JP3347865B2 - How to align radiographic images - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a method for correcting the positional deviation of a plurality of images to be subjected to a superposition process or a subtraction process of a radiographic image, and to align the images. It relates to a method for performing matching.

[0002]

2. Description of the Related Art Using a stimulable phosphor, radiation image information of a subject such as a human body is temporarily recorded on a stimulable phosphor sheet (hereinafter referred to as a stimulable phosphor sheet), and this is recorded with excitation light. 2. Description of the Related Art There is known a method in which a stimulating light is emitted by scanning, and the stimulating 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 back as a hard copy or on a CRT.

On the other hand, superposition processing of radiation images has been conventionally known (see, for example, JP-A-56-11399).
Generally, a radiographic image is used for diagnosis and other purposes, and in using the radiographic image, it is required to detect a minute radiation absorption difference of a subject satisfactorily. The degree of this detection in a radiographic image is referred to as contrast detectability or simply detectability, and the higher the detectability, the higher the diagnostic performance,
It can be said that the radiation image has a high practical value. Therefore, in order to enhance the diagnostic performance, it is desired to increase the detectability, but the biggest obstacle is various noises. The superposition process is a method of greatly 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 (stored and recorded) on a plurality of stimulable phosphor sheets, and a plurality of image signals obtained by subjecting the plurality of sheets to a reading process are added. Decrease.

Conventionally, in order to actually perform this superposition processing, for example, two stimulable phosphor sheets are placed in a cassette in a superimposed state, and an object is photographed. A method of sequentially performing normal reading processing to obtain two sets of image signals and adding the two sets of image signals is used.

On the other hand, subtraction processing of a radiation image has been conventionally known. This subtraction of a radiation image means that two radiation images captured under different conditions are photoelectrically read out to obtain digital image signals, and then these digital image signals are subjected to subtraction processing in correspondence with each pixel of both images. This is a method of obtaining a difference signal for extracting a specific structure in a radiation image. By 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) an image signal of a radiographic image to which a contrast agent is not injected from an image signal of a radiological image in which a specific structure is enhanced by injection of a contrast agent. So-called time subtraction processing, and (2)
The same object is irradiated with radiation having different energy distributions, or the radiation transmitted through the object is irradiated to two radiation detecting means with different energy distribution states, thereby forming an image in which a specific structure is different. This is a so-called energy subtraction process in which an image of a specific structure is extracted by presenting a signal between two radiographic images and then performing appropriate subtraction (subtraction) between the image signals of the two radiographic images and then performing subtraction. .

Since this subtraction processing is a particularly effective method for medical diagnosis, it has attracted much attention in recent years, and its research and development have been actively promoted by making full use of electronic engineering technology.

[0008] However, the following problems occur in the method of superimposing and subtracting radiation images using the stimulable phosphor sheet as described above.

That is, in each of the above-described processing methods using the stimulable phosphor sheets, two (or three or more) stimulable phosphor sheets are sequentially or simultaneously inserted into the photographing table and overlapped or overlapped. A radiation image to be subtracted is taken, and then the stimulable phosphor sheet is individually inserted into the reader, and each time the stimulable phosphor sheet is irradiated with excitation light to detect the stimulated emission light. Radiographic images are read out, but in this process, even if the mechanical accuracy of all the devices related to imaging and reading is increased, there is a positional shift and a rotational shift between the images to be superimposed or subtracted. Becomes As a result, in the superimposition process, although various noises are averaged and reduced by this process, blurring occurs in the entire image including the edges of the structures in the image, and the image to be observed is not suitable for observation, Further, in the subtraction processing, an image to be erased is not erased, or an image to be extracted is erased, and a false image is generated, so that an accurate subtraction image cannot be obtained. As described above, it has been found that the positional deviation and the rotational deviation described above cause serious problems in diagnosis.

If 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 film capable of capturing an X-ray image as a visible image, it is not possible to visually match two X-ray photographs, and it is extremely difficult to correct the displacement.

Further, even if the positional deviation and the rotational deviation occurring between the two radiographic images can be detected by some means, if a conventionally known arithmetic processing is performed to correct the data of the read radiographic image, the rotational deviation is particularly reduced. A great deal of time is spent on correction, which is a very serious problem in practical use.

