JPH07262346A - Aligning method for radiograph - Google Patents

Abstract

PURPOSE:To accurately carryout alignment at high speed. CONSTITUTION:Reference correspondent points 40A-40D are set on a radiograph, and template areas 41A-41D are set with these reference correspondent points as centers. The correspondent points are set by executing template matching in any area on the other radiograph while using these template areas 41A-41D. Next, affin transformation is executed to match the correspondent points with the reference correspondent points 40A-40D, new correspondent points are set by moving the template areas 41A-41D inside any area smaller than the area on the other radiograph and the new correspondent points are matched with the reference correspondent points 40A-40D by the expression of affin transformation so that the position of the radiograph can be matched with that of the other radiograph.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a method for aligning images by correcting positional deviations of a plurality of images which are subjected to a radiation image superimposing process or a subtraction process, and more specifically, to quickly align the images. It's about how to do it.

[0002]

2. Description of the Related Art A stimulable phosphor is used to temporarily record radiation image information of an object such as a human body on a stimulable phosphor sheet (hereinafter referred to as a stimulable phosphor sheet), and this is recorded by excitation light. A method is known in which scanning is performed to cause stimulated emission of light, the stimulated emission of light is photoelectrically read to obtain an image signal, and the image signal is processed to obtain a radiation image of a subject having good diagnostic suitability. This final image can be played as a hard copy or on a CRT.

On the other hand, overlay processing of radiation images has been conventionally known (see, for example, JP-A-56-11399).
Generally, radiation images are used for diagnostic purposes and other purposes, but in using them, it is required to satisfactorily detect minute radiation absorption differences of a subject. The degree of this detection in the radiographic image is called contrast detectability or simply detectability. The higher the detectability, the higher the diagnostic performance.
It can be said that the radiographic image has high practical value. Therefore, in order to improve the diagnostic performance, it is desired to increase this detectability, but the biggest obstacle factor is various noises. The superposition processing is a method of significantly reducing the various noises, making it possible to clearly observe even a slight radiation absorption difference of the subject in the final image, and greatly improving the detectability. That is, a radiation image is captured (accumulated and recorded) on a plurality of stacked stimulable phosphor sheets, and a plurality of image signals obtained by subjecting the plurality of sheets to a reading process are subjected to an addition process. To reduce.

Conventionally, in order to actually carry out this superimposing processing, for example, two stimulable phosphor sheets are put in a cassette in an overlapping manner to photograph an object, and the two stimulable phosphor sheets are attached to the cassette. A method is used in which normal reading processing is sequentially performed to obtain two sets of image signals, and the two sets of image signals are added.

On the other hand, subtraction processing of radiation images has been conventionally known. The subtraction of the radiographic image means that two radiographic images captured under different conditions are photoelectrically read to obtain digital image signals, and then these digital image signals are subjected to subtraction processing by associating each pixel of both images with each other. A method of obtaining a difference signal for extracting a specific structure in a radiation image, and using the difference signal thus obtained, a radiation image in which only the specific structure is extracted can be reproduced.

There are basically the following two methods for this subtraction processing. That is, (1) a specific structure is extracted by subtracting (subtracting) the image signal of the radiation image in which the contrast agent is not injected from the image signal of the radiation image in which the specific structure is emphasized by the injection of the contrast agent. So-called temporal subtraction processing, which (2)
Radiation having different energy distributions is radiated to the same subject, or radiation after passing through the subject is radiated to the two radiation detecting means by changing the energy distribution state, whereby images having different specific structures are displayed. This is a so-called energy subtraction process in which an image of a specific structure is extracted by allowing it to exist between two radiographic images, and then performing an appropriate weighting between the image signals of the two radiographic images and performing subtraction. .

Since this subtraction processing is an extremely effective method especially for medical diagnosis, it has received a great deal of attention in recent years, and its research and development have been actively promoted by making full use of electronic technology.

However, the following problems occur in the radiation image superposition processing method and subtraction processing method using the above-described stimulable phosphor sheet.

That is, in each of the above-described processing methods using the stimulable phosphor sheet, two (or sometimes three or more) stimulable phosphor sheets are sequentially or simultaneously inserted into the photographing table and overlapped with each other. The radiation image to be subtracted is taken, and then the stimulable phosphor sheet is individually inserted into the reading device, and each time the stimulable luminescent light emitted by irradiating the stimulable phosphor sheet with excitation light is detected. The radiographic image is read out by this, but in this process, even if the mechanical accuracy of all devices involved in imaging and reading is increased, positional deviation and rotational deviation occur between the images to be superposed or subtracted. Becomes As a result, although various noises are averaged and reduced in the superposition processing by this processing, blurring occurs in the entire image including the edge portion of the structure in the image, and the image to be observed becomes unsuitable for observation, In addition, in the subtraction process, the image to be erased is not erased, or conversely, the image to be extracted is erased to generate a false image, which makes it impossible to obtain an accurate subtraction image. As described above, it has been found that the positional deviation and the rotational deviation described above cause serious trouble in diagnosis.

