JP2981694B2 - Energy subtraction image generation method - 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 of generating a subtracted image over the entire image more accurately, and more particularly, to a method of correcting the spatial frequency of an original image so as to be independent of a place. The present invention relates to a method for generating an energy subtraction image using a constant energy subtraction coefficient.

[0002]

2. Description of the Related Art In various fields, a recorded radiation image is read to obtain image data, and after appropriate image processing is performed on the image data, the image is reproduced and recorded. For example, an X-ray image is recorded using an X-ray film having a low gamma value designed to be compatible with the subsequent image processing, and the X-ray image is read from the film on which the X-ray image is recorded, and is converted into an electric signal. After conversion, the electric signal (image data) is subjected to image processing, and then reproduced as a visible image in a copy photograph or the like, a reproduced image with good image quality performance such as contrast, sharpness, and graininess can be obtained. A system has been developed (see Japanese Patent Publication No. 61-5193).

[0003] Further, according to the present applicant, radiation (X-ray, α
Radiation, β-rays, γ-rays, electron beams, ultraviolet rays, etc.), a part of this radiation energy is accumulated, and then, when irradiated with excitation light such as visible light, the amount of stimulated emission light corresponding to the accumulated energy is increased. Phosphor that emits light (stimulable phosphor)
Utilizing, a radiation image of a subject such as a human body is once photographed and recorded on a sheet-shaped stimulable phosphor, and the stimulable phosphor sheet is scanned with excitation light such as laser light to generate stimulated emission light,
A radiation recording / reproducing system for photoelectrically reading the obtained stimulated emission light to obtain an image signal, and outputting a radiation image of the subject as a visible image to a recording material such as a photographic photosensitive material or a CRT based on the image signal. Have already been proposed (JP-A-55-12429, JP-A-56-11395, JP-A-55-0163472,
Nos. 56-164645 and 55-116340).

This system has the practical advantage of being able to record images over a very wide radiation exposure area as compared to conventional radiographic systems using silver halide photography. That is, it has been recognized that the amount of stimulating light emitted by excitation after accumulation with respect to the amount of radiation exposure is proportional over a very wide range, and thus the amount of radiation exposure varies considerably depending on various imaging conditions. However, the stimulated emission light emitted from the stimulable phosphor sheet is read, the gain is set to an appropriate value, read by photoelectric conversion means, and converted into an electric signal (image data). By outputting a radiation image as a visible image to a display device such as a photosensitive material or a CRT, a radiation image which is not affected by a change in radiation exposure amount can be obtained.

In a system using an X-ray film, a stimulable phosphor sheet, or the like as described above, a plurality of recorded radiation images are read to obtain a plurality of image data, and based on the image data, the radiation image is obtained. Image subtraction processing may be performed.

Here, the subtraction processing of the radiation images refers to processing for obtaining an image corresponding to the difference between a plurality of radiation images taken under different conditions. By reading at a sampling interval, obtaining a plurality of digital image signals corresponding to each radiation image, and performing a subtraction process for each corresponding sampling point of the plurality of digital image signals,
It refers to a process of obtaining a radiation image in which only a specific subject portion in the radiation image is emphasized or extracted.

This subtraction processing basically includes the following two. That is, the subtraction (subtract) of a radiographic image in which a contrast agent is not injected is performed by subtracting a radiographic image in which a contrast agent is not injected from a radiographic image in which a specific portion of a subject (for example, a blood vessel when a human body is set as a subject) is enhanced by injection of a contrast agent. Using the so-called time subtraction to extract a specific part (for example, a blood vessel) and the fact that the specific part of the subject has different radiation absorptivity for radiation having different energies, the same subject can be compared with each other. By irradiating radiations having different energies, a plurality of radiation images of each radiation having different energies are obtained, and a specific portion of the subject is extracted by appropriately weighting the plurality of radiation images and calculating a difference therebetween. There is so-called energy subtraction. Further, the energy subtraction processing includes the following two types. One is a so-called one-shot energy subtraction in which the energy distribution of radiation transmitted through a subject is changed and a subtraction is performed between a radiation image before the change and a radiation image after the change, and the other is a one-shot energy subtraction. This is a so-called two-shot energy subtraction in which subtraction is performed between radiation images of a plurality of radiations, which have transmitted through the subject separately and have different energy distributions. For example, in the former method, at least two stimulable phosphor sheets are laminated with a radiation separating filter such as a copper plate interposed therebetween, or at least two stimulable phosphor sheets having different radiation absorption characteristics are laminated to perform imaging. The different methods are performed by recording different radiation images on the stimulable phosphor sheet. In the latter method, the stimulable phosphor sheet at the imaging position is replaced at a high speed, and high-energy and low-energy radiation is applied to the subject with an X-ray tube. Irradiation is performed by switching the tube voltage at high speed, and a radiation image by each radiation is recorded on each sheet. The present applicant has also proposed energy subtraction using a stimulable phosphor sheet (JP-A-59-834).
No. 86, JP-A-60-225541).

[0008]

With respect to the above-described energy subtraction processing, when radiation having a predetermined energy distribution is applied to the subject when performing radiation imaging of the subject, the radiation transmittance differs for each part of the subject.
In addition, since the transmittance of the lower-energy radiation is lower, a phenomenon called beam hardening occurs in which the energy distribution of the radiation as a whole shifts toward the higher energy side as the radiation passes through the subject, and the degree of the shift is determined by the degree of the shift of the subject. It will be different for each part.

On the other hand, in the energy subtraction processing, an image in which a shadow of a desired tissue of the subject is extracted or emphasized using the difference in the radiation transmittance of each tissue constituting the subject to radiation having different energies. Therefore, if beam hardening of different degrees occurs in each part of the subject, unnecessary tissues are clearly erased in a certain area of the image, and only desired tissues can be extracted. However, a phenomenon occurs in which unnecessary tissue shadows remain in other partial regions of the image, and this has been a factor in lowering the image quality of the image after the subtraction processing.

[0010] The image after the energy subtraction processing is an image obtained by subtracting a plurality of radiation images before the processing (hereinafter, the radiation image before the energy subtraction processing is referred to as an “original image”). Therefore, there is a problem that the S / N ratio is lower than the original image, and the image becomes hard to see.

The applicant assigns a parameter for each pixel in accordance with image data representing an average image (so-called blurred image) of a plurality of radiation images obtained based on a plurality of image data on which energy subtraction is performed. By performing the subtraction processing of each pixel while changing, it is possible to sufficiently reduce the influence of different beam hardening for each part of the subject and propose an energy subtraction image generation method capable of obtaining a high-quality subtraction image. (See JP-A-3-289277).