The applicant of the present invention has proposed in Japanese Patent Application Laid-Open No. 58-163338 a subtraction processing method for a radiographic image using a marker having a shape that provides a reference point or a reference line. In this method, a marker is recorded on two stimulable phosphor sheets at a fixed position with respect to a radiographic image, and the marker is detected when reading the radiographic image.
One of the radiation images to be subtracted is rotated and / or moved on the digital data by calculating the positional deviation and the rotational deviation, and the image data is subtracted between the corresponding pixels of the radiation image. . The positioning process in the subtraction processing method of the radiological image using the marker can also be applied to the above-described superposition processing method. In this case, after the alignment, 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-described marker must be stored and recorded on the stimulable phosphor sheet together with the subject. Then, there is a problem that image information of the subject cannot be obtained from a portion overlapping the position of the marker of the accumulated and recorded radiographic image.

Therefore, the applicant of the present invention has proposed a method of aligning radiographic images 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 a plurality of radiation images to be registered, and template matching is performed on another radiation image using the template region, and at least two corresponding radiation images are assigned to each radiation image. In this method, points are obtained, the corresponding points are affine-transformed so that the corresponding points of the respective radiation images match, and rotational movement correction, enlargement or reduction rate correction, and translation correction are performed on each radiation image.

According to this method, quick and accurate alignment can be performed without recording a marker or the like together with the subject for alignment.

[0016]

However, in the above-described radiographic image alignment method, for example, as shown in FIG. 13, a corresponding point for performing affine transformation may be found near the center of the radiographic image. Thus, even if the corresponding point is obtained near the center of the radiation image, the alignment can be performed with a certain degree of accuracy. For example, when the corresponding point 50 is converted to rotate around the point O shown in FIG. ,
The conversion amount of the point 51 located at a portion distant from the corresponding point 50 is larger than the degree of conversion of the corresponding point 50, and the accuracy of the alignment is reduced. Therefore, it is desirable that the corresponding points are apart from each other in order to improve the accuracy of the alignment of the radiation image.

The present invention has been made in view of the above circumstances, and has as its object to provide a radiation image positioning method capable of more accurately positioning a radiation image.

[0018]

According to the present invention, there is provided a method for aligning a plurality of radiographic images for superimposing or subtracting radiographic images as described above. A plurality of feature points of the one radiation image are extracted from one radiation image of the image, and among the plurality of extracted feature points,
A template region is set around at least two feature points near the edge of the one radiation image and at positions separated from each other, and the template region is matched with another radiation image other than the one radiation image. By performing template matching, at least two corresponding points of each radiation image corresponding to each other are obtained, and the corresponding points of one radiation image in the radiation images are set as reference corresponding points,

[0019]

(Equation 2)

(Where u and v are the coordinates of the reference corresponding point, x,
y is the coordinates of the corresponding point to be transformed, a, b, c, and d are the coefficients indicating the rotational movement correction and the enlargement or reduction ratio correction, and e and f.
Is a coefficient indicating translation correction), and calculates a coefficient of the affine transformation represented by the following equation. Using the coefficient, the coordinate value of the corresponding point is set such that the reference corresponding point coincides with the corresponding point of another radiographic image. An affine transformation for converting the coordinate value of the reference corresponding point into a coordinate value is performed.

Here, a feature point is defined as a point at which one radiation image is filtered by a filter such as a Laplacian filter or a cross-type filter, and a point at which a processing value obtained by the filtering processing becomes a predetermined value or more, a radiation image. A point representing a characteristic part of the radiation image, such as a point having a value equal to or more than a predetermined value among values representing an edge part of the radiation image obtained by the differentiation processing.

The "feature points near the edge of the radiation image" are, for example, a plurality of feature points within a predetermined distance from the end of the radiation image among the feature points extracted by filtering processing or differentiation processing. The feature point having the largest processing value, the feature point having the largest processing value among the several feature points located close to the end of the radiation image, and the feature located closest to the four corners of the radiation image Points and so on.

Further, template matching is a process of, when a template region is set on one radiographic image as described above, moving the template region on another radiographic image to search for the best matching location. The point representing that location gives the coordinates of the corresponding point.

In such template matching, an evaluation scale representing the degree of matching includes a correlation method and an SSDA (Sequential Similarity Detection Algorithm).
thms).

