JPH0743767B2 - Radiation image processing method and apparatus - Google Patents

Description

Description: TECHNICAL FIELD The present invention relates to frequency processing of a radiation image signal, and particularly to frequency processing that suppresses deterioration of graininess of a reproduced image due to reduction of a radiation dose applied to a subject. The present invention relates to a radiation image processing method for performing the method and an apparatus for performing the method.

(Prior Art) When a certain kind of phosphor is irradiated with radiation (X-ray, α-ray, β-ray, γ-ray, electron beam, ultraviolet ray, etc.), a part of this radiation energy is accumulated in the phosphor, It is known that when a phosphor is irradiated with excitation light such as visible light, the phosphor exhibits stimulated emission depending on the stored energy, and a phosphor having such a property is a stimulable phosphor (luminescent material). Exhaustible phosphor).

Using this stimulable phosphor, a subject such as a human body is irradiated with radiation to be photographed, the radiation image information of this subject is once recorded on a sheet of the stimulable phosphor, and the stimulable phosphor sheet is used for laser light irradiation. 2D scans with excitation light such as to generate stimulated emission light, and the obtained stimulated emission light is photoelectrically read by a photodetector to obtain an image signal. Based on this image signal, a photographic light-sensitive material The applicant has already proposed a radiation image information recording / reproducing system that outputs a radiation image of a subject as a visible image on a recording material such as CRT or a display device such as a CRT. (JP-A-55-12429, JP-A-56-11395, etc.) This system is practical in that it can record an image over an extremely wide radiation exposure area as compared with a radiographic system using a conventional silver salt photograph. Have advantages. That is, in the stimulable phosphor, it has been recognized that the amount of emitted light stimulated by excitation after storage is proportional to the radiation exposure amount over a very wide range,
Therefore, even if the radiation exposure amount fluctuates considerably due to various photographing conditions, the amount of stimulated emission light emitted from the stimulable phosphor sheet is read by the photoelectric conversion means by setting the reading gain to an appropriate value. A radiation image that is not affected by fluctuations in radiation exposure is obtained by converting it into an electrical signal and using this electrical signal to output a radiation image as a visible image on a recording material such as a photographic photosensitive material or a display device such as a CRT. be able to.

(Problems to be Solved by the Invention) When the radiation image information recording / reproducing system is used for diagnosing a human body, the exposure dose to the human body can be significantly reduced as compared with a conventional X-ray imaging diagnosis system.

However, as the amount of radiation applied to the subject at the time of shooting is reduced, the effect of quantum noise of radiation on the radiographic image increases, and the graininess of the image deteriorates, resulting in a rough and grainy reproduced image.

Among the methods for improving this graininess, the device's idea is to thicken the stimulable phosphor sheet or enlarge the particles of the stimulable phosphor used in this sheet to store and record a blurred image at the time of shooting. It is conceivable to increase the diameter of the excitation light to be scanned and read the image while blurring the image, input the read analog image signal to the analog filter, and blur the image. Despite the need for delicate control in order to improve graininess and suppress deterioration of other image quality performance such as sharpness as much as possible, in the above method, the types of sheets must be increased respectively. The degree of freedom that can be controlled is limited even if the number is increased, the degree of freedom that can be controlled is extremely low in spite of the complicated mechanism, and only the time-series image signal flow direction (main scanning direction) can be controlled. I have a problem. Also, as a method of improving this graininess by image processing, FFT
(Fast Fourier Transform) is used for frequency processing, and a method of digitally blurring each scanning point by obtaining an average value of image signals around the scanning point can be considered. The method using the FFT has a very large degree of controllable degree of freedom, but has a problem that the processing speed is too slow to be applied to a large-capacity image signal, and a high cost is required to increase the processing speed. The above-mentioned digitally blurring method has a short processing time, but delicate control cannot be performed.
There is a problem that it is usually too blurry.

In view of the above problems, the present invention is capable of improving the graininess of a radiation image and suppressing deterioration of other image quality performance to a minimum, and further, without complicating the apparatus, a sufficient calculation time is acceptable. It is an object of the present invention to provide a radiation image processing method therein and an apparatus capable of implementing the method.