However, in order to perform the subtraction processing using the image data representing the blurred image as described above, it is necessary to use different energy subtraction coefficients for each location, and it is very difficult to set the coefficient optimally. Have difficulty.

Therefore, the present invention has been made in view of the above circumstances,
Even if the energy subtraction coefficient can be easily set, and even if the degree of beam hardening differs for each part of the subject, the effect can be sufficiently reduced, and therefore the energy that can obtain a high-quality subtraction image can be obtained. It is an object to provide a subtraction image generation method.

Another object of the present invention is to provide a method for generating a subtraction image having excellent observability and having reduced noise to substantially the same level as the original image before the subtraction processing, while achieving the above objects. To provide.

[0015]

An energy subtraction image generating method according to the present invention is obtained by recording radiation having different energy distributions transmitted through an object composed of a plurality of tissues having different radiation absorptances from each other. 2
For at least one of the two original image data representing each of the two radiation images, an unsharp mask signal Sus corresponding to a very low spatial frequency is obtained, the original image data is Sorg, and the enhancement coefficient is β When the reproduced image data is represented by S ', a frequency processing operation represented by an arithmetic expression S' = Sorg + β (Sorg-Sus) is performed, and the image data obtained by this frequency processing and the frequency processing are performed. By performing subtraction processing according to the original image data that has not been performed or according to two image data obtained by the frequency processing, a subtraction in which a shadow of a desired tissue in the subject is extracted or emphasized. An energy subtraction image generating method for obtaining image data, wherein the emphasis coefficient β If the original image data is an original image data representing an image due to radiation of higher energy than the radiation image, the original image data is negative in a low density range, otherwise it is 0, and the original image data is the two radiation images. If the original image data represents an image with radiation of lower energy than the original image data, the original image data is changed according to the density of the original image data so as to be positive in a low density range of the image data and to be 0 otherwise. It is a feature.

Further, the second energy subtraction image generating method of the present invention is obtained by recording radiation transmitted through a subject composed of a plurality of tissues having different radiation absorptances and having different energy distributions. For at least one of the two original image data representing each of the two radiation images, an unsharp mask signal Sus corresponding to a very low spatial frequency is obtained, the original image data is Sorg, and the enhancement coefficient is β And the reproduced image data is S '
Then, an operation of frequency processing represented by an arithmetic expression S ′ = Sorg + β (Sorg−Sus) is performed, and the image data obtained by the frequency processing and the original image data not subjected to the frequency processing are calculated. Based on, or based on two image data obtained by the frequency processing, first image data representing a first image of the subject, in which mainly a first tissue is recorded, is obtained, and the first image data is obtained. The first smoothed image data representing the first smoothed image in which the noise component of the first image has been reduced or removed by processing the image data of the first image is obtained, and the smoothed image data is obtained from the original image data. By subtraction processing, an energy subtraction image generation method for obtaining second image data representing a second image in which a second tissue of the subject is mainly recorded. Thus, when the original image data is original image data representing an image with radiation of higher energy than the two radiation images, the enhancement coefficient β is negative in a low density range of the original image data. Otherwise, it is 0, and when the original image data is original image data representing an image with radiation of lower energy than the two radiation images, the original image data is positive in the low density range of the image data, and otherwise 0. It is characterized in that it is changed according to the density of the original image data so that

In carrying out the second method, various modified methods are employed on the surface, such as decomposing the second method into smaller steps and changing the order of the operations. In some embodiments, the substantially same method as the second method can be realized, and the present invention is understood as a concept including various methods that are substantially the same.

Further, the third energy subtraction image generating method of the present invention is obtained by recording radiation transmitted through a subject composed of a plurality of tissues having different radiation absorptances and having different energy distributions. For at least one of the two original image data representing each of the two radiation images, an unsharp mask signal Sus corresponding to a very low spatial frequency is obtained, the original image data is Sorg, and the enhancement coefficient is β And the reproduced image data is S '
Then, an operation of frequency processing represented by an arithmetic expression S ′ = Sorg + β (Sorg−Sus) is performed, and the image data obtained by the frequency processing and the original image data not subjected to the frequency processing are calculated. Or a first process for obtaining first image data representing a first image of the subject, in which a first tissue is mainly recorded, based on two image data obtained by the frequency processing. After performing the first image data, the first image data is processed to obtain first smoothed image data representing a first smoothed image in which a noise component of the first image has been reduced, and the original image is obtained. By performing a subtraction process of the first smoothed image data from the data, a second process of obtaining second image data representing a second image in which the second tissue of the subject is mainly recorded is performed. After the second processing, the second image data is processed to obtain a second smoothed image data representing a second smoothed image in which the noise component of the second image has been reduced, A third process of obtaining new first image data representing a new first image in which the first tissue of the subject is mainly recorded by subtracting the second smoothed image data from the original image data; An energy subtraction image generation method for performing the processing of
The emphasis coefficient β is negative in the low density range of the original image data when the original image data is original image data representing an image due to radiation of higher energy than the two radiation images, and otherwise. Is 0, if the original image data is original image data representing an image with radiation of lower energy than the two radiation images, it is positive in the low density range of the image data, and 0 otherwise.
It is characterized in that it is changed according to the density of the original image data so that

Here, the second processing and the third processing in the third energy subtraction image generating method are repeatedly performed, whereby it is possible to obtain an image having better image quality performance. That is, the fourth energy subtraction image generation method of the present invention, after performing each process in the third energy subtraction image generation method, the new first image data obtained by the third process By performing the second processing again as the first image data in the second processing, a new second image representing a new second image in which mainly the second tissue of the subject is recorded By performing the third process again as a new second process for obtaining image data and the new second image data as the second image data in the third process, mainly the subject Repeating one or more times a new third process for obtaining new first image data representing a new first image in which the first tissue is recorded It is an feature.

Further, it is also possible to finally obtain the second image data of the subject by applying the third or fourth energy subtraction image generation method. That is, the fifth energy subtraction image generation method of the present invention, after performing the processing in the third or fourth energy subtraction image generation method,
The new first image data obtained by the third process or the new third process is used again as the first image data in the second process or the new second process. Performing a second process or the new second process to obtain new second image data representing a new second image in which the second tissue of the subject is mainly recorded. Is what you do.

Here, the third to fifth energy subtraction image generation methods include the same steps as those of the second energy subtraction image generation method. As described above, the above-described third to fifth energy subtraction image generation methods can be grasped as concepts including various aspects that are substantially the same. Further, the present invention is naturally included in the present invention as long as it includes each of the above methods including substantially the same method. For example, before carrying out the present invention, steps such as noise reduction processing by another method are included. Other steps may be included to further reduce noise after implementing the present invention.

