make the grid density of the grid coincide with the pixel pitch of the FPD, or to adopt a grid density that is so high that the input signal cannot be converted by the FPD. Note that the maximum lattice density that can be stably manufactured with the current technology is approximately 80 LP / cm.

  When matching the lattice density to the pixel pitch, very high pitch accuracy is required, so the intermediate material of the grid is limited to aluminum. Further, when a high lattice density such as 80 LP / cm is employed, if the lattice ratio is freely specified to some extent, the intermediate material is still limited to aluminum. In any case, the reason is that 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.

In the case of using a grid having such a limited specification, if the following normal exposure reduction measures are implemented, the image quality deteriorates in any case.・ Reduce the grid ratio.
In this case, since the ratio of scattered rays increases, noise increases and the contrast of the image decreases.・ Reduce X-ray fluoroscopy and imaging dose.
In this case, the signal component decreases and the quantum noise increases.・ Improve X-ray fluoroscopy and radiographic quality (increase tube voltage, thicken quality filter).
In this case, the ratio of scattered radiation increases and the signal component decreases.

  As described above, in an X-ray diagnostic apparatus using an FPD as an X-ray detector, it is difficult to reduce exposure while maintaining image quality.

  As a related technique, there is JP-A-5-244508.

  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 according to the embodiment 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. An image generation unit, a plurality of noise reduction processing units that individually perform noise reduction processing on a plurality of second images, a correction unit that performs scattered radiation correction processing on one of the plurality of second images, and a noise reduction processing And 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 embodiment will be described with reference to the drawings.
In the present embodiment, 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.

(X-ray diagnostic equipment)
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 corresponding 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. For example, the X-ray movable diaphragm 4 has four slidable diaphragm blades, and narrows the X-rays by sliding these diaphragm blades.

  The top 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. However, 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 image processing for display (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 according to 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)
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 the number of lead foils per unit length, and is defined as N = 1 / (D + d) where D is the distance between the lead foils and d is the thickness of the lead foil. When the lattice density N is higher than the Nyquist frequency fa, interference fringes are generated at a frequency f (f = 2 · fa−N) where N is turned back by fa. Therefore, for example, when the pixel size of the FPD 7 is 0.148 mm, the lattice density N is preferably closer to 34 LP / cm within a range of about 34 to 50 LP / cm. By satisfying this condition, processing related to the removal of interference fringes described later is facilitated.

  Lattice ratio r, intermediate material: Direct line transmittance as high as possible. The lattice ratio r is defined as r = h: D, where D is the distance between the lead foils and h is the height of the lead foil. For example, the lattice ratio r is about 6: 1 to 10: 1, and the intermediate material is a fiber.

  Further, 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. Set the dose and quality. For example, when the specification of the grid 6 is r = 6: 1 to 10: 1, N = 34 to 50 LP / cm, and the intermediate material is fiber, the conventional general grid specification is r = 15: 1, N = 80 LP. / Cm and intermediate material = aluminum, the dose of X-rays generated in the X-ray tube 3 is set to about 70 to 90% of the conventional value, and the tube voltage is increased by about 10 to 20 kV compared to the conventional value to harden the radiation quality.

  Adopting the specifications, dose, and radiation quality of the grid 6 as described above can reduce the exposure dose of the subject P, but causes an increase in scattered radiation, resulting in a decrease in SN ratio and a decrease in contrast. Furthermore, it is expected that interference fringes appear in the X-ray image data as in the conventional case.

  Therefore, in the present embodiment, 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.

(Image correction unit)
Details of the image correction unit 10 will be described.
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.

  Hereinafter, processing executed by the interference fringe removal processing unit 100, the noise reduction processing unit 101, and the scattered radiation correction processing unit 102 will be described.

[Interference fringe removal processing section]
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.

  A 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.

  First, 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. The kernel size of the LPF process S11 is, for example, (x, y) = (31, 1). 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. According to the specifications of the grid 6 described above, the frequency f of the interference fringes is in the range from ½ 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. In the LPF process S11, 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. By such LPF processing S11 in the x direction, LPF processed image data I1 from which mainly interference fringe components have been removed is obtained.

  Usually, 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 ½ 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.

  After the LPF process 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.