(Means for Solving the Problems) The radiation image processing method of the present invention scans a stimulable phosphor in which radiation image information is accumulated and recorded with excitation light, and emits light emitted from each scanning point by this excitation light. When the radiation image is reproduced as a visible image on the recording medium after photoelectrically reading the exhausted light to obtain the original image signal, the original image signal within the surrounding predetermined range corresponding to each scanning point or this original image signal is reproduced. One blurring mask signal obtained by averaging the image signals obtained by subjecting the image signals to the intermediate processing or a plurality of blurring mask signals obtained by changing the predetermined range is S.
_us.k (k = 1,2, ..., n; n is an integer indicating the number of blur mask signals), the original image signal or an image signal obtained by performing an intermediate process on the original image signal S _b1 , S _b2 , one of the above Alternatively, when one or more attenuation coefficients respectively corresponding to a plurality of blur mask signals are β _k (k = 1, 2, ..., N) and the image signal after the arithmetic processing is S ′, the attenuation coefficient β _At least one attenuation coefficient β ₁ (l is an integer within 1 to n) of _k (k = 1, 2, ..., N) is always within the range of 0 ≦ β ₁ <1 and each radiation image Which is a variable that changes within and is a function of the original image signal or an image signal obtained by subjecting the original image signal to intermediate processing, and using this attenuation coefficient β _l It is characterized in that the spatial frequency component higher than the spatial frequency component _included in the blur mask signal S _us.l corresponding to the attenuation coefficient β _l is attenuated by performing the calculation according to the equation ( ₁ ).

Further, the radiation image processing apparatus of the present invention for carrying out the radiation image processing method scans a stimulable phosphor in which radiation image information is stored and recorded with excitation light, and emits from each scanning point by this excitation light. After photoelectrically reading the stimulated emission light obtained to obtain an original image signal, the original image signal is processed by a calculation unit, and a radiation image is reproduced as a visible image on a recording medium based on the processed image signal. In the radiation image processing device in the radiation image recording / reproducing system, the arithmetic unit averages an original image signal in a predetermined surrounding range corresponding to each scanning point or an image signal obtained by performing an intermediate process on the original image signal. a plurality of unsharp mask signals calculated by changing one or the predetermined range determined by _{S us.k (k = 1,2, ...} , n; n represents a number of unsharp mask signals Number), the original image signal or an image signal subjected to intermediate treatment in the original image signal S _b1,
S _b2 , one or a plurality of attenuation coefficients respectively corresponding to the one or a plurality of blur mask signals, β _k (k = 1,
2, ..., N), where S ′ is the image signal of the arithmetic processing, at least one of the attenuation coefficients β _k (k = 1, 2, ..., N)
The attenuation coefficient β ₁ (l is an integer within 1 to n) is a variable that is always in the range of 0 ≦ β ₁ <1 and changes in each radiographic image, and the original image signal or this original image signal Is a function of the image signal that has undergone intermediate processing in
Using this damping coefficient β _l The calculation is performed according to the equation

(Operation) In the radiation image processing method of the present invention, as described above, at least one of the attenuation coefficients β _k (k = 1, 2, ..., N) has an attenuation coefficient β _l of 0 ≦ β _l <1. A variable having a value within the range, and a function of the original image signal or an image signal obtained by subjecting the original image signal to intermediate processing, The calculation is performed according to the formula.