The "first image" (including the "new first image") and the "second image" (including the "new second image") in each of the energy subtraction image generation methods. Are two images obtained by energy subtraction processing, in which the shadows of different tissues of the same subject are emphasized or extracted, and are not limited to specific ones. It refers to a bone image, an image in which a breast is emphasized, and an image in which a malignant tumor is emphasized when a human breast is taken as a subject.

[0023]

The beam hardening is determined by the total transmittance in the thickness direction of the subject between the front of the subject on which the radiation is incident and the back exiting from the subject. Therefore, the difference of each part of the beam hardening is Has a correlation with the density (the value of the image data) on the radiation image. Therefore, according to the present invention, an unsharp mask signal Sus corresponding to an ultra-low spatial frequency is obtained for at least one original image data, the original image data is Sorg, the enhancement coefficient is β, and the reproduced image data is S ′. At this time, an arithmetic operation of frequency processing represented by an arithmetic expression S '= Sorg + β (Sorg-Sus) is performed, and thereafter, a subtraction processing of each pixel is performed. What is characteristic in this frequency processing is that, when the original image data is original image data representing an image by radiation of higher energy than the two radiation images, the original image data It is negative in the low-density region of the data, and 0 otherwise. When the original image data is original image data representing an image by radiation of lower energy than the two radiation images, the low level of the image data is low. That is, the density is changed according to the density of the original image data so that the density is positive in the density range and 0 otherwise.

Here, by correcting the high spatial frequency in the low density range of the original image data as described above, the contrast of the bones can be enhanced and the effect of beam hardening as described above can be reduced. Can be obtained. Further, in the present invention, unlike a blurred image using a coefficient depending on each place, no artifact occurs.

The threshold values of the low-density region and other densities and the magnitude of the enhancement coefficient β are usually determined experimentally and empirically.

The image obtained by the ordinary energy subtraction processing has a reduced S / N ratio due to the execution of the subtraction processing. The second energy subtraction image generation method of the present invention has been studied with attention paid to this point.

That is, in this method, after the above-described frequency processing, subtraction processing is performed.
The first image of one of the two images (the first image and the second image) is obtained, and a first smoothed image in which a noise portion is reduced or removed is obtained from the image. Since the one smoothed image is subjected to the subtraction processing, a second image with reduced noise and excellent observation suitability is generated to the same degree as the original original image.

Here, in order to obtain a high-quality second image, it is necessary to delete only the noise component while maintaining the shadow of the first tissue of the subject when obtaining the first smoothed image. . However, the shading of the first tissue and the noise component are partially different from each other in their spatial frequency components,
For this reason, even if a non-linear filter that removes only the noise component as much as possible is used, there is naturally a limit to complete separation between the first tissue shadow and the noise component.

Therefore, the third to fifth energy subtraction image generation methods of the present invention abandon the complete separation of noise in a single noise reduction process, and repeatedly perform the noise reduction process. This is to generate an image with excellent observation suitability with reduced noise.

That is, in the third energy subtraction image generation method of the present invention, the noise component is reduced by processing the first image data, and then the second image data is obtained to obtain the second image data. In this method, new first image data is obtained by further reducing the noise component by performing the processing. As a result, an image with further reduced noise compared to the above-described second energy subtraction image generation method and having more excellent observation suitability is generated.

Further, the fourth energy subtraction image generation method of the present invention is intended to further reduce noise by repeating the third energy subtraction image generation method. Noise reduction processing can be shared, and an image with further reduced noise is generated.

Further, the fifth energy subtraction image generating method of the present invention is obtained by performing the third or fourth energy subtraction image generating method and then performing the third or fourth energy subtraction image generating method. Since the noise reduction processing is performed on the new first image data to perform the subtraction processing from the original image, a new second image with reduced noise components is generated.

[0033]

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS Hereinafter, the present invention will be described in detail based on embodiments shown in the drawings. Here, an example in which the above-described stimulable phosphor sheet is used will be described.

FIG. 1 is a schematic diagram of an X-ray imaging apparatus.

An object (the chest of a human body) 4 is irradiated with X-rays 3 emitted from an X-ray tube 2 of the X-ray imaging apparatus 1. The X-rays 3a transmitted through the subject 4 are irradiated on the first stimulable phosphor sheet 5, and X-rays having relatively low energy among the energies of the X-rays 3a are accumulated on the first stimulable phosphor sheet 5. Thus, an X-ray image of the subject 4 is stored and recorded on the sheet 5. The X-rays 3b that have passed through the sheet 5 pass through a filter 6 that further cuts low-energy X-rays, and the relatively high-energy X-rays 3c that have passed through the filter 6 irradiate the second stimulable phosphor sheet 7. Is done. Thus, the X-ray image of the subject 4 is also stored and recorded on the sheet 7. The subject 4 is provided with two marks 8 that serve as references for aligning two X-ray images when performing the subtraction processing. The “relatively low energy X-ray” means an X-ray having a lower energy than the X-rays incident on the two stimulable phosphor sheets 5 and 7, and the “relatively high energy X-ray” The “X-ray” means an X-ray having a higher energy than the X-rays incident on the two stimulable phosphor sheets 5 and 7.

It should be noted that the above-mentioned X-ray imaging apparatus requires only two
The X-ray images are stored and recorded on the two sheets 5 and 7, one at each of two temporally successive timings.
The photographing may be performed one by one.

FIG. 2 is a perspective view showing an example of an X-ray image reading apparatus and an image processing and display apparatus for implementing the energy subtraction image generating method of the present invention.

After the photographing is performed by the X-ray photographing apparatus 1 shown in FIG. 1, the first and second stimulable phosphor sheets 5, 7 are set one by one at a predetermined position of the X-ray image reading apparatus 10. You. Here, the case of reading the first X-ray 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 X-ray image is stored is recorded by an endless belt or other sheet conveying means 15 driven by a driving means (not shown) in the direction of arrow Y. (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 rotated at a high speed in the direction of arrow Z, passes through a focusing lens 20 such as an fθ lens, And enters the stimulable phosphor sheet 14 in the sub-scanning direction (arrow Y direction).
The main scanning is performed in the direction of the arrow X substantially perpendicular to. From the portion of the stimulable phosphor sheet 14 where the light beam 17 is irradiated, stimulated emission light 22 is emitted in an amount corresponding to the accumulated and recorded X-ray image information. It is guided by a guide 23 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 so that a linear incident end face 23a extends along a main scanning line on the stimulable phosphor sheet 14. 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 by the photomultiplier 24 into an electric signal.