When such a deviation occurs between the radiation image information stored and recorded on the stimulable phosphor sheet, the radiation image is stored and recorded in the stimulable phosphor as a latent image.
Unlike the case of an X-ray photo film capable of capturing an X-ray image as a visible image, it is impossible to visually match two X-ray images, and it is extremely difficult to correct the deviation.

Further, even if the positional deviation and the rotational deviation that occur between the two radiation images can be detected by some means, if the conventionally known arithmetic processing is performed to correct the data of the read radiation image, especially the rotational deviation occurs. A great amount of time is spent in the correction, which is a very serious problem in practical use.

Therefore, the applicant of the present invention has proposed a subtraction processing method for a radiation image using a marker having a shape which provides a reference point or a reference line in Japanese Patent Application Laid-Open No. 58-163338. This method records a marker on two stimulable phosphor sheets at a fixed position with respect to the radiographic image, detects the marker when reading the radiographic image,
Positional deviation and rotational deviation are calculated to rotate and / or move one of the deviations of the radiation image to be subtracted on the digital data, and the image data is subtracted between the corresponding pixels of this radiation image. . The alignment step in the subtraction processing method for a radiation image using this marker can also be applied to the above-described overlay processing method. In that case, after the alignment is performed, the addition processing of the image data may be performed between the corresponding pixels of the radiation image.

However, in this method, each time a radiographic image is taken, the above-mentioned marker must be stored and recorded together with the subject in the stimulable phosphor sheet. Then, there is a problem that the image information of the subject cannot be obtained from the portion of the accumulated and recorded radiation image that overlaps the position of the marker.

Therefore, the applicant of the present application has proposed a method of aligning a radiation image without using a marker or the like for alignment (Japanese Patent Application No. 4-318533). In this method, a template region is set in one of the plurality of radiation images to be aligned, and template matching is performed for other radiation images using the template region, and at least two correspondences are made to each radiation image. This is a method in which points are obtained and the corresponding points are repeatedly affine-transformed so that the corresponding points of the radiographic images match, and rotational movement correction, enlargement or reduction ratio correction, and parallel movement correction are performed for each radiographic image.

According to this method, it is possible to perform quick and highly accurate alignment without recording a marker or the like together with the subject for alignment.

[0016]

[Problems to be Solved by the Invention]
With the method described in 4-318533, it is necessary to repeat the template matching for the area in the same range as the area for which the template matching is performed again each time the affine transformation is repeated, and this template matching operation requires time. However, it took a long time to perform the calculation for the alignment.

In view of the above circumstances, it is an object of the present invention to provide a method of aligning a radiation image, which can further shorten the calculation time and perform the alignment quickly.

[0018]

A method of aligning radiation images according to the present invention is a method of aligning a plurality of radiation images for overlay processing or subtraction processing of radiation images. A template region is set on one of the radiographic images and
At least two corresponding points of each of the radiation images are obtained by performing template matching for matching the template region with the other radiation image within a predetermined range of the radiation image other than the one radiation image. , Each of the corresponding points of one radiographic image among the plurality of radiographic images is a reference corresponding point,
formula

[0019]

[Equation 2]

(Where u and v are the coordinates of the reference corresponding point, x and
y is the coordinate of the corresponding point to be converted, a, b, c, d are coefficients indicating rotational movement correction and enlargement or reduction correction, and e, f
Is a coefficient indicating parallel movement correction), and the coefficient is used for the other radiographic image so that the reference corresponding point and the corresponding point of the other radiographic image match. The first affine transformation for converting the coordinate value of the corresponding point into the coordinate value of the reference corresponding point is performed, and the other predetermined radiation image that has undergone the first affine transformation has the template region including the matched region. By performing the template matching again in the area smaller than the area of the range, a new corresponding point corresponding to the reference corresponding point of the one radiation image is obtained in the other radiation image, and the new corresponding point is obtained. Based on the new corresponding points, the coefficient of the affine transformation represented by the above formula is obtained, and the coefficient is used to newly update the reference corresponding point and another radiographic image. It is characterized in that the corresponding point is performed a second affine transformation that transforms the coordinates of the new corresponding points to match the coordinates of the reference corresponding point.