By transforming (1) above, Becomes

Focusing on the second term β ₁ (S _b2 −S _us.l ) of the equation (2), for example, S _b2 −S _us.l in the parentheses of this term causes the blur mask signal S from the original image signal S _{b2. us.l}
By subtracting S _b2 from the blur mask signal S
_The low spatial frequency components that _us.l has are subtracted.
By subtracting from this _{S b2} -S us.l to 0 ≦ β _l β by multiplying the _<1 attenuation coefficient _{β l l (S b2 -S us.l} ) a further example original image signal S _b1, In the region of β ₁ ≠ 0 (β ₁ is a variable that changes in the radiation image) in the radiographic image, the high spatial frequency component of S _b2 −S _us.l can be attenuated from the signal of S _b1 . This high spatial frequency component is a variable that matches the granular noise of the image, and the attenuation coefficient β ₁ is changed within the range of 0 ≦ β ₁ <1, and the original image signal or this original image signal is subjected to intermediate processing. By setting an appropriate value as a function of the image signal, it is possible to attenuate the granular noise of the image according to the state of each area in one image and to minimize the deterioration of other image quality performance such as sharpness. it can. Further, a radiation image processing apparatus for realizing this calculation method is the same as the radiation image processing apparatus in the radiation image information recording / reproducing system proposed by the present applicant in the above-mentioned JP-A-55-12429 and JP-A-56-11395. In comparison, an apparatus for implementing the radiation image processing method can be realized without making the apparatus particularly complicated, and the calculation time can be set within a sufficiently allowable range. Image signal S
_b1, as the S _b2, may be used an original image signal read photoelectrically to both, subjected to intermediate image processing to the original image signal, even if a signal subjected to the image processing to one or both Good.

Next, the third and fourth terms of the above equation (2) will be described.
Granular noise has a fairly wide range of spatial frequency components. Therefore, if the granular noise cannot be suppressed sufficiently by the combination of the first term and the second term in the above equation (2), or if the spatial frequency band is changed for each area within one screen, more detailed image processing can be performed. When it is desired to perform, for example, the second term is such that the spatial frequency band is changed so that the same operation as the second term can be performed in the third term or the fourth term. In addition, the attenuation coefficient β _m (m ≠ l) is set to β _m <0 in the third and fourth terms, and a specific spatial frequency component proposed by the applicant in JP-A-55-163472 is emphasized. The operations may be combined.

Here, the above-mentioned image processing method is described in
Assuming that the non-sharp mask signal is S _us , the original image signal is S _org , the enhancement coefficient is β, and the processed signal is S ′, proposed in −163472 and the like, S ′ = S _org + β (S _org -S _us ) ... The basic difference from the case of performing the operation of emphasizing a specific spatial frequency component according to the equation (3) will be described.

The simplest formula for the present invention is the first formula of the above formula (2).
Only the term and the second term, that is, S ′ = S _b1 −β ₁ (S _b2 −S _us.l ) ... (4).

As described above, this equation (4) indicates that the spatial frequency component of the granular noise is positively attenuated.

However, it has been clarified by the inventors of the above-mentioned JP-A-55-163472 that the spatial frequency of the particle noise is simultaneously overlapped with the spatial frequency that affects other image quality performances such as sharpness. It can be fully imagined that other image quality performance will be deteriorated to an unrecoverable level if the spatial frequency of the particle noise is positively attenuated. Therefore, conventionally, the granular performance can be improved without actively attenuating the spatial frequency component of particle noise. The image quality is improved by emphasizing the spatial frequency components having a relatively large contribution to the image quality performance such as sharpness rather than the contribution to the image quality.

As a result of further detailed study of the properties of the granular noise, the present inventors subtly select the spatial frequency to be attenuated and the degree of attenuation of this spatial frequency, and positively suppress the spatial frequency component of the granular noise. By doing so, it has been found that the granular noise can be made inconspicuous and deterioration of other image quality performance such as sharpness can be suppressed to the minimum.

By changing the attenuation coefficient β ₁ for performing the above attenuation within the range of 0 ≦ β <1, it is possible to optimize each region in the image for almost all images. This attenuation coefficient β _l is, for example, in a radiation image,
The degree of attenuation is made larger and blurred for areas of low image density where grain noise is relatively conspicuous, and the degree of attenuation is increased for areas of high image density where grain noise is relatively inconspicuous so that the detailed structure becomes clearer. As a function of the image signal, such as reducing the bone part, the lung part, the heart part, etc. in the radiographic image of the chest of the human body,
Various functional forms are selected according to the purpose of the image processing, such as changing the optimum image processing for each object according to each object in one image.

(Examples) Examples of the present invention will be described below with reference to the accompanying drawings.

FIG. 2 is a perspective view showing an example of a radiation image processing apparatus using the radiation image processing method of the present invention.