  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. In this LPF process S13, the interference fringe removal processing unit 100 may use the same kernel size (for example, (x, y) = (1, 31)) and gain as in the LPF process S11, or a frequency different from that of the LPF process S11. For example, a kernel size or gain that passes only lower frequency components may be used. By such LPF processing S13 in the y direction, interference fringe image data I3 representing the interference fringe component more accurately is obtained.

  After the LPF process S13, 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.

  This completes the process for removing the interference fringes. The interference fringe removal processing unit 100 outputs the X-ray image data I4 to the noise reduction processing unit 101.

[Noise reduction processing section]
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.

  As a process for reducing noise, various methods can be employed. For example, as a process for reducing noise, a “coherent filter” disclosed in Japanese Patent No. 4170767 may be employed. The coherent filter can effectively reduce noise while maintaining the resolution. In the coherent filter, 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, brain tissues (capillaries) under the control of the same cerebral artery, By giving a high weight to pixels with a high degree of similarity and conversely giving a low weight close to zero to a pixel with a low degree of similarity, while suppressing noise, the reduction in spatial resolution is suppressed. Making it possible.

  Furthermore, when performing fluoroscopy of the subject P, in addition to the processing related to the “coherent filter”, processing for reducing noise in the time direction may be performed on the X-ray image data I4. As such processing, for example, a technique disclosed in Japanese Patent Application No. 2011-250066 can be adopted.

  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.

[Scattered radiation correction processing section]
The scattered radiation correction processing unit 102 reduces components based on scattered radiation included in the X-ray image data I5.
However, it is generally known that when a strong correction is performed so that the influence of the scattered radiation becomes zero, adverse effects such as all the low pixel value portions in the X-ray image data I5 become zero. Therefore, 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.

First, the theory of correction executed by the scattered radiation correction processing unit will be described.
When the X-ray image data I5 input from the noise reduction processing unit 101 is expressed as q1 (x, y) and the X-ray image data obtained when the target grid is used is expressed as q0 (x, y), Equations (1) and (2) hold.

(1): p (x, y) * (SPR1 · psf1 (x, y) + δ) = q1 (x, y) (2): p (x, y) * (SPR0 · psf0 (x, y) + δ ) = Q0 (x, y)
Here, 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. δ is a delta function. Further, “*” represents convolution and “·” represents a product.

  Since the target grid is a grid with higher scattered ray removal performance than the grid 6, the relationship 0 <SPR0 <SPR1 is established. In other words, the target grid may be selected based on this relationship.

  SPR0 and SPR1 vary depending on the tube voltage, the area of the irradiation field, and the 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. The subject thickness can be estimated by an empirical formula using statistical values of pixel values 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.

  When the above equation (1) is Fourier transformed, the following equation (3) is obtained, and when the equation (2) is Fourier transformed, the following equation (4) is obtained.

(3): P (u, v) · (SPR1 · PSF1 (u, v) +1) = Q1 (u, v)
(4): P (u, v) · (SPR0 · PSF0 (u, v) +1) = Q0 (u, v)
Where P (u, v), PSF1 (u, v), Q1 (u, v), Q0 (u, v), and PSF0 (u, v) are p (x, y) and psf1, respectively. represents the Fourier transform of (x, y), q1 (x, y), q0 (x, y), and psf0 (x, y), u represents the spatial frequency in the x direction, and v represents the spatial frequency in the y direction. Represents.

  From the equations (3) and (4), the following equation (5) is obtained.

(5): Q0 (u, v) = Q1 (u, v). (SPR0.PSF0 (u, v) +1) / (SPR1.PSF1 (u, v) +1)
X-ray image data q0 (x, y), which is a target image (final image), can be obtained by performing inverse Fourier transform on the equation (5).

  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.

Next, the flow of processing related to correction of scattered radiation will be described.
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 and tube current, and 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. .

  With the implementation of fluoroscopy and imaging, the system control unit 13 calculates the statistics of pixel values such as the tube voltage, the time product of tube current, the X-ray focus-X-ray detector distance, 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 tube current time product, 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 flow of processing executed by the scattered radiation correction processing unit 102 is shown in FIG.

  First, 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).

  Further, 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.

  Thereafter, the scattered radiation correction processing unit 102 uses the SPR0, SPR1, PSF0 (u, v), and PSF1 (u, v) acquired in the acquisition processes S22 and S23 to obtain coefficients: (SPR0 · PSF0 (u, v)). An arithmetic processing S24 for calculating (+1) / (SPR1 · PSF1 (u, v) +1) is executed.