Further, the second affine transformation may be repeated a plurality of times while gradually reducing the region of the predetermined range for each affine transformation.

Here, the template matching is a process in which, when a template region is set on one radiographic image as described above, the template region is moved on another radiographic image to find the best matching place. , The point that represents the location gives the coordinates of the corresponding point.

In such a template matching, the correlation method and the SSDA (Sequential Similarity Detection Algoli) are used as evaluation scales showing the matching degree.
thms).

The correlation method is a method in which a product is calculated for each corresponding pixel and a value obtained by standardizing the sum of the products (hereinafter referred to as a standardized value) is used as a scale of superposition. In this standardization, the sum of the products (squares) of the pixels themselves in each area is calculated, the product of each sum is further calculated, and the square root of this product is calculated as the denominator of the sum of the products for each pixel. It is carried out by When the superposition is perfect, the numerator products are not all sums of squares due to noise, etc.
Therefore, it is considered that the standardized value does not become 1, but becomes the maximum value closest to 1. Therefore, it is considered that the template region is moved variously on the radiographic image and the superposition is achieved with the above-described movement that maximizes the standardized value. However, the movement having the maximum standardized value cannot be determined unless all the movements are completed. Details of this method can be found in, for example, Smith et al., "Automated
cloud tracking using precisely aligned digital ATS
pictures ”ibid. , July 1972, volume c-21, pages 715-729.

Further, SSDA is a measure of superimposing the sum (residual difference) of absolute values of differences for each pixel.
When the superposition is perfect, it is considered that the residual is minimized even if it is not zero due to noise or the like. Therefore, it is considered that the template region is moved variously on the image and the overlay is achieved with the movement that minimizes the residual error. At this time, if the overlay is deviated, the residual error rapidly increases when the pixels are sequentially added. Therefore, if the residual exceeds a certain threshold in the middle of the addition, the method of aborting the addition immediately and moving to the next movement is this SSD.
It is A. Since the calculation used is only addition, and in many cases the calculation is aborted halfway, the calculation time is greatly shortened. For details of this method, see, for example, Barnea et al. "A class.
of algorithms for fast digital image registration
IEEE. Trans. , Vol. C-21, February 1972, pp. 179-186.

[0026]

In the method of aligning the radiation images by repeating the affine transformation described above, the precision of the alignment is improved each time the affine transformation is repeated, so that even if the area for template matching is made small. The alignment accuracy is almost unchanged. In addition, the smaller the area for matching the template, the shorter the calculation time for performing the template matching. The radiation image registration method according to the present invention is made by paying attention to this point, and when performing the second affine transformation,
The range of the region for matching the template region in another radiographic image is made small. Therefore, the calculation time for performing template matching can be shortened, and the calculation time for position alignment can be shortened while ensuring the accuracy of position alignment.

If the above-mentioned second affine transformation is repeated, it is possible to shorten the calculation time for performing template matching each time the affine transformation is repeated, and further improve the alignment accuracy. it can.

[0028]

Embodiments of the present invention will now be described in detail with reference to the drawings.

FIG. 1 is a schematic view of a radiation imaging apparatus which is an embodiment of an apparatus for recording a radiation image used in the present invention. The radiation image obtained by this imaging is used for energy subtraction processing.

The stimulable phosphor sheets 5 and 7 are stacked with the sheet 7 facing down with the filter 6 sandwiched therebetween. On this, a radiation source 2 that emits radiation 3 through a subject 4
Are arranged. The radiation imaging apparatus 1 is configured as described above.

Radiation 3 emitted from the radiation source 2 is applied to the subject 4. The radiation 3a that has passed through the subject 4 is applied to the first stimulable phosphor sheet 5, and a part of the energy of the radiation 3a is recorded in the first stimulable phosphor sheet 5, whereby the sheet 5 is exposed. Radiation images are accumulated and recorded. The radiation 3b transmitted through the sheet 5 is further transmitted through the filter 6, and the radiation 3c transmitted through the filter 6 is second.
The stimulable phosphor sheet 7 is irradiated. As a result, the radiation image of the subject 4 is also stored and recorded on the sheet 7.

FIG. 2 shows each stimulable phosphor sheet 5 and 7.
It is the figure which represented typically the radiographic image accumulated and recorded in. The radiation images 4a and 4b of the subject 4 are accumulated and recorded on substantially the entire surfaces of the stimulable phosphor sheets 5 and 7. That is, the radiation image 4a is obtained from the stimulable phosphor sheet 5 on the upper side,
The radiation image 4b is a radiation image obtained from the lower stimulable phosphor sheet 7.