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.
A digital image signal S0 is obtained. This image signal S0
Represents a first X-ray image stored and recorded on the first stimulable phosphor sheet 5, and is referred to as a _first image signal S01. The first image signal S0 ₁ image processing and displaying apparatus 30
Is temporarily stored in the internal memory.

The image processing and display device 30 includes a keyboard 31 for inputting various instructions, a CRT display 3 for displaying auxiliary information for instructions and a visible image based on image signals.
2. A floppy disk drive unit 33 for loading and driving a floppy disk as an auxiliary storage medium, and a main unit 34 having a CPU and an internal memory are provided.

[0042] Next, in the same manner as described above, the second image signal S0 ₂ representing the second X-ray image stored recorded in the second stimulable phosphor sheet 7 is obtained, the second image signal S0
_{2 is} also temporarily stored in the internal memory in the image processing display device 30.

FIG. 3 shows an image processing and display device based on _two image signals SO ₁ and SO ₂ representing first and second X-ray images stored in an internal memory in the image processing and display device. It is a figure showing the flow of the process performed.

The first and second X-ray image signals SO ₁ and SO stored in the internal memory of the image processing display device.
₂ are the first X-ray image 41 and the second X-ray image 41 shown in FIG.
This is a signal carrying the line image 42. The first X-ray image 41 is an image based on relatively low energy X-rays, and the second X-ray image 42 is an image based on relatively high energy X-rays. Both are original images in which both the soft part and the bone part are recorded.

The first and second X-ray image signals SO
₁ and SO ₂ are read from an internal memory in the image processing display device 30 shown in FIG. 2, and at least one image signal is subjected to the following frequency processing. That is, an unsharp mask signal Sus corresponding to an ultra-low spatial frequency is obtained, and when the original image data is Sorg, the enhancement coefficient is β, and the reproduced image data is S ′, S ′ = Sorg + β (Sorg−Sus) by using an arithmetic expression expressed frequency processing, (1) above in the low density region of the emphasis coefficient β the X-ray image signal SO ₁ positive, 0 in other cases, with respect to the X-ray image signal SO ₁ (2) The emphasis coefficient β is negative in the low-density region of the X-ray image signal SO ₂ , and 0 otherwise.
As, it operates on the X-ray image signal SO _2;
(3) above in the low density region of the emphasis coefficient β the X-ray image signal SO ₁ positive, 0 in other cases, performs an operation on the X-ray image signal SO _1, further wherein said emphasis coefficient β Any one of the following frequency processes is performed: the X-ray image signal SO ₂ is set to be negative in a low density range of the X-ray image signal SO ₂ and set to 0 otherwise. FIG.
Is a graph showing an enhancement coefficient in an arithmetic expression for performing the above-described frequency processing.
(B) shows an emphasis coefficient in an arithmetic expression for performing frequency processing of an image with relatively low energy. Such frequency processing is performed by subtraction processing using an X-ray image in which a low-density region of an X-ray image having a relatively low energy has a high density due to a bias toward the high energy side due to the above-described beam hardening. Can be corrected. Thus, the frequency-processed X
The line image signals SO ₁ ′ and SO ₂ ′ are obtained (if only one of them is subjected to the frequency processing, the other is used without being subjected to the frequency processing, but here the X-ray image signal SO ₁ ′ and SO ₂ ′).

Next, these two image signals SO ₁ ',
The relative positioning of each of the X-ray images 41 and 42 carried by SO ₂ 'is performed on the image signal (Japanese Patent Laid-Open No. 58-163).
No. 338). This alignment is performed by relatively linearly moving and rotating the two X-ray images so that the two marks 8 shown in FIG. 1 overlap. It should be noted that the relative positional relationship in each image is not substantially changed by the above-described gradation processing, and only the density is changed. Therefore, this alignment may be performed before the frequency processing.

As described above, after the frequency processing is performed to obtain the image signals SO ₁ ′ and SO ₂ ′ and the alignment is performed, the equation is calculated for each pixel.

[0048]

(Equation 1)

A subtraction process is performed in accordance with the above, and a soft part image 43 (see FIG. 3) is obtained in which the shadow of the bone of the subject 4 is eliminated and only the soft part is extracted. Here, Ka and Kb are two image signals SO ₁ ',
A parameter for determining the weight of SO ₂ ′, Kc is a parameter for determining the bias, and in the present embodiment, each of the parameters Ka, Kb, and Kc is a constant.

These constants are stored in advance in a memory (parameter storage means in the present invention) in the image processing display device 30, and are referred to when performing the subtraction processing according to the above equation (1).

The soft part image signal S1 obtained in accordance with the above equation (1) is used for the CRT display 32 of the image processing and display device 30.
And a visible image based on the soft part image signal S1 is represented by C
It is reproduced and displayed on the RT display 32.

The above embodiment is an example in which the soft part image signal S1 is obtained. However, when the bone part image 44 (see FIG. 3) is to be observed, the bone part image signal S2 is obtained using the bone part image parameters. Should be obtained.

Next, a second energy subtraction image generation method of the present invention will be described with reference to embodiments. This embodiment is similar to the above-described embodiment of the first energy subtraction image generation method of the present invention,
The description will be made using the X-ray imaging apparatus shown in FIG. 1 and the X-ray image reading apparatus and the image processing display apparatus shown in FIG.

[0054] FIG. 5 is based on the first and second X-ray image two image signals S0 ₁₁ representing a, S0 ₁₂ stored in the internal memory of the image processing display device, in the image processing and displaying apparatus FIG. 9 is a diagram illustrating an example of a flow of processing performed.

[0055] stored in the internal memory of the image processing display device, first and second X-ray image signal S0 _11, S0
₁₂ is a first X-ray image 51 and a second X-ray image 51 shown in FIG.
This is a signal carrying the line image 52. The first X-ray image 51 is an image based on relatively low energy X-rays, and the second X-ray image 52 is an image based on relatively high energy X-rays. Both are original images in which both the soft part and the bone part are recorded.

The first and second X-ray image signals SO
₁₁ and SO ₁₂ are read from an internal memory in the image processing and display device 30 shown in FIG. 2, and at least one of the image signals is subjected to the same frequency processing as in the above-described embodiment. As described above, the frequency-processed X-ray image signals SO ₁₁ ', SO _11
12 'when the frequency process is performed only on the obtained (either, but the other is frequency processing is used as it is without subjected, for convenience the X-ray image signal SO ₁₁ here',
Both are represented by SO ₁₂ ').