In the correlation method, 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 measure of superposition. In this standardization, the sum of the product (square) of the pixel itself 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 product for each pixel. Is performed. When the superposition is complete, the products of the molecules are not all sums of squares due to noise and the like.
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 was moved variously on the radiographic image, and the superimposition was achieved with the movement that maximizes the above-described standardized value. However, the movement with the maximum standardized value cannot be determined unless all the movements are completed. For details of this method, see, for example, Smith et al., “Automated
cloud tracking using precisely aligned digital ATS
pictures "ibid. , July 1972, c-21, pp. 715-729.

The SSDA uses the sum of the absolute values of the differences (residuals) for each pixel as a measure of superposition.
When the superposition is complete, it is considered that the residual is minimized even if it does not become 0 due to noise or the like. Therefore, it is considered that the superposition has been achieved by moving the template region variously on the image and moving the template region to minimize the residual. At this time, if the superposition is shifted, the residual increases rapidly when the pixels are sequentially added. Therefore, if the residual exceeds a certain threshold value in the middle of the addition, the addition is stopped as soon as possible, and the SSD moves to the next movement.
A. The calculation to be used is only addition, and in many cases is terminated halfway, so that the calculation time is greatly reduced. For details of this method, see, for example, “A class
of algorithms for fast digital image registration
IEEE. Trans. , February 1972, c-21, pp. 179-186.

[0027]

According to the radiographic image registration method of the present invention, a template is selected so as to include feature points near the edge of the radiographic image and apart from each other from among the feature points of the radiographic image for setting the template area. Since an area is set and template matching is performed on another radiographic image to obtain a corresponding point, the corresponding point is located on the radiographic image near the edge of the radiographic image and separated from each other. By affine-transforming the corresponding points at positions apart from each other, an image located between the corresponding points is also accurately converted to a corresponding portion of the radiation image in which the reference point is set.

[0028]

Embodiments of the present invention will be described below 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 radiographic image obtained by this imaging is used for energy subtraction processing.

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

The subject 3 is irradiated with radiation 3 emitted from the radiation source 2. The radiation 3a transmitted 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 on the first stimulable phosphor sheet 5, whereby the sheet 5 A radiation image is stored and recorded. The radiation 3b that has passed through the sheet 5 further passes through the filter 6, and the radiation 3c that has passed through the filter 6 is the second radiation 3b.
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 of the stimulable phosphor sheets 5 and 7.
FIG. 3 is a diagram schematically illustrating a radiation image stored and recorded in the storage device. Radiation images 4a and 4b of the subject 4 are accumulated and recorded on substantially the entire surface of each 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 embodiment of a radiation image reading apparatus which is an embodiment of a reading unit for reading a radiation image used in the present invention, and an embodiment of an arithmetic unit which carries out a subtraction process by executing the alignment method of the present invention. It is a perspective view of an image processing display.

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

The stimulable phosphor sheet 5 set at a predetermined position and on which the first radiation image is stored is recorded in the direction of arrow Y by a sheet conveying means 15 such as an endless belt driven by a driving means (not shown). It is conveyed (sub-scan). On the other hand, a light beam 17 emitted from a laser light source 16 is reflected and deflected by a rotary polygon mirror 19 driven by a motor 18 and rotated at a high speed in the direction of arrow Z, and transmitted through a focusing lens 20 such as an fθ lens. In the direction of sub-scan (arrow Y direction)
The main scanning is performed in the direction of arrow X, which is substantially vertical. From the portion of the stimulable phosphor sheet 5 where the light beam 17 has been irradiated, stimulated emission light 22 having a light amount corresponding to the radiation image information stored and recorded is emitted. And is photoelectrically detected by a photomultiplier (photomultiplier tube) 24. The light guide 23 is formed by molding a light-guiding material such as an acrylic plate, and is arranged such that a linear incident end face 23a extends along a main scanning line on the stimulable phosphor sheet 5. The injection end face 23 formed in an annular shape
The light receiving surface of the photomultiplier 24 is connected to b. The stimulated emission light 22 that has entered the light guide 23 from the incident end face 23a travels inside the light guide 23 by repeating total reflection, exits from the emission end face 23b, and exits from the photomultiplier 24.
The photostimulated luminescence 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 a log amplifier 25, and then input to an A / D converter 26 where it is sampled to obtain a digital image signal SO.
Is obtained. This image signal SO represents the first radiation image stored and recorded on the first stimulable phosphor sheet 5, and is referred to as a _first image signal SO1. This first image signal SO ₁ is temporarily recorded in an 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 auxiliary information for instructions and a visible image based on image signals,
Floppy disk drive 33 loaded with and driven by a floppy disk as an auxiliary storage medium, and CPU
And a main unit 34 having a built-in internal memory.