FIG. 3 shows an example of a radiation image reading apparatus which is an example of a reading unit for reading a radiation image used in the present invention, and an example of an arithmetic unit which executes the alignment method of the present invention and performs subtraction processing. It is a perspective view of an image processing display device.

After radiography is performed by the radiation imaging apparatus 1 shown in FIG. 1, the first and second stimulable phosphor sheets 5 and 7 are set one by one at a predetermined position of the radiation image reading apparatus 10. Here, the case of reading the first radiation image accumulated and recorded in the first stimulable phosphor sheet 5 will be described.

The stimulable phosphor sheet 5 in which the first radiation image is stored and recorded, which is set at a predetermined position, is moved in the direction of arrow Y by the sheet conveying means 15 such as an endless belt driven by a driving means (not shown). It is transported (sub-scan). On the other hand, a light beam 17 emitted from a laser light source 16 is reflected and deflected by a rotating polygon mirror 19 driven by a motor 18 and rotating at a high speed in the arrow Z direction, and after passing through a focusing lens 20 such as an fθ lens, an optical path is passed by a mirror 21. Incident on the stimulable phosphor sheet 5 by changing the direction, and the sub-scanning direction (arrow Y direction)
The main scanning is performed in the direction of the arrow X, which is substantially vertical. From the place where the light beam 17 of the stimulable phosphor sheet 5 is irradiated, stimulated emission light 22 having a light amount corresponding to the stored and recorded radiation image information is emitted. And is photoelectrically detected by a photomultiplier (photomultiplier tube) 24. The light guide 23 is made by molding a light guide material such as an acrylic plate, and is arranged so that the linear incident end face 23a extends along the main scanning line on the stimulable phosphor sheet 5. , Injection end face 23 formed in an annular shape
The light receiving surface of the photomultiplier 24 is coupled to b. The photostimulated luminescent light 22 that has entered the light guide 23 through the incident end face 23a travels through the inside of the light guide 23 after repeating total reflection, and then exits from the exit end face 23b to exit from the photomultiplier 24.
The photostimulated luminescent light 22 representing the radiation image is converted into an electric signal by the photomultiplier 24. The analog signal S output from the photomultiplier 24 is logarithmically amplified by the log amplifier 25, is then input to the A / D converter 26, is sampled, and is a digital image signal SO.
Is obtained. This image signal SO represents the first radiation image accumulated and recorded on the first stimulable phosphor sheet 5, and is referred to as the first image signal SO ₁ . The first image signal SO ₁ is temporarily recorded in the internal memory in the image processing display device 30.

The image processing display device 30 includes a keyboard 31 for inputting various instructions, a CRT display 32 for displaying a visible image based on auxiliary information for the instructions and an image signal,
Floppy disk drive device 33 in which a floppy disk as an auxiliary storage medium is loaded and driven, and a CPU
A main body 34 having a built-in memory and an internal memory is provided.

Next, in the same manner as above, the second image signal SO ₂ representing the second radiation image accumulated and recorded on the second stimulable phosphor sheet 7 is obtained, and the second image signal SO ₂ is obtained.
_{2 is} also temporarily stored in the internal memory in the image processing display device 30.

When the two image signals SO ₁ and SO ₂ to be subjected to the subtraction operation are stored in the internal memory in this manner, these two image signals SO ₁ and SO ₂ are read out and these two images are read. The images are aligned so that a corresponding subtraction operation is performed between the pixels of each radiation image carried by the signals SO ₁ and SO ₂ .

Here, the image signal SO in this embodiment is
_A method of aligning two radiographic images represented by ₁ and SO ₂ will be described.

In the method of aligning a radiation image according to the present invention, the radiation image 4a obtained from the stimulable phosphor sheet 5 located near the radiation source 2 in FIG. 1 is divided into four equal parts as shown in FIG. Four areas 14A, 14B, 14C, 14
Determine D. Then, for each area, an image is filtered using a cross-type filter to detect a point that gives the maximum output value of the filter.

Here, the filter matrix is Aij (i =
1,2, ... n, j = 1,2, ... n), the filter expression is

[0042]

[Equation 3]

The cross-type filter is the element a.
= Ai, j (i = j or i = nj-1), b = Ai, j (i ≠ j and i ≠ nj
-1) means a matrix such that a ≠ b (a> b), and in the present embodiment, for example,

[0044]

[Equation 4]

The following matrix is used.