Next, these two image signals SO ₁₁ ',
The relative positioning of each of the X-ray images 51 and 52 carried by SO ₁₂ ′ is performed on the image signal (Japanese Patent Laid-Open No. 58-163).
No. 338). This alignment is performed by relatively linearly moving and rotating the two X-ray images so that the two marks 8 shown in FIG. 1 overlap. It should be noted that the relative positional relationship in each image is not substantially changed by the above-described gradation processing, and only the density is changed. Therefore, this alignment may be performed before the gradation processing.

Thereafter, a subtraction process is performed.

Here, the absorption coefficient μ of X-rays is determined as follows separately for the soft part and the bone part of the subject and for the low-energy X-rays and the high-energy X-rays.

Μ _L ^T : Absorption coefficient of soft part by low energy X-ray μ _H ^T : Absorption coefficient of soft part by high energy X-ray μ _L ^B : Absorption coefficient of bone by low energy X-ray μ _H ^B : High energy X Absorption coefficient of bone by line At this time, for each pixel of two image signals SO ₁₁ ′ and SO ₁₂ ′ corresponding to each other, the equation

[0061]

(Equation 2)

However, C is weighted and subtracted according to the bias component to obtain a bone image signal S11 representing the bone image 53 (see FIG. 5) from which the shadow of the bone is extracted.

Also, the equation

[0064]

(Equation 3)

However, the soft image signal S12 representing the soft image can be obtained by weighting and subtracting C 'according to the bias component, but this operation is unnecessary in the present embodiment.

Further, the expression

[0067]

(Equation 4)

The superimposition image 54 of the two X-ray images 51 and 52 is generated by performing the addition process for each pixel corresponding to each other according to the following. This superimposed image 54 is also an original image in which both the soft part and the bone part are recorded. Although the X-ray image 51 or the X-ray image 52 can be used instead of the superimposed image 54, the superimposed image 54
Since the line images 51 and 52 are superimposed, the noise component is reduced compared to any of these X-ray images.
Therefore, it is advantageous for the subsequent processing.

Next, the noise component contained in the bone image 53 is extracted by processing the bone image signal S11.

FIG. 6 is a diagram showing a spectrum with respect to a spatial frequency f of an image obtained by processing a bone image and a bone image signal.

The graph 61 shown in the figure represents the spectrum of the bone image 53, and includes a noise component 63.

Here, first, the bone image signal S11 is subjected to a smoothing process. As this smoothing processing method, for example, for each pixel, an average of image signals corresponding to each pixel in a predetermined area centered on the pixel is obtained, and this average value is used as a simple average as an image signal of the central pixel. A method using a median filter that sets the median value (median) of the image signal in the predetermined area as the image signal of the pixel at the center, and further divides the predetermined area into a plurality of small areas, And a method using an edge-preserving filter (V-filter) that uses the average value of the small area having the smallest variance as the value of the image signal of the central pixel, and performs a Fourier transform on the image signal to obtain the high Although a method of performing an inverse Fourier transform after removing a spatial frequency component can be used, the above-described blur mask processing method has a drawback that an edge is blurred. In this method, pixels are exchanged, so that contour-like artifacts may occur.In addition, when the edge-preserving filter is used, honeycomb-like artifacts may occur. There's a problem. Therefore, in this embodiment, smoothing is performed using the following histogram adaptive filter which is different from any of the above methods. By using this method, it is possible to remove noise without preserving the necessary edges (step-like density changes that define the boundaries of two different tissue shadows) as image information and eliminate the artifacts. And has the advantage that noise can be removed in a short time.

First, for each pixel of the bone image, a histogram of the image signal S11 of a large number of pixels in a predetermined area centered on the pixel is created.

FIGS. 7A and 7B show the appearance frequencies of the image signal S11 corresponding to a number of pixels in a predetermined area centered on a certain pixel (image signal S11 ') obtained as described above. FIG. 8 is a diagram illustrating an example of a function using a difference between the image signal S11 and the image signal S11 ′ of the central pixel as a variable.

A function representing a histogram as shown in FIGS. 7A and 7B is generally represented by h (S11), and monotonically decreases as the absolute value | S11-S11 '| increases. For example, as shown in FIG. A function such as f (S11-S1
1 '). At this time, the expression

[0076]

(Equation 5)

A function g (S
11) is obtained. This function g (S11) is equivalent to the function h (S
If 11) has a plurality of peaks as shown in FIG. 7A, it has the effect of extracting only the peak to which the image signal S11 'of the central pixel belongs.

After obtaining the function g (S11) according to the above equation (5), the average value <S11> of the image signal S11 weighted by the function g (S11) is obtained. That is,
Specifically, for example, the first moment of the function g (S11) is obtained according to the following equation.

[0079]

(Equation 6)

The processing according to the above equations (5) and (6) is performed with each pixel of the bone image as the center pixel, whereby the smoothed image signal <S11> (for simplicity, the image corresponding to each pixel is processed). The same symbol is used for the signal and the image signal representing the entire image.) The smoothed image signal <S11> is a signal mainly obtained by removing the high spatial frequency components of the original bone part image signal S11 as shown in a graph 62 of FIG. As shown in a), since the signal is obtained by calculating the average value after extracting only the mountain to which the pixel belongs, the edge in the original bone image is stored without blurring.

Next, the smoothed image signal S11 is weighted and subtracted from the superimposed image signal S0 (see the above equation (4)) representing the superimposed image 44 for each pixel, that is,

[0082]

(Equation 7)

Here, C ″ represents a bias component.

As a result, the image information carries substantially the same information as the soft part image represented by the above equation (3), and the noise component is reduced compared to the soft part image represented by the above equation (3). Soft part image 56 (see Fig. 5)
Is required.

The image signal S1 obtained according to the equation (7)
2 'is sent to the CRT display 32 of the image processing and display device 30, and a visible image based on the image signal S12' is displayed on the CRT.
It is reproduced and displayed on the display 32.

In the above embodiment, the bone image signal S11 is smoothed and subtracted from the original image to obtain the soft image signal S1.
In the example of obtaining 2 ′, when a bone image is to be observed, a soft image signal S12 is obtained based on the above equation (3), and the soft image signal S12 is smoothed and subtracted from the original image. What is necessary is just to obtain the bone image in which the noise component is reduced.

Next, processing that is substantially the same as that of the above-described embodiment described with reference to FIG. 9 and is therefore included in the present invention will be described.

