KR20140090831A - THE REDUCTION METHOD OF X-ray SCATTER FROM DIGITAL RADIOGRAPHY IMAGE BY IMAGE PROCESSING AND DIGITAL RADIOGRAPHY SYSTEM USING THE SAME - Google Patents

Abstract

The present invention relates to a method of removing scattered rays, capable of minimizing exposure of a patient to X-rays and obtaining X-rays having excellent quality by removing scattered rays using image processing without using grid, and to a digital radiography system using the same. The method of removing scattered rays according to the present invention includes a first step of imaging a subject with a predetermined X-ray level by the digital radiography system; a second step of acquiring an original image from a detector of the digital radiography system; a third step of calculating a single scatter component as a fixed Gaussian width (σ) according to the X-ray energy level; a fourth step of calculating a scattered ray weight map (σ map) from the original image of the detector; a fifth step of generating a primary X-ray image by removing the scattered rays included in the original image of the detector by performing convolution with respect to the distribution of the single scattered ray to the weight map; and a sixth step of outputting the primary X-ray image from which the scattered rays are removed. Since the scattered rays are removed using image processing without using a grid or air gap, the exposure of a patient is reduced and an X-ray image having excellent quality can be obtained as compared with a case of using the grid.

Description

BACKGROUND OF THE INVENTION 1. Field of the Invention [0001] The present invention relates to a method of removing a scattered ray and a digital radiating system using the same,

The present invention relates to a digital radiation system, and more particularly, to a digital radiation system capable of minimizing radiation exposure to a patient and obtaining a high-quality X-ray image by removing scattering lines using image processing without using a grid And a digital radiation system using the same.

Generally, for medical imaging using X-rays, X-ray energies of 10 keV to 200 keV are used, in which only a part of the energy of an X-ray photon is transmitted as an electron, or a Compton effect, Is electronically transmitted and the object material is imaged through the photoelectric effect in which the X-ray is completely absorbed.

In a digital X-ray imaging apparatus for obtaining an X-ray image, an X-ray passing through a subject impinges on a scintillator to produce a visible ray, which is converted into a CCD (Charge Coupled Device) or a TFT Thin Film Transistor), and direct-type equipment in which an X-ray passing through a subject is directly irradiated to a TFT having a photoconductor (or Photoresistor) Respectively.

Such an X-ray imaging apparatus is irradiated on a large area at a time in a conical shape of an X-ray emitted from an X-ray tube to cause image distortion due to scattering of X-rays, and a method An anti-scatter grid or an air gap as shown in FIG. 1 has been mainly used.

Referring to FIG. 1, in the grid system, the primary radiation is incident almost perpendicularly to the detector (film / detector), whereas the scatter radiation component is incident on the detector in a random direction Point, an anti-scattering grid consisting of lead and aluminum is placed between the patient and the detector so that the X-ray passing through the patient can be lead And physically shields scatter radiation incident obliquely from passing through the grating to prevent it from reaching the detector (film / detector).

The air gap method is a method in which a space is placed between the subject and the detector. Since the traveling direction of the scattered ray is obliquely incident on the detector rather than perpendicular to the detector, if an air gap is provided between the detector and the subject, One of the photons in the first X-ray reaches the detector, while the scattered ray component is scattered around.

However, these conventional methods can reduce the detection efficiency by attenuating the dose reaching the detector due to the loss of X-rays in itself, and as a result, the sufficient signal-to-noise ratio is obtained and the quality of the image reaches a certain level In this case, it is inevitable that the radiation exposure to the patient is increased. Even if the grid is used, there is a problem that the scattering line can not be completely removed. In addition, the conventional method requires a relatively long processing time by performing a plurality of calculations for each thickness, and thus it is difficult to apply to a region having a density or thickness such as a chest.

SUMMARY OF THE INVENTION The present invention has been made to solve the above problems and it is an object of the present invention to provide a method and apparatus for minimizing radiation exposure to a patient by removing scattering lines using image processing without using a grid, And a digital radiation system using the same.

That is, according to the present invention, in a digital radiation system, by removing image scattering lines only by using image processing without using a grid or a separate device, the image quality similar to that of the image captured using the grid is reduced, It is an object of the present invention to provide an improved type digital radiation system capable of reducing the inconvenience of a grid at the same time.