The stimulable phosphor sheet 1 on which the radiation image information of the subject is stored and recorded is conveyed (sub-scanned) in the arrow Y direction by the sheet conveying means 3 such as an endless belt driven by a motor 2. On the other hand, the excitation light 5 emitted from the laser light source 4 is reflected and deflected by a rotary polygon mirror 6 driven by a motor 13 and rotating at a high speed in the direction of the arrow, and after passing through a focusing lens 7 such as an fθ lens, the optical path is changed by a mirror 8. Instead, the light is incident on the sheet 1 and the main scanning is performed in the arrow X direction substantially perpendicular to the sub-scanning direction (arrow Y direction). From the portion of the sheet 1 irradiated with the excitation light 5, a stimulating luminescent light 9 of a light amount corresponding to the accumulated and recorded radiation image information is diverged, and the stimulating luminescent light 9 is collected by the condenser 10. The light is emitted and photoelectrically detected by a photomultiplier (photomultiplier tube) 11 as a photodetector. The light collector 10 is formed by molding a light guide material such as an acrylic plate, and the linear incident end face 10a extends along the main scanning line on the stimulable phosphor sheet 1. The light receiving surface of the photomultiplier 11 is coupled to the emitting end surface 10b which is arranged and has a ring shape. The stimulated emission light 9 that has entered the light collector 10 from the incident end face 10a proceeds through the inside of the light collector 10 by repeating total reflection, is emitted from the emission end face 10b, and is received by the photomultiplier 11. The photomultiplier 11 detects the amount of stimulated emission light 9 carrying the radiation image information.

The analog output signal S output from the photomultiplier 11 is amplified by the amplifier 16 and digitized by the A / D converter 17 with a predetermined recording scale factor.

The digitized original image signal S _org thus obtained is input to the calculation unit 18, and the calculation unit 18 averages the image signals in the surrounding predetermined range corresponding to each scanning point. Blurred mask signal S _us.k (k = 1,2, ..., n; n
Is the number of blur mask signals obtained by changing the above predetermined range)
Original image signal that is calculated and input to the calculation unit 18
S _org or a signal obtained by subjecting this original image signal S _org to intermediate processing is added to S _b1 , S _b2 and the blur mask signal S.
_Using the attenuation coefficient β _k (k = 1,2, ..., n) prepared in advance as a function of the original image signal S _org in correspondence with _us.k (k = 1,2, ... n), respectively. , The image signal S'after the arithmetic processing is obtained in accordance with the equation.

The simplest calculation process of the formulas (5) is one blur mask signal S _us.l and attenuation coefficient β _l (0 ≦
β _l <1: β _l is a variable and an original image signal or a function of an image signal obtained by performing intermediate processing on the original image signal) is used, and S ′ = S _b1 −β _l (S _b2 −S _us.l ) ... It is a calculation process according to the equation (6). This arithmetic processing means that the spatial frequency component higher than the spatial frequency component of the _blur mask signal S _us.l is attenuated. Apparently, by appropriately selecting the spatial frequency component to be attenuated and the degree of attenuation. It is possible to improve the graininess of an image and suppress deterioration of other image quality performance such as sharpness to a minimum.

The image signal S'computed by the arithmetic unit 18 is stored in the memory 19
The radiographic image based on this image signal is reproduced and displayed on the image display device 20 as necessary.

3A to 3C are block diagrams showing different configuration examples of the arithmetic unit 18 shown in FIG. 2, respectively.

In the configuration example of FIG. 3A, the original image signal S _org is input to the storage means 21 from the left side of the drawing and temporarily stored. The original image signal S _org temporarily stored in the storage means 21 is directly input to the subtraction means 24 described later, and at the same time n blur mask signal calculation means 2 such as the first blur mask signal calculation means 22a.
Input to 2a, 22b, ..., 22n in parallel. In these blur mask signal calculation means 22a, 22b, ..., 22n, N ₁ × N ₁ pieces, N ₂ × N ₂ pieces, ..., N _n × N _n pieces of surrounding scanning points are respectively corresponding to the respective scanning points. Image signals are averaged and the blur mask signals S _us.1 , S
_us.2 , ..., S _us.n are required. These blur mask signals S _us.1 , S _us.2 , ..., S _us.n are input to n attenuation term calculation means 23a, 23b, ..., 23n such as the first attenuation term calculation means 23a, respectively. The attenuation terms β ₁ (S _org −S _us.1 ) and β ₂ (S
_org- S _us.2 ), ..., β _n (S _org- S _us.n ) is calculated. These attenuation terms and the original image signal S _org are input to the subtracting means 24, Is calculated, and the image signal S ′ after the arithmetic processing is obtained.