By using such a cross-type filter, as shown in FIG. 4, a complex portion having a structure such as a cross edge where ribs intersect with each other, that is, the density is extremely high when viewed from the surrounding area. Can be detected, and in this embodiment, four points 40A, 40 can be detected.
B, 40C, 40D can be detected. Although there are a plurality of such cross edges in the four areas 14A to 14D, in the present embodiment, the points 40A, 40B, 40C and 40D at which the output of the filter has the maximum value are detected in each area. And

When the four points 40A to 40D are detected in this way, a template region centered on each point is set on the radiation image 4a. That is, as shown in FIG. 5, template regions 41A to 41D centering on the points 40A to 40D are set in the respective regions 14A to 14D divided into four equal parts.

Next, the template regions 41A to 41D are moved within the regions 43A to 43D divided into four equal parts on the radiation image 4b shown in FIG. 6 to perform template matching. Here, template matching is performed using the above-mentioned correlation method or SSDA. In the correlation method, as described above, the point having the maximum standardized value gives the coordinates of the corresponding points described below. Also in SSDA, as described above, the point where the sum of residuals is the minimum gives the coordinates of the corresponding point.

By performing the template matching in this manner, four corresponding points 44A to 44D are respectively assigned to the four areas 43A to 43D of the radiation image 4b as shown in FIG.
Is required.

Next, each point 40A on the radiation image 4a
-40D is set as the reference corresponding point, and the coordinates of each reference corresponding point are (ui
, Vi) (i = 1 to 4), and affine transformation

[0051]

[Equation 5]

However, m is the enlargement / reduction ratio, θ is the rotation amount,
A and B are the first radiation image by converting the coordinates of each corresponding point (xi, yi) according to the coefficient indicating the rotational movement correction and the enlargement or reduction ratio correction, and C and D according to the coefficient indicating the parallel movement correction. The 4a and the second radiation image 4b are superposed. Here, in the coordinate transformation based on the equation (3), the entire second radiographic image is enlarged or reduced independently in the X direction and the Y direction, the second radiographic image is rotationally moved, and the second radiographic image is rotated. The parallel translation of the radiation image in X direction and Y direction is performed at the same time.

Here, the coefficients A, B, and
A method of obtaining C and D will be described.

In the present invention, the coefficients A, B, C and D are determined by the method of least squares. First, the positional relationship between the reference corresponding point (ui, vi) and the corresponding point (xi, yi) is determined. The error E is represented by E = Σ (u−ui) ² + Σ (v−vi) ² = Σ (A · xi −B · yi + C−ui) ² + Σ (B · xi + A · yi + D−vi) ² ... 4) and solving equation (4) to minimize the squared error of the error E,

[0055]

[Equation 6]

It becomes Therefore, when the linear equation is solved from the equation (5), the coefficients A, B, C and D are obtained as follows.

A = [d · (e + f) −b · i−c · j] / Δ B = [d · (g + h) + c · i−b · j] / Δ C = [− b · (e + f) + c・ (−g + h) + a · i] / ΔD = [− c · (e + f) −b · (−g + h) + a · j] / Δ where Δ = a · d−b ² −c ² a = Σ ( xi ² + yi ² ), b = Σxi, c = Σyi, d = Σ e = Σxi · ui, f = Σyi · vi, g = Σyi · ui, h = Σxi · vi, i = Σui, j = Σvi ... 6) Using the coefficients, A, B, C, and D obtained in this way, coordinate conversion is performed using equation (3),
The second radiographic image 4b can be approximately equalized to the first radiographic image 4a so that the error between A to 44D and the reference corresponding points 41A to 41D is minimized.

Next, with respect to the radiation image 4b thus affine-transformed, the template regions 41A to 41D are moved again on the radiation image 4b in order to perform the second affine transformation for further improving the alignment accuracy. And perform template matching. In this case, template area
The range in which 41A to 41D are moved includes the area matched with the template areas 41A to 41D when the template matching was performed the last time, and is within a narrower area than the areas 43A to 43D divided into four parts of the radiation image 4b. . That is, as shown in FIG. 7, the regions 43A to 43D of the radiographic image 4b.
Template area within the narrower area 45A-45D
Template matching is performed by moving 41A to 41D.

By performing the template matching in the areas 45A to 45D narrower than the areas 43A to 43D in this manner, the calculation time for the template matching is shortened as compared with the case of performing the template matching in the areas 43A to 43D. can do.

By performing this template matching, as shown in FIG. 7, four regions 45A to 45D are formed.
, Four new corresponding points 44A 'to 44D' are obtained for each.