According to an aspect of the present invention, there is provided a system comprising: an X-ray tube generating X-rays at a predetermined energy level; A detector for detecting an X-ray passing through a subject and converting the detected X-ray into an electric image signal; A scattering line weight map (? Map) is obtained from the image signal of the detector to obtain a Gaussian width? Of a single scattering line component depending on an X-ray energy level, a single scattering line component is convoluted to a weight map, A digital image processor for removing a scattering line included in a video signal of the video signal; And output means for outputting an X-ray image from which the scattering line input from the digital image processing unit is removed.

The digital image processing unit includes a look-up table storing a Gaussian width (sigma) value according to an energy level of an X-ray, a sigma value selecting unit for selecting a Gaussian width value according to an X- And generating a weight map (? Map) for the scattered line from the original image; a single scattered ray generating unit for generating a single scattered ray component by calculating a value of the scattered ray from the original image; A scattering line distribution image generating unit for generating a scattering line distribution image by convoluting the weight map and the single scattering line component to generate a scattered ray distribution image; - a scattering line image removing unit for generating a line image, and an image output unit for outputting a primary X-ray image from which scattering lines have been removed.

According to another aspect of the present invention, there is provided a method for acquiring an image of a subject at a predetermined X-ray energy level using a digital radiography system. A second step of obtaining an original image from a detector of the digital radiation system; A third step of calculating a single scattering ray component at a fixed Gaussian width (?) Value according to the X-ray energy level; A fourth step of calculating a scattered ray weight map (? Map) from the original image of the detector; A fifth step of generating a first X-ray image by removing the scattering line included in the original image of the detector by convoluting the single scattering line component with the weight map; And a sixth step of outputting the primary X-ray image from which the scattering line is removed.

에 따라 단일 산란선 성분을 산출한다.In the third step, a Gaussian width (?) According to an X-ray energy level is obtained using a lookup table,

To calculate a single scatter line component.

In the fourth step, a scattered ray weight map is calculated using a lookup table in an original image,

와 수학식

에 따라 역퓨리에변환을 통해 산란선 분포영상을 생성한다.
The fifth step is a step

And Equation

To generate a scattered ray distribution image through inverse Fourier transform.

The scattering line elimination method and the digital radiation system using the same according to the present invention can reduce the dose to the patient by about 40% to 50% by removing the scattering line by using image processing without using a method such as grid or air gap, It is possible to obtain a high-quality X-ray image.

In addition, in the conventional scattering line elimination method using image processing, there is a case in which a change in the scattering line distribution due to a change in X-ray energy to be irradiated is not considered, or a shape or size of a single scattering ray component is different from the actual one , It is difficult to apply to a region having various densities and thicknesses such as a chest, since it requires a relatively long processing time by performing a plurality of calculations per thickness. However, in the present invention, a single scattering line component having a proper shape and sufficient size is used, By omitting the elements and improving the processing speed, it can be applied to real X-ray imaging systems for areas with complex anatomical structures such as the chest.

BRIEF DESCRIPTION OF THE DRAWINGS FIG. 1 is a schematic view showing a concept of removing a scattering sheet using a grid in an X-
FIG. 2 is a block diagram showing the overall configuration of a digital radiation system according to the present invention,
FIG. 3 is a detailed block diagram of scattering line removal shown in FIG. 2,
4 is a flowchart showing an operational procedure of the digital radiation system according to the present invention,
5 is a flowchart illustrating a procedure for removing a scattered ray through image processing according to the present invention.
Figure 6 is an example of a pixel value profile applied to the present invention,
7 shows an example of an acrylic plate used in the experiment of the present invention,
8 shows an example of the beam stop array structure used in the experiments of the present invention,
Figure 9 is an example of an acrylic step edge used in the experiments of the present invention,
10 to 12 show an example of an alpha value lookup table according to the present invention,
13 shows an example of a pixel value profile of a step image according to the present invention,
Fig. 14 is a graph of irradiation dose comparison at the time of grid mounting and grid removal.

BRIEF DESCRIPTION OF THE DRAWINGS The above and other objects, features, and advantages of the present invention will become more apparent from the following detailed description of the present invention when taken in conjunction with the accompanying drawings. The following examples are merely illustrative of the present invention and are not intended to limit the scope of the present invention.

FIG. 2 is a block diagram showing the overall configuration of the digital radiation system according to the present invention, and FIG. 3 is a detailed block diagram of the scattering line removing unit shown in FIG.