FIG. 3B is a block diagram showing a configuration example of the calculation unit 18 different from that in FIG. 3A. 3A for the same parts as in FIG. 3A
The same numbers as in the figure are attached and the description is omitted.

In the blur mask signal calculation means 22 'in this configuration example,
First, an average value of 3 × 3 scanning points centering on each scanning point is obtained, and then the average value of these average values is obtained to obtain 9
By calculating the average value of the scanning points of × 9, 15 × 15, etc., the blur mask signal corresponding to each attenuation term calculation means 23a, 23b, ..., 23n is calculated, and each attenuation term calculation means 23a, 23b. , ..., 23n. By doing so, the blur mask signal can be efficiently calculated.

FIG. 3C is a block diagram showing a further different configuration example of the arithmetic unit 18 shown in FIG.

The original image signal S _org is temporarily stored in the storage means 21 ″ and then sent to the blur mask signal calculation means 22 ″. The blur mask signal calculation means 22 ″ calculates the blur mask S _us.1 corresponding to the attenuation coefficient β ₁ based on the original image signal S _org . This blur mask signal S _us.1 is calculated as the attenuation term calculation means 2
Sent to 3 ″, and β ₁ (S _org −S
_us.1 ) is calculated and sent to the subtraction means 24 ″.
At 24 ″, an image signal S ₁ = S _org −β ₁ (S _org −S _us.1 ) obtained by _performing intermediate processing on the original image signal S _org is calculated.

The image signal S ₁ of this calculation result is returned to the storage means 21 ″,
Original image signal S _org stored in the storage means 21 ″
Is stored instead of. The image signal S ₁ is sent to the unsharp mask signal calculating means 22 ', this time the unsharp mask signal S _Us.2 corresponding to the attenuation coefficient beta ₂ on the basis of the image signals S ₁ is calculated, this unsharp mask signal S _Us.2 It is sent to the attenuation term calculation means 23 ″ and β ₂ (S ₁ −S _us.2 ) is calculated. The result of this calculation is sent to the subtraction means 24 ″, and the image signal S ₁ is further subjected to the second intermediate processing. The image signal S ₂ = S ₁ −β ₂ (S ₁ −S
_us.2 ) is calculated.

By repeating the above loop n times, the final processed signal S ′ is S ′ = S _n−1 −β _n (S _n−1 −S _us.n ) (7) Is obtained as.

Thus, the image subjected to intermediate processing signals _{S 1, S 2, ...,} S n-1
The deterioration of image quality performance such as sharpness is minimized by calculating the _blur mask signals S _us.1 , S _us.2 , ..., S _{us.n and} the calculation represented by the equation (7) using The granular noise can be effectively attenuated while being held down.

The above equation (7) uses the same image signal S _n-1 as the image signals S _b1 and S _b2 as compared with the above equation (5), but for example, the attenuation term calculating means shown in FIG. 3C is used. The original image signal S _org is also directly input to and stored in 23 ″, and the image signals S ₁ , S ₂ , ..., S _n−1 after the intermediate processing should be used in the calculation in this attenuation term calculation means 23 ″. Instead, the original image signal S _org is always used to calculate β ₁ (S _org −S _us.1 ) β ₂ (S _org −S _us.2 ) ……………………………… , Finally, when the calculation of S ′ = S _n-1 −β _n (S _org −S _us.n ) (8) is performed, the image signals S _b1 and S _b2 may be different.