  After the calculation process S24, 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.

  Finally, 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 computation result of the computation 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. As described above, 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.

  If 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.

Furthermore, by introducing such correction, it is possible to employ the grid 6 having a specification capable of reducing the exposure dose. (Modification of the first embodiment)
The components disclosed in the first embodiment can be modified as appropriate.

  For example, as processing executed by the interference fringe removal processing unit 100, processing using wavelet transform disclosed in Japanese Patent Application Laid-Open No. 2011-10829 may be employed.

  Further, 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. In this modification, 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. In addition, 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.

Even when the circuit as described above is used, 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. (Second Embodiment)
A second embodiment will be described. The same components as those in the first embodiment are denoted by the same reference numerals, and redundant description is omitted.

  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.
As shown in the figure, 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. In the present embodiment, 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 that reduces the resolution by down-sampling, an up-sampling processing unit 112-1 that returns 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.

  For the LPF processing by the reduction filter 111 and the complementer 112, for example, a Gaussian filter of about 5 × 5 can be employed.

The second-stage low-pass filter 111-1 and the down-sampling processing section 111-2, the LPF processing the low-resolution image data generated in the previous stage instead of the original image data g _0, i.e. a low-resolution image data g ₁ as the target perform downsampling to produce a low-resolution image data g _2. 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. As a result, 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. In the following description, 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.} Frequency band data b ₀ ~b ₅ shows the data of a plurality of images having different spatial frequency bands.

An example of a frequency characteristic graph of the frequency band data _b 0 ~b ₅ and background data _{g 6} shown in FIG. Further, FIG. 9 shows a graph obtained by extracting only the frequency band data b ₀ in FIG. According to the specification of the grid 6 described in the first embodiment, the frequency f of the interference fringes is in the range from ½ 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.

And 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. As shown in the figure, in this embodiment it is assumed that the frequency range of background data g ₆ is covered PSF1 (u, v). Therefore, the sixth-stage reduction filter 111 outputs the generated background data g ₆ to the scattered radiation correction processing unit 102. Furthermore, 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. Hereinafter, 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. Moreover, 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 a process for reducing noise on each of the input frequency band data b _{0 to} b ₅ . For this process, for example, a process using the coherent filter described in the first embodiment can be employed. Hereinafter, 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 ′. By this complementation, 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 second-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 addition data g ₅ ′, and the addition data after this complementation subjected to the LPF processing in g ₅ '. By this complementation, the addition data g ₅ ′ has the same size as the frequency band data b ₄ ′.

The second stage adder 122 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 upsampling processing unit 121-1, the low-pass filter 121-1, and the adder 122 in the third and subsequent stages perform the same processing. As a result, the addition data g ₃ ′ to g ₀ ′ are sequentially generated by the adders 122 of each stage. Hereinafter referred 'the 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. As described in the first embodiment, 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.

According to the configuration of the present embodiment described above, similarly to the first embodiment, 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. In particular, if multi-resolution analysis as in the present embodiment is used, noise can be removed at each resolution level, which leads to a significant improvement in the SN ratio. (Modification of the second embodiment)
The components disclosed in the second embodiment can be modified as appropriate.

  For example, the number of obtaining frequency band data may be more than 6 layers or less than 6 layers as long as the purpose of image correction can be achieved.

Further, although the present embodiment illustrates the case of applying the process for reducing the component based on the scattered radiation relative to the background data g _6, the sum data g _{1 '~g} 5' or the point spread function of 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.

Furthermore, 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.