FIG. 9 shows two image signals S0 representing the first and second X-ray images stored in the internal memory of the image processing and display device for explaining this substantially identical embodiment.
_11, S0 based on ₁₂ is a diagram showing another example of a flow of processing performed by the image processing and displaying apparatus. The same elements as those in FIG. 5 are denoted by the same reference numerals and symbols as those in FIG. 5, and descriptions of the parts described with reference to FIG. 5 are omitted here to avoid redundant description.

From the two X-ray images 51 and 52, the above equation (2)
Bone image 53 (bone image signal S11) based on equation (3)
And the soft part image 57 (soft part image signal S12) are obtained.

Next, the noise component contained in the bone image 53 is reduced by processing the bone image signal S11 based on the above equations (5) and (6) in the same manner as in the above-described embodiment. A smoothed image signal <S11> is obtained, and then the smoothed image signal <S11 is calculated from the bone image signal S11 for each pixel.
11>, a noise image 48 (noise signal S _N ) from which only the noise component is extracted is obtained.

[0091]

(Equation 8)

The noise signal _SN is a signal obtained by extracting a noise component of the bone image as shown in a graph 63 of FIG.
Here, in the smoothed image signal <S11>, since the information on the edge of the bone image is stored even if the spatial frequency is as high as the noise component, the bone image signal S11 and the bone image signal S11 are calculated according to the above equation (8). By obtaining the difference from the smoothed image signal <S11>, the edge information is neatly canceled,
Therefore, the noise signal _SN becomes a signal carrying only the noise component of the bone image more purely as compared with the case where the smoothing process for losing edge information is performed.

Next, the noise signal S _N thus obtained and the soft part image signal S representing the soft part image 57 (see FIG. 9) are obtained.
12 is weighted and added for each pixel, whereby image information carries substantially the same information as the soft part image 57 and a noise component is lower than that of the soft part image 57. In the present embodiment, this weighted addition is represented by an equation

[0094]

(Equation 9)

The noise component is further reduced.

Here, it will be described that the embodiment described with reference to FIG. 5 and the embodiment described with reference to FIG. 9 are substantially the same.

The soft image signal S12 shown by the above equation (3) and the noise signal _SN shown by the above equation (8) are substituted into the above equation (9). It should be noted that the bias component (C 'in the above equation (3)) adjusts the finally obtained density of the entire image (including the luminance when displayed on a CRT display device or the like). Will be omitted.

By substituting equations (3) and (7) into equation (9),

[0099]

(Equation 10)

By substituting the bone image signal S11 represented by the above equation (2) into the equation (10), the bias component C
Is ignored),

[0101]

[Equation 11]

By rearranging this equation (11),

[0103]

(Equation 12)

Then, if the above equation (3) is substituted,

[0105]

(Equation 13)

The following is obtained. This equation (13) is the same as the above equation (6) except for the bias component. That is, substantially the same processing is performed in the embodiment described with reference to FIG. 5 and the embodiment described with reference to FIG.

FIG. 10 is a diagram showing a flow of processing according to another embodiment of the present invention. FIG. 11 is a diagram schematically showing a profile of each image shown in FIG. 10 in one predetermined direction. .

In FIG. 10, the elements corresponding to those in FIG. 5 or FIG. 9 are denoted by the same reference numerals and symbols as in FIG. 5 and FIG.

FIGS. 11 (a) and 11 (b) are diagrams schematically showing X-ray images (original images) 51 and 52, respectively, in a predetermined direction (x-direction) on the X-ray images 51 and 52. a plot of the values of the image signal S0 ₁₁ ', S0 _12' along, these image signals S0 ₁₁ ', S0 _12' facilities shaded in uniform soft (FIG although different from the value one another ) And a signal component representing a stepped bone part are superimposed, and a random noise component is superimposed.

[0110] Two X-ray image (original image) 51 and 52 These two image signals S0 ₁₁ representing the ', S0 _12' the basis of the equation (3) weighting subtraction process on the basis of - a (symbols in) As a result, a soft part image signal S12 representing the soft part image 57 is obtained, and two image signals S0 ₁₁ ′ and S
A superimposed image signal S0 representing the superimposed image 54 is obtained by performing an addition process (represented by a symbol +) based on the expression (4) based on 0 ₁₂ '.

FIG. 11 (c) is a diagram schematically showing the superimposed image signal S0. As in FIGS. 11 (a) and 11 (b), a uniform signal component representing the soft part (the hatched portion in FIG. 11A), a signal component representing a bone part changed stepwise, and a random noise component are superimposed. The noise component is composed of two X-rays shown in FIGS. 11A and 11B. The images are reduced as compared with the images 51 and 52.

FIG. 11D is a diagram showing the soft part image signal S12 obtained based on the above equation (3). Although only signal components representing uniform soft parts are extracted, random noise components are extracted from the two X-ray images 51 and 52.
(FIGS. 11A and 11B).

Although it is not necessary to obtain the bone image signal S11 in this embodiment, it is assumed that the bone image signal S11 is obtained based on the above equation (2). Although a signal component representing a bone changed in a step shape is extracted, a random noise component is generated in the two X-ray images 5 as in the soft image signal S12 (FIG. 11D).
1, 52 (FIGS. 11A and 11B).

Here, the soft part image 57 (soft part image signal S12,
11D) is subjected to a smoothing process 71 (see FIG. 10), and a smoothed soft part image signal <S representing a smoothed soft part image 81 is obtained.
12> (FIG. 11F). This smoothing process
At 81, a high spatial frequency component of, for example, 1.0 cycle / nm or more of the soft part image 57 is cut.

Next, the smoothed soft part image signal <S12> is weighted and subtracted from the superimposed image signal S0, whereby the bone part image signal S11 'representing the bone part image 82 is obtained. As shown in FIG. 11 (g), the bone part image signal S11 ′ has a reduced random noise component compared to the bone part image signal S11 (FIG. 11 (e)).
The effect of performing the smoothing process on 57 appears, and the high spatial frequency component of the soft part image is slightly mixed.

Next, a smoothing process 82 is performed on the bone image signal S11 'obtained as described above. In the smoothing process 82 performed here, only low-contrast shadows (small changes in the bone image signal S11 ') in the spatial frequency band of, for example, 0.5 cycles / mm or more of the bone image 62 are cut. . As this processing method, for example, a window having an area corresponding to 0.5 cycle / mm with respect to a predetermined pixel P _O is considered, and a predetermined pixel P _O among the signals S11 ′ corresponding to each pixel in this window is considered. new signal S of the pixels P _O with an average value of a predetermined the average value of the 'signal S11 is within the value ± predetermined value' corresponding signal S11 to
A method of scanning on the bone image 82 using a filter of 1 _O ′ is adopted. By this smoothing processing 72, a smoothed bone image signal <S11 ′> representing the smoothed bone image 83 is obtained. This smoothed bone image signal <S11 ′> is
As shown in FIG. 11 (i), although the noise component and the high frequency component of the mixed soft part image are reduced, the rising part is also dull.