Next, each coefficient of the affine transformation shown in the equation (3) is obtained by the least squares method, and new corresponding points 44A'-44.
By performing coordinate conversion of D ', each corresponding point 44A'-44
The second radiographic image 4b can be matched to the first radiographic image 4a substantially equally so that the error between D'and the reference corresponding points 41A to 41D is further reduced.

Here, the principle of improving the alignment accuracy by repeating the affine transformation a plurality of times will be described with reference to FIGS. 8 and 9.

FIG. 8 is a graph showing the relationship between the relative inclination between a plurality of images and the error between the positions of the corresponding points obtained by template matching and the positions of the true corresponding points. This graph shows that the smaller the inclination of the image, the smaller the positional deviation of the corresponding points. Also,
For example, when the relative inclination of the image is 2 °, it indicates that the error between the corresponding points obtained by the template matching as described above and the true corresponding points is about 1 pixel. The actual deviation between the images is within ± 3 ° because it occurs inside the cassette, the photographing device, and the reading device, but it is confirmed that the relationship shown in FIG. 8 is effective over a range of approximately ± 5 °. ing.

In FIG. 5, the corresponding points are 100 in the X and Y directions.
FIG. 9 is a diagram showing a true corresponding point in the case where the pixels are separated from each other and a position of the corresponding point obtained by template matching. This FIG. 9A shows the true corresponding points, and FIG.
Indicates the corresponding points obtained. The corresponding points shown in FIG. 9B show a state in which the position is displaced by one pixel from the true corresponding point when the relative inclination of the image is 2 ° as described above.

Here, in FIG. 9A, the inclination of the straight line connecting the true corresponding points is θr = tan ⁻¹ (100/100) = 45 °. On the other hand, in FIG. 9B, the slope of the straight line connecting the corresponding points estimated by the template matching is θm = tan ⁻¹ (101/99) = 45.6 ° even in the worst case, which is true. It shows that the maximum deviation from the slope of the line connecting the corresponding points is 0.6 °. That is, it is shown that the relative inclination of the image was originally 2 ° but was improved to 0.6 ° by one affine transformation. Next, when template matching is performed again by the method of the present invention and the second affine transformation is further performed, it can be seen from FIG. 8 that the shift of corresponding points can be accommodated within 0.3 pixels. In this case, the slope of the straight line connecting the corresponding points estimated by the template matching is θm ′ = tan ⁻¹ (100.3 / 99.7) = 45.2 ° even in the worst case, which is the straight line connecting the true corresponding points. The maximum deviation from the tilt is 0.2 °. Ie 2
It is shown that the degree is improved to 0.2 by affine transformation. Therefore, it will be understood that by repeating the affine transformation a plurality of times as described above, the relative deviation regarding the rotational movement of the radiation image is gradually reduced, that is, the alignment accuracy is gradually improved.

Here, the corresponding points are in the X and Y directions.
An example is shown in which the pixels are separated by 100 pixels, but the present invention is not limited to this. As another example, a case will be described in which corresponding points are located at positions separated by 1000 pixels in the X and Y directions and the relative inclination of the image is 2 °. The slope of the straight line connecting the true corresponding points is also 45
°. On the other hand, the slope of the straight line connecting the corresponding points estimated by template matching is θm = tan ⁻¹ (1001/999) = 45.06 ° even in the worst case, which is the same as the straight line connecting the true corresponding points. It shows that the maximum deviation is 0.06 °. That is, it shows that the relative inclination of the image was 2 °, but was improved to 0.06 ° by one affine transformation. From this, it can be seen that the accuracy of alignment improves as the distance between corresponding points increases.

FIG. 10 shows the result of the alignment performed by performing the affine transformation in this way.

As shown in FIG. 10, looking at the data obtained by aligning 55 sets of radiographic images by the alignment method of the present invention, the reference level is | R | = 4.5, max | Ri | =
When set to 1.00, all the data are within the reference level, and are within | R | = 3.15 and max | Ri | = 0.7. Compared with the alignment result by the conventional method shown in FIG. It can be seen that the alignment is performed with high accuracy.

[0069] After this manner affine transformation is performed, subtraction processing, i.e. the image signal after the alignment of the image signal SO ₂ SO ₂ 'when _{the, S1-Wa · SO 1 -Wb} · SO 2' + C (7) where Wa and Wb are weighting coefficients, and C is a bias component.

The weighted subtraction is performed according to the image signal S1 corresponding to the image of the difference between the two radiation images.
Is generated. This image signal S1 is used for the image processing display device 30.
Sent to the CRT display 32, and a visible image (energy subtraction image) based on the image signal S1 is reproduced and displayed on the CRT display 32. The main body
The function (combination of hardware and software) for performing the above subtraction processing executed in 34 is considered as an example of the arithmetic unit of the present invention.