FIG. 1A shows that n = 2 using the radiation image processing method of the present invention.
6 is a graph showing an example of calculation in the spatial frequency domain in the case of (a blur mask signal and two attenuation coefficients).
The horizontal axis represents the spatial frequency, and the vertical axis represents the relative value with the DC component being 1. For simplification, the image signal S'after the arithmetic processing is Fourier transformed and the signal shown in the spatial frequency domain is similarly expressed by S '.

Graph A is an ideal graph showing optimum spatial frequency characteristics for suppressing granular noise and suppressing deterioration of other image quality performance such as sharpness to a minimum for a certain radiation image. In contrast to the graph A, the graph A ′ is the blur mask signals S _us.1 and S _us.2 , which are 15 × 15 around each scanning point.
, 5 × 5 scanning points are used as average values, and the attenuation coefficients β ₁ and β ₂ are used.
Using β ₁ = 0.1 and β ₂ = 0.4 respectively, the calculation of S ′ = S _org −β ₁ (S _org −S _us.1 ) −β ₂ (S _org −S _us.2 ) …… (9) It is a graph showing the result in the spatial frequency domain, which is sufficiently similar to graph A.

Graph B is an ideal graph showing optimum spatial frequency characteristics for other radiation images. In contrast to the graph B, the graph B ′ uses the average values of 15 × 15 scanning points and 3 × 3 scanning points around each scanning point as the _blur mask signals S _us.1 and S _us.2 , respectively, and the attenuation coefficient β ₁ , β ₂ = β ₁ = 0.1, β ₂ = 0.
8 is a graph showing the calculation result of S ′ = S _org −β ₁ (S _org −S _us.1 ) −β ₂ (S _org −S _us.2 ) ... (10) in the spatial frequency domain. Is. Also in this case, the graph B ′ is sufficiently close to the graph B.

As shown in FIG. 1A, the damping coefficient β _k (k = 1, 2, ...,
The optimum value of n) is first determined by the type of radiographic image.

FIG. 1B is a graph showing an example of a function of the attenuation coefficient β _{1 with} the image signal as a variable. This graph shows that the attenuation coefficient β for the region C where the image density where the granular noise is relatively conspicuous is low.
Blur with _l = α, and stop blurring with β = 0 so that the detailed structure becomes clear in the area E with high image density where grain noise is relatively inconspicuous, and with respect to the intermediate area D, the higher the image density becomes It means to reduce β _l . As shown in FIG. 1B, by changing β _k (k = 1,2, ..., n) in one radiation image according to the image signal, β of the same value can be obtained for one radiation image as a whole. Compared with the case of using _k (k = 1,2, ..., n),
More detailed image processing can be performed.

Note that the graph of FIG. 1B is merely one example, and for example, the attenuation coefficient β ₁ may be changed in a curve with respect to the image signal, and it is appropriate depending on the type of radiation image, the purpose of image processing, and the like. Different functional forms are defined.

In addition, β _k (k = 1, 2, ..., N) does not have to be a function of the image signal, but can be changed as a function of the image signal even if it is changed for each subject in one image as described above. As in the case
More detailed image processing can be performed.

In this way, the radiation image is divided according to the type of the subject in the entire image (for example, the chest of the human body, the head, etc.), the intensity of the radiation applied to the subject, and the blur mask signal is calculated so as to match each radiation image. By determining the method and the function of the attenuation coefficient and performing the arithmetic processing according to the above-mentioned method, the granular noise of the radiation image is effectively attenuated according to each area in the image and other image quality performance such as sharpness is obtained. It is possible to obtain a reproduced image in which the deterioration of is minimized.

(Effect of the Invention) The present invention scans a stimulable phosphor in which radiation image information is stored and recorded with excitation light, photoelectrically reads stimulated emission light emitted from each scanning point by this excitation light, and reads the original. After obtaining the image signal, the attenuation coefficient β _k (k = 1, 2, ...,
n), at least one damping coefficient β _l is always 0 ≦
It is a variable that is within the range of β ₁ <1 and changes in each image, and is a function of the original image signal or an image signal obtained by subjecting this original image signal to intermediate processing, and using this attenuation coefficient β _l hand Since the calculation is performed according to the equation, the spatial frequency component higher than the spatial frequency component of the _blur mask signal _Sus.l can be attenuated, and the granularity of the radiation image can be changed according to each region in the radiation image. The noise can be effectively attenuated, and deterioration of other image quality performance can be suppressed to a minimum. Further, the apparatus for carrying out this method is not particularly complicated, and the calculation time can be set within a sufficiently acceptable range.