Next, the smoothed bone image signal <S11 '> is weighted and subtracted from the superimposed image signal S0 to obtain a soft image signal S12' representing the soft image 84. As shown in FIG. 11 (h), the soft part image 84
Although the noise component is reduced as compared with (FIG. 11D), the smoothed bone image signal <S11 ′> (FIG. 11I)
Since the rising part of the part is dull, the information of the bone part image of that part is superimposed as noise. However, the information of the random noise portion and the bone image as noise is considerably small. Therefore, a series of processing is stopped at this stage, and the soft image signal S12 'is converted to the CR of the image processing display device 30.
The visible image based on the soft part image signal S12 'may be sent to the T display 32 (see FIG. 2) to be reproduced and displayed on the CRT display for observation.

However, in the present embodiment, the same processing as described above is further repeated to further improve the image quality.

A soft part image signal S12 'representing the soft part image 84
Is obtained, a smoothing process 73 is performed on the soft part image signal S12 'to obtain a smoothed soft part image signal <S12'> (FIG. 11 (j)) representing the smoothed soft part image 85. As the smoothing process 73, for example, a process of cutting a spatial frequency component of 1.5 cycles / mm or more is performed.

The smoothed soft part image signal <S12 ′> is subjected to weighted subtraction processing from the superimposed image signal S0 to obtain a bone part image signal S11 ″ representing the bone part image 86. As shown in FIG.
82 (FIG. 11 (g)), random noise and information of the soft part image mixed as noise are also reduced. When a bone image is to be observed, a visible image based on the bone image signal S11 ″ may be reproduced and displayed on the CRT display 32.

In this embodiment, the bone image signal S11 ″ obtained as described above is further subjected to a smoothing process 74, and the smoothed bone image signal <S1 representing the smoothed bone image 87 is obtained.
1 ″> (FIG. 11 (m)).
As 74, for example, a low contrast component of 1.0 cycle / mm or more is cut.

Next, the smoothed bone image signal <S11 ″> is weighted and subtracted from the superimposed image signal S0 to obtain a soft image signal S12 ″. This soft part image signal S
As shown in FIG. 11 (l), 12 ″ has further reduced both the random noise and the information of the bone image as noise as compared with the previously obtained soft part image signal S12 ′ (FIG. 11 (h)). Signal.

By repeating the smoothing process and the weighted subtraction of the superimposed image (original image) in this manner, it is possible to alternately obtain a bone image and a soft image in which noise is sequentially reduced.

FIG. 12 is a diagram showing the flow of another process substantially the same as the embodiment described with reference to FIG. The same elements as those in FIG. 10 and the like are denoted by the same reference numerals and symbols as in FIG. 10 and the like, and description thereof is omitted.

The processing shown in FIG. 12 is the processing up to obtaining the bone image 82 shown in FIG. 10 (the processing described with reference to FIG. 5 (however, the bone image and the soft image are different from those in FIG. 5). )) Is replaced by the processing described with reference to FIG. 6 (however, the bone image and the soft image are replaced in FIG. 6), and as described above, these are substantially the same as each other. Processing.

In the processing shown in FIG. 12, only the first stage is replaced by the processing method described with reference to FIG. 6, but this replacement can be performed at any stage of the repeated processing. The processing is substantially the same, and the present invention includes all substantially the same processing modes changed in any one or more of these steps.

In each of the above embodiments, the X of the chest of the human body is used.
Although an example of obtaining a soft part image or a bone part image based on a line image is described, the present invention is not limited to obtaining a soft part image or a bone part image. For example, an image in which a mammary gland is emphasized or a malignant tumor is obtained. May be broadly applied to obtain one or both of two images in which two different tissues in a subject are respectively emphasized or extracted.

Further, although the above embodiment is an example using a stimulable phosphor sheet, the present invention is not limited to a stimulable phosphor sheet, but is applied to an X-ray film (generally sensitized for photographing). (Combined with a screen).

[0129]

As described in detail above, the energy subtraction image generating method of the present invention obtains an unsharp mask signal Sus corresponding to an ultra-low spatial frequency for at least one original image data, and Sor image data
g, the emphasis coefficient is β, and the reproduced image data is S ′. The emphasis coefficient β is calculated using the arithmetic expression of frequency processing expressed by S ′ = Sorg + β (Sorg−Sus). Is negative in the low-density region of the original image data when the image is a relatively high-energy image, and 0 otherwise. When the original image data is an image with a relatively low energy, the image data This frequency processing is performed in accordance with the density of the original image data so as to be positive in the low density range of 0, and 0 otherwise, and then the subtraction processing of each pixel is performed. The contrast of the image due to high spatial frequency can be enhanced without changing the contrast of the image due to the frequency, and the subject can be captured without the artifacts that occur when using blurred images. Sufficiently reduces the negative effects of different beam hardening for each part of, it is possible to obtain a subtraction image of high quality.

Further, another energy subtraction image generating method of the present invention represents a first image in which mainly a first tissue in a subject is recorded after performing frequency processing by the above-described energy subtraction image generating method. The first image data is obtained, the noise component of the first image data is reduced or removed to obtain the first smoothed image data, and the first smoothed image data is subtracted from the original image data. Since the second image data is obtained, an image with reduced noise components and excellent observation suitability is generated.

Further, the first image and the second image are alternately smoothed and the subtraction process from the original image is repeated to generate a first image and a second image in which noise components are further reduced. be able to.

The above-described present invention has a remarkable effect particularly on the one-shot energy-sub, which is greatly affected by the beam hardening.

[Brief description of the drawings]

FIG. 1 is a schematic diagram of an X-ray imaging apparatus.

FIG. 2 is a perspective view showing an example of an X-ray image reading apparatus and an image processing and display apparatus for implementing the energy subtraction image generation method of the present invention.

FIG. 3 is a diagram showing a flow of processing performed in the image processing display device.

FIG. 4 is a diagram illustrating frequency processing in the energy subtraction image generation method according to the present invention.

FIG. 5 is a diagram showing a flow of another process performed in the image processing display device.

FIG. 6 is a diagram illustrating a spectrum with respect to a spatial frequency f of a bone image and an image obtained by processing the bone image signal.