2, the digital radiation system 100 according to the present invention includes an X-ray tube 110 for generating X-rays at a predetermined energy level according to a control signal, an X- A controller 130 for controlling an X-ray energy level according to an operation input and providing an X-ray condition to a digital image processing unit, An image input unit 142 for receiving a video signal from the detector 120, a detector correction unit 144 for correcting the map and gain of the detector, and a correction unit 144 for removing the scattered line from the corrected image of the detector, An edge and contrast adjusting unit 148 for adjusting the edge and contrast of the primary X-ray image, and a display unit 150 for displaying the final X-ray image on a monitor . The image input unit 142, the detector correction unit 144, the scattering line removal unit 146 and the edge and contrast adjustment unit 148 are digital image processing units 140 and implemented as software executed by a computer (processor) And the final X-ray image may be output in film form.

In the embodiment of the present invention, the detector 120 uses an amorphous selenium (a-Se) -based direct mode detector, and the maximum expressive power per pixel of the detector is 16383 Lt; / RTI > The size of one pixel is 168um, and the size of the acquired image is 2560x2560 pixels, and the image of 43x43cm area can be obtained. This image is appropriately processed according to each shooting region selected by the user, and then stored as a DICOM file together with the processed image. The detector 120 is mounted on a mechanical structure (for example, a bed, a pedestal, or the like) for supporting a subject according to a subject and a photographing part, and the mechanical structure includes a motor or the like driven under the control of the controller 130 can do.

As shown in FIG. 3, the scattering line removing unit 146, which removes the scattering line by using an image processing method without using a grid according to the present invention, can calculate the Gaussian width of a single scattering line component according to the energy level of the X- a σ value selection unit 146-2 for selecting a Gaussian width value according to the X-ray energy level, and a selected σ value according to a predetermined formula A single scattering line generation unit 146-3 for generating a single scattering line component, an original image input unit 146-4 for receiving an original image of the detector, and an alpha A scattering line distribution image generation unit 146-6 for convoluting a weight map and a single scattering line component to generate a scattering line distribution image, A scattered ray image removing unit 146-7 for generating a primary X-ray image by removing the scattered ray, It consists of the first image output unit (146-8) for outputting an X- ray imaging.

The operation of the digital radiation system according to the present invention will now be described with reference to FIGS. 4 and 5. FIG.

FIG. 4 is a flowchart illustrating an operation procedure of a digital radiation system according to the present invention, and FIG. 5 is a flowchart illustrating a procedure of removing a scattered ray through image processing according to the present invention.

2 and 4, a digital radiation system 100 according to the present invention includes a step S1 of generating a Gaussian width? Lookup table, a step S2 of generating a Gaussian width? (S4) of generating a primary X-ray image by removing a scattered ray from an original image through image processing, and a step (S4) of generating a primary X- , And outputting a primary X-ray image from which scattered rays are removed to a monitor or a film (S5).

5, the step S4 of removing the scattered rays through the image processing according to the present invention includes a step S41 of collecting the condition (energy level) of the X-ray tube, (S42) of calculating a Gaussian width (sigma) value of a single scattered line portion through a lookup table according to a predetermined formula, a step (S43) of generating a single scattered line component by a calculation according to a predetermined equation with the calculated sigma value, (S45) of generating a weight map (? Map) for a scattering line, a step (S45) of receiving an original image of a detector, generating a scattering line distribution image by convoluting a weight map and a single scattering line component A step S47 of generating a primary X-ray image by removing a scattered ray distribution image from the original image, a step S48 of outputting a primary X-ray image from which a scattered ray is removed, .

First, in the present invention, a single scattering line component is modeled as a general Gaussian function as shown in the following Equation 1 and expressed as hs (r). You can change the shape by changing the width of the function by adjusting σ, and r represents the distance from the origin of the polar coordinate system.

The distribution of the scattered line components of the acquired image is the result of the single scattered line components being convolved with respect to the entire image region. In the acquired image, there exists the zero point offset value of the detector itself in the case of the primary X-ray as well as the scattering ray component and the digital radiation system. Figure 6 is a pixel value profile in the presence of scattered lines according to the present invention, where the profile is generally non-planar and curved, and can be seen as a result of single scattered line components being convolved over the entire image. In the present invention, as in the pixel value profile shown in Fig. 6, the portion having the maximum value of the central portion of the image is set as a weight α for the scattering line component, the first X-ray portion is β, the offset of the system itself is γ And weights were calculated based on the SF values measured using a beam stop and used for image processing.