Although the affine transformation is repeated twice in the above-mentioned embodiment, the accuracy of alignment can be further improved by repeating the affine transformation a plurality of times. In this case, for each affine transformation,
The affine transformation is repeated a plurality of times while gradually reducing the region for moving the template region. That is, as shown in FIG. 12, every time the affine transformation is repeated,
The regions 43A to 43D, 45A to 45D, 47A to 47D may be gradually reduced in size. In this case, template matching is performed until the region becomes the size of the template region non-linearly as shown in FIG. 13, for example. The area to be performed may be made smaller. Further, the area for template matching may be linearly reduced.

Further, in the above embodiment, the corresponding points detected by the template matching are affine-transformed from the radiation image obtained from the stimulable phosphor sheet 7 located far from the radiation source 2 in FIG. However, this is due to the following reasons. That is, when a radiation image is stored and recorded on two stimulable phosphor sheets by so-called one shot as shown in FIG. 1, the two sheets 5 and 7 are spatially different from each other with the filter 6 interposed therebetween. Therefore, the distance between the two sheets is different from the radiation source 2 and the subject 4, and therefore the size of the radiation image recorded on each sheet 5 and 7 is different, and the sheets are far from the radiation source 2. The radiation image 4b accumulated and recorded on the sheet 7 at the position becomes an image with more blurring and more scattered rays than the radiation image 4a accumulated and recorded on the sheet 5. Further, when the above-mentioned affine transformation is performed, the image quality of the transformed image deteriorates to some extent because it is necessary to perform some kind of interpolation between pixels. Therefore, a high-quality radiation image 4a
Radiation image rather than affine transformation to deteriorate image quality
Converting the radiation image 4b, which is inferior in image quality to 4a,
As a result, the quality of the subtraction image obtained is guaranteed. Therefore, the radiation image 4b obtained from the stimulable phosphor sheet 7 located far from the radiation source 2 is affine-transformed.

Further, in the above-mentioned embodiment, four corresponding points are set on the radiation image to perform the affine transformation, but the number of these points may be any number as long as it is two or more. The more corresponding points there are, the higher the accuracy of alignment becomes. However, if the number of corresponding points is large, the calculation time becomes long. Therefore, it is preferable to set the number of corresponding points in consideration of the alignment accuracy and the calculation time.

Further, in the above-mentioned embodiment, the two radiation images are aligned for performing the energy subtraction process, but the two radiation images are aligned for performing the superimposing process. You may do it. That is, as shown in FIG. 14, the radiation image of the subject 4 is accumulated and recorded on the two stimulable phosphor sheets 5'and 7'without using the filter 6 in the radiation image capturing apparatus 1 similar to that shown in FIG. Fluorescent phosphor sheet 5 ',
Image signals SO ₁ and SO ₂ representing two radiation images are obtained from 7 ′ by the radiation image reading apparatus shown in FIG. 3, and the radiation images obtained from the two stimulable phosphor sheets 5 ′ and 7 ′ are 'after the alignment is subjected to affine transformation similar to the embodiment described above the radiation image obtained from the superposition processing, i.e. the image signal sO ₂ after alignment of the image signal sO _2' sheet 7 was At this time, S2 = Wc · SO ₁ + Wd · SO ₂ ′ (8) However, Wc and Wd are weighted and added by a weighting coefficient, whereby an image signal S2 corresponding to an image of the sum of two radiation images is obtained. To generate.

Further, in the above-described embodiment, the radiation image is filtered by the cross filter to detect the point that gives the maximum value of the filter output, but this point is, for example, divided. Any point may be used, such as a point representing the maximum value in each area or a point having a certain value or more.

Further, in the above-mentioned embodiment, the alignment of two radiation images is explained, but the number of radiation images to be aligned may be three or more. In this case, of the three or more radiation images, one of the radiation images has the above-mentioned reference corresponding point set, the other radiation images have corresponding points set by template matching, and the corresponding points of each radiation image are set as reference points by affine transformation. It may be adapted to the corresponding points.

Further, in the above-mentioned embodiment, the alignment is performed by the affine transformation shown in the equation (3), but this is a general equation.

[0078]

[Equation 7]

(However, u and v are the coordinates of the reference corresponding point, x and
y is the coordinate of the corresponding point to be converted, a, b, c, d are coefficients indicating rotational movement correction and enlargement or reduction correction, and e, f
Is any coefficient as long as it performs the affine transformation represented by (coefficient indicating parallel movement correction).