Although several embodiments of the present invention have been described, these embodiments are presented by way of example and are not intended to limit the scope of the invention. These novel embodiments can be implemented in various other forms, and various omissions, replacements, and changes can be made without departing from the scope of the invention. These embodiments and modifications thereof are included in the scope and gist of the invention, and are included in the invention described in the claims and the equivalents thereof.
Hereinafter, the invention described in the scope of claims of the present application will be appended.
[1] An X-ray source that generates X-rays, a grid that removes scattered rays from X-rays transmitted through a subject, an X-ray detector that detects X-rays transmitted through the grid, and the X-ray detector A first image generating unit that generates data of a first image based on the detection result of the second, and a second that generates data of a plurality of low resolution images having a lower resolution than the first image from the data of the first image At least one second image corresponding to a target grid having a second ratio that is lower than a first ratio of scattered dose to a direct dose specific to the grid, wherein at least one of the image generator and the data of the low resolution image An X-ray diagnostic apparatus comprising a correction unit that corrects the data.
[2] A frequency band data generation unit that generates a plurality of frequency band data based on the data of the plurality of low resolution images, and a noise reduction processing unit that reduces noise included in the plurality of frequency band data. The X-ray diagnostic apparatus according to Supplementary Note [1], comprising:
[3] The supplementary note, further comprising a third image generating unit that generates data of a third image by combining the plurality of frequency band data with reduced noise and the data of the second image. 2].
[4] The correction unit performs, on the data obtained by Fourier transforming the second image data, the ratio relating to the first grid and the second grid and the point spread of the scattered radiation that passes through the first grid and the second grid. The X-ray diagnostic apparatus according to [1], wherein the low-resolution image is corrected by performing inverse Fourier transform on data obtained by multiplying a coefficient defined using a function.
[5] The correction unit adds the ratio of the first grid and the second grid to the data obtained by Fourier transform of the second image data, and the point spread of scattered rays that pass through the first grid and the second grid. The data obtained by multiplying a coefficient defined by using a function is subjected to inverse Fourier transform to correct the data of the second image corresponding to the target grid. Line diagnostic equipment.
[6] The X-ray according to [5], wherein the coefficient is determined based on a tube voltage applied to the X-ray source, an irradiation field of the X-ray, and a body thickness of the subject. Diagnostic device.
[7] An X-ray source that generates X-rays, a grid that removes scattered rays from X-rays transmitted through a subject, an X-ray detector that detects X-rays transmitted through the grid, and the X-ray detector Based on the detection result of the above, the ratio of the scattered dose to the direct dose included in the transmitted X-ray is smaller than the ratio in the grid from the image generation unit that generates the data of the first image and the data of the first image An X-ray diagnostic apparatus comprising: a correction unit that corrects second image data corresponding to a target grid that is a value.
[8] The correction unit performs, on the data obtained by Fourier transforming the first image data, the ratio relating to the first grid and the second grid, and the point spread of scattered rays that pass through the first grid and the second grid. The X-ray diagnostic apparatus according to appendix [7], wherein the second image data is obtained by performing inverse Fourier transform on data obtained by multiplying a coefficient defined using a function.
[9] The X-ray according to [8], wherein the coefficient is determined based on a tube voltage applied to the X-ray source, an irradiation field of the X-ray, and a thickness of the subject. Diagnostic device.
[10] The image processing apparatus further includes a removal processing unit that removes interference fringes generated due to the grid from the first image data, and the correction unit performs the first processing after the interference fringes are removed by the removal processing unit. The X-ray diagnostic apparatus according to [9], wherein image data is corrected to the second image data.