The obtained image can be expressed by the following equation (2), and only the scattering line distribution among these can be expressed by the following equation (3).

In clinical images, α and β are values that change according to position, but they always keep the same value regardless of their position in an object of constant thickness. In other words, the distribution of the scattered rays is a result of convoluting hs with a uniform image having only an alpha value irrespective of the position, and β is added thereto. The result obtained from the actual experiment is a system in which the system function of the detector is convoluted with the energy reaching the detector. In addition, there is always noise generated by various factors. Therefore, the finally obtained image I can be expressed by the following equation (4).

In Equation (4), f represents a function for the system, and sigma represents noise. However, the system function f of the detector does not need to be considered because its width is small, and the noise can be ignored by the curve fitting method. In order to obtain the weight of each component, the relation between? And?,? And SPR is first defined as shown in Equation (5).

The SPR can be thought of as α / β. If the actual measured value is expressed more accurately, the value considered as the primary X-ray component excluding the scattering line includes the offset of the detector system. Since SPR and SF have the same value in all regions in an object of uniform thickness, when M and N are assumed to be the sizes in the x and y directions of the image, respectively, the points of M / 2 and N / 2 If the pixel value is Icenter, Icenter is expressed by the following Equation 6, and by summarizing it,? Is expressed by Equation 7 below.

By the formulas (6) and (7),?, Which is the size of the primary X-ray, is expressed by the following equation (8).

Then,? Is derived as shown in Equation (9) using Equations (6) and (8).

Referring again to FIG. 5, the energy level of the X-ray is received, the corresponding sigma value is obtained from the lookup table, the single scattered ray component is calculated according to Equation (10) after modifying Equation 1 (S41 to S43).

The weight map? (M, n) for the scattered ray component at each position is obtained by obtaining the corresponding value from the lookup table dedicated to the weight map using the pixel value at each position after receiving the original image (S44).

In the step S46 of generating the scattered ray distribution image, an algorithm for obtaining the scattered ray distribution image is designed by partially modifying the conventional inverse Fourier transform method. If the σ value of the single scattering line component hs is fixed, the scattering line distribution image can be calculated by the following calculations.

Here, when m 'and n' are defined as M / 2 and N / 2, respectively, | m'-n | and | n'-n | If the value calculated in Equation (12) is again subjected to inverse Fourier transform, a scattered ray distribution image can be obtained as shown in Equation (13).

When the scattered ray distribution image is obtained as described above, the image in which the scattered ray component is finally removed by removing the scattered ray distribution image from the original image can be obtained as shown in the following Equation 14 (S47).

On the other hand,? (M, n) is a weight for the scattering line component at each position and can be obtained by the above-described method in the single scattering line component measurement method. In the present invention, one of the methods for improving the operation speed of the scattering line component is to change the sigma value of hs only for the irradiated X-ray energy, and the change of the sigma value with respect to the thickness is the difference between the predicted value and the measured value And the error is fixed to a value that minimizes the error. We use the pixel values to generate an array of weights α (m, n) at each position and apply it to the above equation. That is, in the present invention, an image having an alpha value corresponding to each pixel value is generated using an original image, and a predicted image convolved with a suitable scatter point spread function (sPSF) is removed from the original image, .

A Fast Fourier Transform (FFT) algorithm was used to perform the inverse transform and the Fourier transform used in the image processing proposed in the present invention. Since the size of the acquired image is 2560 × 2560 pixels and the size of the image that can be processed is limited to 2n due to the nature of the FFT, it is allocated to the size of 4096 × 4096 pixels to accommodate both the original image and the single scattering line component function array. And the remaining area was zero filled and used for image processing.

[Experimental Example]

Fig. 7 is an example of an acrylic plate used in the experiment of the present invention, Fig. 8 is an example of a beam stop array structure used in the experiment of the present invention, and Fig. 9 is an example of an acrylic step edge used in the experiment of the present invention .