Next, in the same manner as described above, a second image signal SO ₂ representing the second radiation image stored and recorded on the second stimulable phosphor sheet 7 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, the two image signals SO ₁ and SO ₂ are read out, and these two image signals SO ₁ and SO ₂ are read out. Image positioning is performed so that the corresponding subtraction operation is performed between each pixel 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 radiation images represented by ₁ , SO ₂ will be described.

In the method of aligning a radiation image according to the present invention, a radiation image 4a obtained from the stimulable phosphor sheet 5 located at a position near the radiation source 2 in FIG. 1 is divided into four equal parts as shown in FIG. Four areas 14A, 14B, 14C, 14
D is determined. For each area, the image is filtered using a cross-type filter, and a point at which the output value of the filter is equal to or more than a predetermined value (for example, 0.8) is detected as a feature point.

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

[0042]

(Equation 3)

Where the cross-type filter is composed of the element a
= Ai, j (i = j or i = nj-1), b = Ai, j (i ≠ j and i ≠ n-
j-1) means a matrix such that a ≠ b (a> b). In this embodiment, for example,

[0044]

(Equation 4)

The following matrix is used.

By using such a cross-type filter, a complicated portion having a structure such as a cross edge where a rib intersects with each other, that is, a point where the density extremely changes when viewed from the surrounding area is detected. In this embodiment, it is possible to detect a plurality of feature points like black points shown in FIG.

When a plurality of feature points are detected as described above,
Of these feature points, four feature points closest to the four corners of the radiation image 4a and separated from each other are extracted. This is a radiographic image
The distance between the feature points detected from the four corners of 4a is calculated, and the four feature points having the shortest distance are extracted. In the embodiment according to the present invention, four points 40A, 40B, 40C and 40D can be detected as shown in FIG.

When the four points 40A to 40D are detected in this way, a template region centering on each point is set on the radiation image 4a. That is, as shown in FIG. 6, template regions 41A to 41D centered on points 40A to 40D are set in the regions 14A to 14D divided into four.

Next, template matching is performed by moving each of the template areas 41A to 41D within a predetermined range on the radiation image 4b. Here, the template matching is performed using the above-described correlation method or SSDA. In the correlation method, as described above, the point at which the standardized value is maximum gives the coordinates of the corresponding point described below. Also, SSDA
As described above, the point at which the sum of the residuals becomes the minimum gives the coordinates of the corresponding point.

By performing template matching in this manner, as shown in FIG. 7, four corresponding points 44A to 44D are respectively assigned to the four regions 43A to 43D of the radiation image 4b.
Is required.

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

[0052]

(Equation 5)

Where m is the enlargement / reduction ratio, θ is the amount of rotation,
A and B are the coefficients indicating the rotational movement correction and the enlargement or reduction ratio correction, and C and D are the coefficients indicating the parallel movement correction, by converting the coordinates of each corresponding point (xi, yi) to obtain the first radiation image. 4a and second radiation image 4b
And superimpose. Here, in coordinate transformation based on equation (3),
Enlarging or reducing the entire second radiation image in the X direction and the Y direction independently of each other; rotating and moving the entire second radiation image;
The translation in the Y direction is all performed simultaneously.

Here, the coefficients A, B,
How to obtain C and D will be described.

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

[0056]

(Equation 6)

Is as follows. 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] / Δ D = [- 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 = Σe = Σxi · ui, f = Σyi · vi, g = Σyi · ui, h = Σxi · vi, i = Σui, j = Σvi. 6) By using the coefficients A, B, C, and D obtained in this way and performing coordinate transformation according to equation (3), each corresponding point 44
The second radiation image 4b can be made substantially equal to the first radiation image 4a so that the error between A to 44D and the reference corresponding points 41A to 41D is minimized. That is, as shown in FIG. 8, when reference is made to data obtained by performing registration of 55 sets of radiation images by the registration method of the present invention, the reference level is | R
| = 4.5, max | Ri | = 1.00, all data are within the reference level, and | R | = 3.15, max | R
i | = 0.7, which indicates that the alignment is performed with extremely high accuracy as compared with the alignment result by the conventional method shown in FIG.

[0059] 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 represent weighting coefficients, and C represents a bias.