[11] The removal processing unit performs low-pass filter processing on the first image data with respect to the direction in which interference fringes are arranged, and subtracts the image data obtained thereby from the first image data, thereby obtaining the interference fringe image data. The obtained image data of the interference fringes is subjected to low-pass filter processing in the direction perpendicular to the arrangement direction of the interference fringes, and the image data obtained thereby is subtracted from the first image data, whereby the first image data is subtracted from the first image data. The X-ray diagnostic apparatus according to Appendix [10], wherein the interference fringes are removed.
[12] The image processing apparatus further includes a noise reduction processing unit that reduces noise included in the first image data, and the correction unit converts the first image data after the noise is reduced by the noise reduction processing unit to the second image data. The X-ray diagnostic apparatus according to Appendix [11], wherein image data is corrected.
[13] An X-ray source for generating X-rays, a grid for removing scattered rays from X-rays transmitted through a subject, an X-ray detector for detecting X-rays transmitted through the grid, and the X-ray detector Based on the detection result, an image generation unit that generates first image data, a low-resolution image generation unit that generates a plurality of low-resolution image data by down-sampling the first image data, and the low-resolution image data A correction unit that corrects the second image data corresponding to the target grid in which the ratio of the scattered dose to the direct dose contained in the transmitted X-ray is smaller than the ratio in the grid, X-ray diagnostic equipment.
[14] A frequency band data generation unit that generates a plurality of frequency band data based on the plurality of low resolution image data, and a noise reduction processing unit that reduces noise included in the plurality of frequency band data. The X-ray diagnostic apparatus according to Supplementary Note [7], wherein
[15] An X-ray source for generating X-rays, a grid for removing scattered rays from X-rays transmitted through the subject, an X-ray detector for detecting X-rays transmitted through the grid, and the X-ray detector A first image generation unit that generates data of the first image based on the detection result of the first, a plurality of frequency band data having different spatial frequency bands from the data of the first image, and a second low-resolution background data A second image generating unit that generates an image, a plurality of noise reduction processing units that individually perform noise reduction processing on the plurality of frequency band data, and a correction unit that performs scattered radiation correction processing on the second image And a final image generating unit that generates a final image by combining the plurality of frequency band data subjected to the noise reduction processing and the second image subjected to the scattered radiation correction processing. Characteristic X-ray diagnosis Location.
[16] The second image generation unit includes a front-stage low-pass filter for passing a low spatial frequency component of the input image, a down-sampling processing unit for reducing the resolution of the output image of the front-stage low-pass filter, and the down-sampling unit An upsampling processing unit for increasing the resolution of the output image of the sampling processing unit, a post-stage low-pass filter for passing a low spatial frequency component of the output image of the up-sampling processing unit, and a post-stage low-pass filter from the input image The X-ray diagnostic apparatus according to Supplementary Note [15], comprising a plurality of sets each including a difference processing unit for differentiating output images.
[17] The X-ray diagnostic apparatus according to [15], wherein the plurality of noise reduction processing units are a plurality of coherent filters having different filter intensities.
[18] The X-ray diagnostic apparatus according to [15], wherein the correction unit reduces a scattered radiation component in a frequency space with respect to the plurality of second images.
[19] The correction unit includes a ratio of a scattered dose to a direct dose inherent to the grid, a ratio related to a target grid having higher performance of removing scattered rays than the grid, a point spread function related to scattered rays passing through the grid, The X-ray diagnostic apparatus according to appendix [18], wherein the scattered radiation component is reduced based on a point spread function relating to the scattered radiation passing through the target grid.