In the experimental example of the present invention, as shown in FIG. 7, a plurality of acrylic plates having a thickness of 20 mm and the beam stop array shown in FIG. 8 were used to measure the thickness gradually. The shape of a single scattering line component was measured by acquiring an image of an acrylic plate having a constant thickness with the grid removed. Weights α and σ of the scattering lines vary depending on the tube voltage and thickness, and the correlation between the pixel values and tube voltage are analyzed. The FOV (field of view) of the image acquisition was made to include the entire area of the acrylic plate, and images were acquired at the acrylic plate thicknesses of 0, 20, 40, 60, 80, 100, 120 and 140 mm. The X-ray was adjusted to 320 mA of tube current and the irradiation time to AEC (automatic exposure control) at 180 cm SID (source to image distance), which is the most commonly used irradiation condition in general chest radiography (Chest PA) 110, and 120kV, respectively. In order to compare the scattered line distribution with the measured values, we obtained the beam stop array for all the conditions except for the images with 0 mm thickness, ie, without any image. It is assumed that the scattering line does not occur at the thickness of 0 mm, and it is used for the purpose of comparing with the pixel value profile in the case of the scattering line.

Obtained pixel value profiles were obtained in the direction perpendicular to the direction of anode placement and cathode placement of the tube, which was less affected by the heel effect. To obtain the center pixel value Icenter and the SF of the central part, the average value of the pixel values for the region of interest (ROI) having a size of 1 mm × 1 mm at the center of the image is obtained, and the average value Respectively. SF was used to calculate the weighting factors α and β for the scattered ray and the primary X-ray, and the correlation between the values and the pixel values was analyzed. The mean square root error rate between the predicted value and the actual value was obtained. Using the obtained α and β values and Equation (2), an It image is created and the pixel value profile is obtained for the same position as the acquired image. Then, the nearest σ value is obtained by using the pixel value profile of the obtained image and the curve fitting method Respectively. In addition, the scattering line measurement values at each position of the beam stop are compared with the scattering line distribution image Is made using α, β, and σ, so that the percentage error (pe) The prediction accuracy was evaluated by measuring the root mean square error (rmspe).

Ibs is the value of the scattered ray component measured at the position where the beam stop is present, and Iconv is the pixel value at the same position in the scattered ray distribution image Is predicted using Equation (3). M and N denote the number of beam stops in the x and y axes, respectively.

On the other hand, in order to quantitatively evaluate the image processing result in the same X-ray irradiation and SID condition as the acrylic plate experiment performed for modeling the single scatter line component, a step image composed of an acrylic plate as shown in FIG. 8 . As shown in FIG. 8, the acrylic step wedge is configured to have a thickness between 60 mm and 140 mm, and each step has a thickness interval of 20 mm. Acrylic step wedge images were acquired for two conditions, with and without grid, respectively. The images captured with the grid and the images without the grid were then quantitatively compared Respectively. To obtain the image of the acrylic step wedge with the grid removed, we compare the error between the scattered line distribution predicted by acquiring a set of images with the beam stop array on the incident surface and the actual measured values Respectively. The method of analyzing the error rate was performed in the same manner as in the experiment for single scattering line modeling. In addition, the dosimetry (Alpha plus, Pehamed, Germany) was placed on the incidence plane of the acrylic step during each experiment and the dose to be irradiated was measured together to compare the case with the case of using the grid. Since the dose varies with the distance from the X-ray tube, the thickness of the acrylic plate 140 mm was measured while maintaining the reference height to maintain the same distance.

[Experiment result]

Table 1 below shows the result of measuring the σ value of the images taken while gradually increasing the thickness of the acrylic plate. For example, when the tube voltage is 110 kV when the thickness is 60 mm, the sigma value obtained through the look-up table is 7.5 cm as shown in Table 1 below. An example of the present invention is an example in which chest radiography is taken as an example, and the value may be changed when photographing another region.

Table 1 shows the correlation coefficient between the measured value and the σ value of each tube voltage and thickness obtained by using the curve fitting. Correlation coefficient shows a value between 0.9 and 0.98, indicating that the fitting was performed well. I could confirm. This value has the same value even if the weights σ and β change. The correlation coefficient σ during curve fitting and σ value of acrylic plate thickness and tube voltage condition increased with increasing thickness and tube voltage, and correlation coefficient ranged from 0.9 to 0.98.

The mean values of SF, SPR, and pixel values were measured and recorded using beam stop at the center of the image to obtain the weights α and β of the scattering line component and the primary X-ray. As a result, as the thickness of the acrylic plate increased, the SF and the SPR showed an increasing tendency as shown in the following Table 2, but they did not show a significant change according to the change of the tube voltage. As the thickness increased, the pixel value tended to decrease as the amount of transmitted X-rays decreased, and the pixel value increased as the tube voltage increased. The weights α and β show a tendency to increase as the tube voltage increases and as the thickness decreases.