[0080]

As described in detail above, the method of aligning a radiation image according to the present invention can reduce the calculation time for performing template matching each time affine transformation is repeated, so that the method is fast and accurate. Good alignment is possible.

[Brief description of drawings]

FIG. 1 is a schematic diagram of a radiographic image recording apparatus that obtains a radiographic image that performs subtraction according to an embodiment of the present invention.

FIG. 2 is a diagram schematically showing a radiation image accumulated and recorded on each stimulable phosphor sheet.

FIG. 3 is a schematic diagram of an image reading apparatus including an apparatus for performing a radiation image registration method according to the present invention.

FIG. 4 is a diagram showing a state in which a radiographic image is divided and reference corresponding points are set.

FIG. 5 is a diagram showing a state in which a template region is set in a radiation image.

FIG. 6 is a diagram showing a state in which corresponding points are set in another radiation image.

FIG. 7 is a diagram showing a state in which a region corresponding to template matching is reduced and template matching is performed to set new corresponding points.

FIG. 8 is a graph showing a relationship between a relative inclination between a plurality of images and a position error of a corresponding point obtained by template matching and a position error of a true corresponding point.

FIG. 9 is a diagram showing the positions of true corresponding points and the positions of corresponding points obtained by template matching when the corresponding points are located 100 pixels apart in the X and Y directions.

FIG. 10 is a diagram showing a result of registration performed by the radiation image registration method according to the present invention.

FIG. 11 is a diagram showing a state in which alignment is performed by a conventional radiation image alignment method.

FIG. 12 is a diagram showing a state in which a region for template matching is gradually reduced.

FIG. 13 is a graph showing the relationship between the size of a region for which template matching is performed and the number of times affine transformation is repeated.

FIG. 14 is a schematic view of a radiographic image recording apparatus for obtaining a radiographic image for superimposing according to an embodiment of the present invention.

[Explanation of symbols]

 1 Radiography apparatus 2 Radiation source 3 Radiation 4 Subject 5 First stimulable phosphor sheet 6 Filter 7 Second stimulable phosphor sheet 10 Radiation image reader 15 Sheet conveying means 16 Laser light source 17 Light beam 18 Motor 19 rotation Polygonal mirror 20 Focusing lens 21 Mirror 22 Excited emission light 23 Light guide 24 Photomultiplier 25 Log amplifier 26 A / D converter 30 Image processing display device 31 Keyboard 32 CRT display 33 Floppy disk drive device 34 Main body 40A-40D Standard Corresponding points 41A to 41D Template area 44A to 44D Corresponding points 43A to 43D, 45A to 45D, 47A to 47D Template matching area

─────────────────────────────────────────────────── ─── Continuation of the front page (51) Int.Cl. ⁶ Identification number Internal reference number FI Technical display location G06T 7/00 7459-5L G06F 15/70 330 P

Claims (2)

[Claims] 1. A method of aligning a plurality of radiation images for overlaying or subtracting the radiation images, wherein a template region is set on one radiation image of the plurality of radiation images, By performing template matching for matching the template region with the other radiation image in a region of a predetermined range in a radiation image other than one radiation image, at least two corresponding points of each radiation image corresponding to each other are identified. Then, each corresponding point of one radiation image among the plurality of radiation images is set as a reference corresponding point, and the following equation is obtained. (However, u and v are the coordinates of the reference corresponding points, x and y are the coordinates of the corresponding points to be converted, a, b, c and d are coefficients indicating rotational movement correction and enlargement or reduction ratio correction, and e and f are The coefficient of the affine transformation represented by (the coefficient indicating the parallel movement correction) is obtained, and the coefficient is used for the other radiographic image so that the reference corresponding point and the corresponding point of the other radiographic image match. First to convert the coordinate value of a point into the coordinate value of the reference corresponding point
Affine transformation of the first affine transformation is performed, and the template matching is performed again in an area smaller than the predetermined area including the area matched with the template area for another radiographic image. According to the above, a new corresponding point corresponding to the reference corresponding point of the one radiographic image in the other radiographic image is obtained, and the coefficient of the affine transformation represented by the above formula based on the new corresponding point obtained by this. And a coordinate value of the new corresponding point is converted into a coordinate value of the reference corresponding point by using the coefficient so that the reference corresponding point and the new corresponding point of another radiographic image match. Alignment method for radiographic images, which is characterized by performing affine transformation of. 2. The method for aligning a radiation image according to claim 1, wherein the second affine transformation is repeated a plurality of times while gradually reducing the region of the predetermined range for each affine transformation. .