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

In the above-described embodiment, the expression (2)
The filtering process is performed on the radiation image using the cross-type filter shown in (1), but is not limited to this. When the matrix of the filter is Aij, the element a =
Ai, j (i = (n + 1) / 2 or j = (n + 1) / 2), b = Ai, j
(I ≠ (n + 1) / 2 and j ≠ (n + 1) / 2), then a ≠
A cross-shaped filter where b (a> b) may be used.

As such a cross filter, for example,

[0063]

(Equation 7)

The following matrix can be used.

By using such a cross-shaped filter, it is possible to detect a portion where the lung field intersects with the diaphragm in the chest radiographic image, and the points 45A and 45B shown in FIG. Can be detected as Although a plurality of such points can be extracted from the radiation image 4a, in this embodiment, the points 45A and 45B closest to the end of the radiation image are detected as feature points for setting the template area.

Further, in the above-described embodiment, the radiation image is filtered by a single cross-type filter. However, the radiation image may be filtered by a plurality of types of filters. . For example, in the case of the above-described radiographic image of the chest, in the upper half area of the radiographic image, a cross-type filter shown in Expression (2) is used to detect a point where a shoulder joint or a rib intersects. In half the area,
In order to extract the boundary between the lung field and the diaphragm, a cross-shaped cross-type filter shown in Expression (8) may be used.

In the above-described embodiment, the filters shown in the equations (2) and (8) are used as the cross-type filters. However, the present invention is not limited to this. Any filter can be used as long as the filter is formed.

In the above-described embodiment, the characteristic point of the radiation image is detected by using the cross-type filter. However, any filter such as a Laplacian filter may be used.

Further, in the above-described embodiment, the feature points of the radiation image are detected by the filtering process. However, the present invention is not limited to this. The point at which the density of the edge portion is equal to or higher than a certain value may be obtained as a feature point.

In the above-described embodiment, the template area is set around the feature point closest to the four corners of the radiation image among the feature points. Among the plurality of feature points within the range, the point at which the above-described filtering processing value is the largest, or among several feature points close to the four corners of the radiation image, the above-described processing value is the largest. Any point may be used as long as it is a characteristic point near the edge of the radiation image and at a position apart from each other, such as a point.

Further, in the above-described embodiment, a point where the processing value obtained by the filtering processing is equal to or larger than the threshold value 0.8 is detected as a feature point. However, the threshold value may be any value. As shown in FIG. 11, the threshold value may be set to be larger as the dose of the radiation image is larger. This is because if the radiation dose is low, noise increases and the output value of the filter decreases, making it difficult to find a feature point.

In the above embodiment, the corresponding points detected by 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 for the following reasons. That is, when a radiation image is accumulated and recorded on two stimulable phosphor sheets in a 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. In this case, the two sheets have different distances from the radiation source 2 and the subject 4. Therefore, the dimensions of the radiation images recorded on each of the sheets 5 and 7 are different and are far from the radiation source 2. The radiation image 4b accumulated and recorded on the sheet 7 at the position is an image having more blur and more scattered radiation than the radiation image 4a accumulated and recorded on the sheet 5. In addition, when the above-described affine transformation is performed, the image quality of the converted image deteriorates to some extent because it is necessary to perform some interpolation between pixels. Therefore, the radiation image 4a is better than the affine transformation of the high quality radiation image 4a to deteriorate the image quality.
By converting the radiation image 4b whose image quality is inferior to that of, the image quality of the resulting subtraction image is ensured. Therefore, the radiation image 4b obtained from the stimulable phosphor sheet 7 located far from the radiation source 2 is affine-transformed.

In the above-described embodiment, the affine transformation is performed by setting four corresponding points on the radiation image. However, the number of these points may be any number as long as it is two or more. As the number of corresponding points increases, the accuracy of alignment improves. However, if the number of the corresponding points is large, the calculation time becomes long. Therefore, it is preferable to set the number of the corresponding points in consideration of the alignment accuracy and the calculation time.

Further, in the above-described embodiment, the two radiation images are aligned for performing the energy subtraction processing. However, the two radiation images are aligned for performing the superimposition processing. You may do so. That is, as shown in FIG. 9, 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 of FIG. Phosphor sheet 5 ',
Image signals SO ₁ and SO ₂ representing two radiographic images are obtained from the radiographic image reading device shown in FIG. 3 from 7 ′, and among the radiographic images obtained from the two stimulable phosphor sheets 5 ′ and 7 ′, '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 ′ (9) where Wc and Wd are weighted and added by weighting factors, whereby an image signal S2 corresponding to the sum of two radiation images is obtained. To generate.