Referring to Table 2, pixel values and SF and SPR measurement results for acrylic plate thickness and tube voltage condition. As the thickness of the acrylic plate increases, the pixel value decreases under the same tube voltage condition, but when the tube voltage increases, the pixel value increases. SF and SPR showed an increasing tendency depending on the acrylic plate thickness and were not related to changes in tube voltage

As a result of measurement of the weight values α and β of the acrylic plate thickness and tube voltage conditions, both values decrease with increasing thickness as shown in Table 3, and the value increases with increasing tube voltage.

10 to 12 show an example of an alpha value look-up table according to the tube voltage, wherein FIG. 10 shows a tube voltage of 100 kV, FIG. 11 shows a tube voltage of 110 kV, and FIG. 12 shows an example of a tube voltage of 120 kV. When the pixel value of an image is converted into an alpha value, a pixel value 1000 is set to about 650 by applying a lookup table (Y = 318 ln (x) -1506.9) based on FIG. 11 in the case of an image photographed under the irradiation condition of 110 kV .

FIG. 13 is an example of a pixel value profile of a step image according to the present invention, and FIG. 14 is a graph of irradiated dose when a grid is mounted and when a grid is removed.

Referring to FIG. 13 and FIG. 14, it can be seen that when using the image processing method according to the present invention, it is possible to obtain a high-quality image using a grid while reducing the amount of exposure compared to using a grid.

While the present invention has been particularly shown and described with reference to exemplary embodiments thereof, it is to be understood that the invention is not limited to the disclosed embodiments, but, on the contrary, is intended to cover various modifications and equivalent arrangements included within the spirit and scope of the appended claims.

20: Subject 100: Digital radiation system
110: X-ray tube 120: detector
130: control unit 140: digital image processing unit
142: Image input unit 144: Detector correction unit
146: scattering line rejection 148: edge and contrast adjustment section
150:

Claims (7)

An X-ray tube generating an X-ray with a predetermined energy magnitude;
A detector for detecting an X-ray passing through a subject and converting the detected X-ray into an electric image signal;
A value of a Gaussian width (sigma) of a single scattering line component according to an X-ray energy size is obtained, a scattering line weight map (? Map) is obtained from an image signal of the detector, a single scattering line component is convoluted to a weight map, A digital image processor for removing a scattered ray included in a video signal; And
And output means for outputting an X-ray image from which the scattering line input from the digital image processing unit is removed. The apparatus of claim 1, wherein the digital image processor
A look-up table storing a Gaussian width (sigma) value of a single scattering line component according to an energy level of an X-ray,
A sigma value selection unit for selecting a Gaussian width value according to the X-ray energy size at the time of photographing,
A single scattering line generator for generating a single scattering ray component by calculating a predetermined sigma value according to a predetermined equation,
An original image input unit for receiving an original image of the detector,
An? Image generating unit for generating a weight map? Map for a scattered ray from an original image,
A scattering line distribution image generator for generating a scattering line distribution image by convoluting the weight map and the single scattering line component;
A scattered ray image removing unit for removing the scattered ray distribution image from the original image to generate a primary X-ray image,
And an image output unit for outputting a primary X-ray image from which scattering lines have been removed. A digital radiography system comprising: a first step of capturing an object at a predetermined X-ray energy level;
A second step of obtaining an original image from a detector of the digital radiation system;
A third step of calculating a single scattering ray component at a fixed Gaussian width (?) Value according to the X-ray energy magnitude;
A fourth step of calculating a scattered ray weight map (? Map) from the original image of the detector;
A fifth step of generating a first X-ray image by convoluting the single scattered ray component on the scattered ray weight map and removing a scattered ray included in the original image of the detector; And
And a sixth step of outputting a primary X-ray image from which the scattering line has been removed. 4. The method of claim 3, wherein the third step is to obtain a Gaussian width (sigma) value according to an X-ray energy level using a lookup table. 4. The method of claim 3, wherein the third step comprises:
Equation

And a single scattering line component is calculated according to the scattering line. 4. The method of claim 3, wherein the fourth step
Wherein the scattered ray weight map is calculated using a lookup table in the original image. 4. The method of claim 3, wherein the fifth step
Equation

And Equation

Wherein the scattered ray distribution image is generated through inverse Fourier transform according to the scattered ray distribution image.

Images