FIG. 7 is a diagram showing two different histograms in which the appearance frequencies of image signals S11 corresponding to a large number of pixels in a predetermined area centering on a certain pixel (image signal S11 ′) are plotted.

FIG. 8 shows an image signal S11 and an image signal S1 of a central pixel.
Diagram showing an example of a function using the difference from 1 'as a variable

FIG. 9 shows two image signals S0 representing first and second X-ray images stored in an internal memory in the image processing display device.
_11, S0 based on _12, illustrates another example of a flow of processing performed by the image processing and displaying apparatus

FIG. 10 is a diagram showing a processing flow of another embodiment of the present invention.

11 is a diagram schematically showing a profile of each image shown in FIG. 10 in one predetermined direction.

FIG. 12 is a diagram showing a flow of another process substantially the same as the process shown in FIG. 10;

[Description of Signs] 1 X-ray imaging apparatus 2 X-ray tube 3 X-ray 4 subject 5 first stimulable phosphor sheet 6 filter 7 second stimulable phosphor sheet 8 mark 16 laser light source 19 rotating polygon mirror 22 luminosity Excitation light 23 Light guide 24 Photomultiplier 25 Log amplifier 26 A / D converter 30 Image display device 41,42,71,72 X-ray image (original image) 43,82,86 Bone image 44 Superimposed image ( Original image) 45,83,87 Smooth bone image 46,47,84 Soft image 48 Noise image 71,72,73,74 Smoothing 81,85 Smooth soft image

Claims (7)

(57) [Claims] 1. Two original image data representing each of two radiation images transmitted through a subject composed of a plurality of tissues having different radiation absorptances and obtained by recording radiation having different energy distributions. For at least one of the original image data, an unsharp mask signal Sus corresponding to an ultra-low spatial frequency is obtained, and the original image data is
g, the emphasis coefficient is β, and the reproduced image data is S ′, a frequency processing operation represented by an arithmetic expression S ′ = Sorg + β (Sorg−Sus) is performed, and the image data obtained by this frequency processing is By performing a subtraction process according to the original image data that has not been subjected to the frequency processing or according to two image data obtained by the frequency processing, a desired tissue shadow in the subject is reduced. An energy subtraction image generation method for obtaining extracted or emphasized subtraction image data, wherein the enhancement coefficient β is determined by using original image data in which the original image data represents an image of radiation of higher energy than the two radiation images. In some cases, the original image data is negative in the low-density range, otherwise it is 0, The two in the case of original image data representing an image by radiation of lower energy than of radiation image positive in the low density region of the image data, in other cases 0
A method according to the density of the original image data so that 2. Original image data representing two radiation images transmitted through a subject composed of a plurality of tissues having different radiation absorptances and obtained by recording radiation having different energy distributions. For at least one of the original image data, an unsharp mask signal Sus corresponding to an ultra-low spatial frequency is obtained, and the original image data
org, the enhancement coefficient is β, and the reproduced image data is S ′, a frequency processing operation represented by an arithmetic expression S ′ = Sorg + β (Sorg−Sus) is performed, and the image data obtained by this frequency processing is A first image in which mainly a first tissue in the subject is recorded, based on the original image data not subjected to the frequency processing or based on two image data obtained by the frequency processing. First smoothed image data representing a first smoothed image in which noise components of the first image have been reduced or removed by processing the first image data. And subtracting the smoothed image data from the original image data to obtain second image data representing a second image of the subject, in which mainly a second tissue is recorded. A method for generating a subtraction image, comprising: when the original image data is original image data representing an image with higher-energy radiation among the two radiation images, Negative in the density region, 0 otherwise, and when the original image data is original image data representing an image with radiation of lower energy than the two radiation images, in the low density region of the image data Positive, otherwise 0
A method according to the density of the original image data so that 3. Original image data of two radiation images transmitted through an object composed of a plurality of tissues having different radiation absorptances and respectively representing two radiation images obtained by recording radiation having mutually different energy distributions. For at least one of the original image data, an unsharp mask signal Sus corresponding to an ultra-low spatial frequency is obtained, and the original image data
org, the enhancement coefficient is β, and the reproduced image data is S ′, a frequency processing operation represented by an arithmetic expression S ′ = Sorg + β (Sorg−Sus) is performed, and the image data obtained by this frequency processing is A first image in which mainly a first tissue in the subject is recorded, based on the original image data not subjected to the frequency processing or based on two image data obtained by the frequency processing. After performing a first process for obtaining first image data representing a first image data, the first image data is processed to represent a first smoothed image in which a noise component of the first image is reduced. By obtaining one smoothed image data and subtracting the first smoothed image data from the original image data, a second image mainly representing a second image of the subject in which a second tissue is recorded is obtained. of A second process for obtaining image data is performed, and after the second process, the second image data is processed to represent a second smoothed image in which a noise component of the second image is reduced. By obtaining the second smoothed image data and subtracting the second smoothed image data from the original image data, a new first image in which mainly the first tissue of the subject is recorded is obtained. An energy subtraction image generation method for performing a third process of obtaining new first image data to be represented, wherein the enhancement coefficient β is set to an image obtained by using higher-energy radiation than the original image data of the two radiation images. Is negative in the low-density region of the original image data, otherwise it is 0, and the original image data is lower than the two radiographic images. Positive in the low density region of the image data in the case of original image data representing an image by radiation of ghee, 0 otherwise
A method according to the density of the original image data so that 4. After performing the processing according to claim 3, the new first image data obtained by the third processing is used again as the first image data in the second processing. By performing the second process, a new second process for obtaining new second image data representing a new second image mainly recording the second tissue of the subject, and the new second process By performing the third process again as the second image data in the third process using the second image data, a second image mainly representing the first tissue of the subject is recorded. A method for generating an energy subtraction image, comprising repeating one or more times a new third process for obtaining new first image data. 5. After performing the processing according to claim 3 or 4, the new first image data obtained by the third processing or the new third processing is subjected to the second processing or By performing the second processing or the new second processing again as the first image data in the new second processing, a new second processing in which mainly the second tissue of the subject is recorded An energy subtraction image generation method, wherein new second image data representing a second image is obtained. 6. The radiation according to claim 1, wherein the radiation is two radiations obtained by transmitting a single radiation through the subject and then separating the energy distribution into components different from each other. 4. The method for generating an energy subtraction image according to claim 2 or 3. 7. The energy according to claim 1, wherein the two radiation images are obtained by separately recording the radiation with recording materials having different recording energy distributions. Subtraction image generation method.