Further, in the above-described embodiment, the point at which the maximum value of the filter output is detected is detected by filtering the radiation image by the cross-type filter. Any point may be used, such as a point representing the maximum value in each region, a point equal to or greater than a predetermined value, and the like.

Further, in the above-described embodiment, the description has been given of the positioning of two radiographic images. However, the number of radiographic images to be positioned may be three or more. In this case, among the three or more radiation images, the above-described reference corresponding points are set for one radiation image, the corresponding points are set for the other radiation images by template matching, and the corresponding points of each radiation image are set to the reference points by affine transformation. What is necessary is just to make it correspond to a corresponding point.

Further, in the above-described embodiment, the alignment is performed by the affine transformation shown in the equation (3).

[0078]

(Equation 8)

(Where u and v are the coordinates of the reference corresponding point, x,
y is the coordinates of the corresponding point to be transformed, a, b, c, and d are the coefficients indicating the rotational movement correction and the enlargement or reduction ratio correction, and e and f.
Any coefficient may be used as long as it performs affine transformation represented by (a coefficient indicating translation correction).

[0080]

As described above in detail, in the radiographic image registration method according to the present invention, the corresponding points for performing the affine transformation are determined at positions near and apart from the radiographic image. Since the accuracy of the alignment can be improved as compared with the case where the corresponding points are close to each other, the alignment can be performed more accurately.

[Brief description of the drawings]

FIG. 1 is a schematic diagram of a radiation image recording apparatus for obtaining a subtraction radiation image 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 diagram schematically illustrating 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 plurality of feature points have been obtained on a radiographic image.

FIG. 5 is a diagram showing a state in which a radiation image is divided and reference corresponding points are set.

FIG. 6 is a diagram illustrating a state where a template area is set in a radiographic image.

FIG. 7 is a diagram showing a state where corresponding points are set in another radiation image;

FIG. 8 is a diagram showing a result of alignment performed by the radiographic image alignment method according to the present invention.

FIG. 9 is a diagram showing a state where alignment has been performed by a conventional radiation image alignment method.

FIG. 10 is a diagram showing a state in which feature points are set by another cross-type filter.

FIG. 11 is a diagram showing a relationship between a radiation dose and a threshold.

FIG. 12 is a schematic diagram of a radiation image recording apparatus for obtaining a radiation image to be superimposed according to an embodiment of the present invention

FIG. 13 is a diagram for explaining a positional shift of a radiographic image;

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 Radiation imaging 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 transport means 16 Laser light source 17 Light beam 18 Motor 19 Rotation Polyhedral mirror 20 Focusing lens 21 Mirror 22 Stimulated light 23 Light guide 24 Photomultiplier 25 Log amplifier 26 A / D converter 30 Image processing display 31 Keyboard 32 CRT display 33 Floppy disk drive 34 Main unit 40A to 40D Corresponding points 41A to 41D Template areas 44A to 44D, 50 Corresponding points

──────────────────────────────────────────────────続 き Continuation of the front page (56) References JP-A-59-137942 (JP, A) JP-A-2-100533 (JP, A) JP-A-3-236831 (JP, A) JP-A 61-137 246876 (JP, A) (58) Fields investigated (Int. Cl. ⁷ , DB name) A61B 6/00

Claims (1)

(57) [Claims] 1. A method for performing registration of a plurality of radiation images for a superposition process or a subtraction process of the radiation images, wherein a plurality of feature points of the one radiation image are extracted from one of the radiation images. Extracting a plurality of extracted feature points, setting a template region around at least two feature points near an edge of the one radiation image and at positions separated from each other; By performing template matching for matching to another radiation image other than the one radiation image, at least two corresponding points of each of the radiation images are obtained, and each of the one of the radiation images in the radiation image is determined. Using the corresponding point as a reference corresponding point, (Where u and v are the coordinates of the reference corresponding point, x and y are the coordinates of the corresponding point to be converted, a, b, c, and d are the coefficients indicating the rotational movement correction and the enlargement or reduction rate correction, and e and f are the coefficients. A coefficient of the affine transformation represented by (a coefficient indicating translation correction) is obtained, and the coordinate value of the corresponding point is calculated using the coefficient so that the reference corresponding point and the corresponding point of another radiographic image match. An alignment method for a radiographic image, comprising performing affine transformation for converting into coordinate values of a reference corresponding point.