  DESCRIPTION OF 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.

Claims (18)

An X-ray source generating X-rays;
A grid that removes scattered radiation from X-rays transmitted through the subject;
An X-ray detector for detecting X-rays transmitted through the grid;
A first image generator for generating data of a first image based on a detection result of the X-ray detector;
A second image generating unit for generating a plurality of low resolution image data having a lower resolution than the first image from the data of the first image;
At least one of the low resolution image data is corrected to at least one second image data corresponding to a target grid having a second ratio that is lower than a first ratio of scattered dose to direct dose inherent to the grid. An X-ray diagnostic apparatus comprising a correction unit. A frequency band data generator for generating a plurality of frequency band data based on the data of the plurality of low resolution images;
The X-ray diagnostic apparatus according to claim 1, further comprising: a noise reduction processing unit that reduces noise included in the plurality of frequency band data.   3. The image processing apparatus according to claim 2, further comprising a third image generation unit configured to generate data of a third image by combining the plurality of frequency band data with reduced noise and the data of the second image. X-ray diagnostic equipment. Wherein the correction unit, the at least one data of the low-resolution image in the Fourier transformed data, the related prior Kigu lid and the target grid first ratio and the second ratio and the front Kigu lid and the target grid 2. The data of the second image is obtained by performing inverse Fourier transform on data obtained by multiplying a coefficient defined using a point spread function of transmitted scattered radiation. X-ray diagnostic equipment. Wherein the first ratio and the second ratio, according to claim 4, wherein the tube voltage applied to the X-ray source, it is determined based on the body thickness of the irradiation field and the object of the X-ray X-ray diagnostic equipment. An X-ray source generating X-rays;
A grid that removes scattered radiation from X-rays transmitted through the subject;
An X-ray detector for detecting X-rays transmitted through the grid;
An image generation unit that generates data of a first image based on a detection result of the X-ray detector;
A correction unit that corrects the data of the first image to the data of the second image corresponding to the target grid in which the ratio of the scattered dose to the direct dose included in the transmitted X-ray is smaller than the ratio in the grid; ,
An X-ray diagnostic apparatus comprising: Wherein the correction unit uses the said data of the first image in the Fourier transform data, point spread function of scattered rays passing through the said ratio before Kigu lid and the target grid for the previous Kigu lid and the target grid by inverse Fourier transform of the data obtained by multiplying the coefficient defined Te, X-rays diagnostic apparatus according to claim 6, characterized in that for obtaining the data of the second image. The ratio, tube voltage applied to the X-ray source, the radiation field of the X-ray, and the test of the body thickness X-ray diagnostic apparatus according to claim 7, characterized in that it is determined based on . A removal processing unit for removing interference fringes caused by the grid from the data of the first image;
Wherein the correction unit, X-rays diagnostic apparatus according to claim 8, characterized in that to correct the data of the first image after the interference fringes are removed by the removing unit into data of the second image . The removal processing unit, said performing low-pass filtering process on the data array direction to the interference fringes of the first image to obtain the image data of the interference fringes by subtracting the image data obtained by this from the data of the first image by applying a low-pass filter processing with respect to the arrangement direction and vertical direction of the interference fringe on the image data of the interference fringes obtained, subtracting the image data obtained by this from the data of the first image data of said first image The X-ray diagnostic apparatus according to claim 9 , wherein the interference fringes are removed from the X-ray diagnostic apparatus. A noise reduction processing unit for reducing noise included in the data of the first image;
Wherein the correction unit, X-rays diagnostic apparatus according to claim 10, characterized in that to correct the data of the first image after the noise has been reduced by the noise reduction processing unit to the data of the second image. An X-ray source generating X-rays;
A grid for removing scattered rays from the X-ray transmitted through the inspected object,
An X-ray detector for detecting X-rays transmitted through the grid;
Based on the detection result of the X-ray detector, an image generation unit for generating data of the first image,
A low-resolution image generation unit for generating data of a plurality of low-resolution image by down-sampling the data of the first image,
The relative data of the low-resolution image, the correction unit for correcting the data of the second image the ratio of scattered dose to direct the dose contained in the X-rays transmitted through the corresponding to the target grid of a smaller value than the ratio in the grid When,
An X-ray diagnostic apparatus comprising: A frequency band data generation unit that generates a plurality of frequency band data based on the data of the plurality of low resolution images;
The X-ray diagnosis apparatus according to claim 12 , further comprising a noise reduction processing unit that reduces noise included in the plurality of frequency band data. An X-ray source generating X-rays;
A grid that removes scattered radiation from X-rays transmitted through the subject;
An X-ray detector for detecting X-rays transmitted through the grid;
A first image generator for generating data of a first image based on a detection result of the X-ray detector;
A second image generation unit that generates a plurality of frequency band data having different spatial frequency bands and a second image that is one low-resolution background data from the first image data;
A plurality of noise reduction processing units that individually perform noise reduction processing on the plurality of frequency band data;
A correction unit that performs a scattered radiation correction process on the second image;
A final image generation unit configured to generate a final image by combining the plurality of frequency band data subjected to the noise reduction processing and the second image subjected to the scattered radiation correction processing; X-ray diagnostic equipment. The second image generation unit includes a front-stage low-pass filter for passing a low spatial frequency component of the input image, a down-sampling processing unit for reducing the resolution of the output image of the front-stage low-pass filter, and the down-sampling processing unit An up-sampling processing unit for increasing the resolution of the output image, a rear-stage low-pass filter for passing a low spatial frequency component of the output image of the up-sampling processing unit, and an output image of the rear-stage low-pass filter from the input image The X-ray diagnostic apparatus according to claim 14, comprising a plurality of sets each including a difference processing unit that performs a difference. The X-ray diagnostic apparatus according to claim 14, wherein the plurality of noise reduction processing units are a plurality of coherent filters having different filter strengths. The correction unit before Symbol X-ray diagnostic apparatus according to claim 14, wherein reducing the scattered radiation component in the frequency space to the second image. The correction unit includes a ratio of a scattered dose to a direct dose inherent to the grid, a ratio related to a target grid having higher performance of removing scattered rays than the grid, a point spread function related to scattered rays passing through the grid, and the target grid. The X-ray diagnostic apparatus according to claim 17, wherein the scattered radiation component is reduced based on a point spread function relating to the scattered radiation passing therethrough.

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