[Brief description of drawings]

FIG. 1A is a graph showing an example calculated in the spatial frequency domain using the radiation image processing method of the present invention, and FIG. 1B is a graph showing an example of a function of the attenuation coefficient β ₁ with the image signal as a variable. 2, FIG. 2 is a perspective view showing an example of a radiation image processing apparatus for carrying out the radiation image processing method of the present invention, and FIGS. 3A to 3C are different configuration examples of the arithmetic units shown in FIG. It is the block diagram shown. 1 ... Accumulative phosphor sheet 2, 13 ... Motor, 3 ... Sheet transport means 4 ... Laser, 6 ... Rotating polygon mirror 9 ... Stimulated emission light, 10 ... Concentrator 11 ... Photo Multiplier 16 …… Amplifier, 17 …… A / D converter 18 …… Calculator, 19 …… Memory 20 …… Image display device

─────────────────────────────────────────────────── ─── Continuation of the front page (51) Int.Cl. ⁶ Identification code Internal reference number FI technical display location 9163-4C A61B 6/00 350 N

Claims (5)

[Claims] 1. An original image signal is obtained by scanning a stimulable phosphor in which radiation image information is stored and recorded with excitation light and photoelectrically reading stimulated emission light emitted from each scanning point by the excitation light. Then, when reproducing the radiation image as a visible image on the recording medium, the original image signal in the surrounding predetermined range corresponding to each scanning point or the image signal obtained by performing the intermediate processing on the original image signal is averaged. One of the blur mask signals obtained by changing the predetermined range or the blur mask signal obtained by
_us.k (k = 1,2, ..., n; n is an integer indicating the number of blur mask signals), the original image signal or an image signal obtained by subjecting the original image signal to intermediate processing S _b1 , S _b2 , 1
1 corresponding to each or a plurality of blur mask signals
Number or a plurality of attenuation coefficient _{β k (k = 1,2, ...} , n), an image signal after the processing is taken as S ', the attenuation coefficient _{β k (k = 1,2, ...} , at least 1 out of n)
The attenuation coefficient β ₁ (l is an integer within 1 to n) is always in the range of 0 ≦ β ₁ <1 and is a variable that changes in each radiographic image, and the original image signal or the original image signal It is a function of the image signal that has undergone intermediate processing on the image signal, and using this attenuation coefficient β _l The radiation image processing method is characterized in that the spatial frequency component higher than the spatial frequency component _included in the blur mask signal S _us.l corresponding to the attenuation coefficient β _l is attenuated by performing the calculation according to 2. The original image signal or image signals S _b1 and S _{b2 obtained} by subjecting the original image signal to intermediate processing,
The radiation image processing method according to claim 1, wherein the original image signals are the same. 3. The original image signal or image signals S _b1 and S _{b2 obtained} by subjecting the original image signal to intermediate processing,
The radiation image processing method according to claim 1, wherein both are the same image signal obtained by subjecting the original image signal to the same intermediate processing. 4. The original image signal or one of the image signals S _b1 and S _{b2 obtained} by subjecting the original image signal to an intermediate process is subjected to a first intermediate process to the original image signal or the original image signal. The radiation image processing method according to claim 1, characterized in that the other is an image signal obtained by performing second intermediate processing on the original image signal. 5. An original image signal is obtained by scanning a stimulable phosphor in which radiation image information is stored and recorded with excitation light and photoelectrically reading stimulated emission light emitted from each scanning point by the excitation light. Then, in the radiation image processing apparatus in the radiation image recording / reproducing system that processes the original image signal in the operation unit and reproduces the radiation image as a visible image on the recording medium based on the processed image signal, the operation unit is , One obtained by averaging an original image signal within a predetermined range of surroundings corresponding to each scanning point or an image signal obtained by subjecting the original image signal to intermediate processing, or a plurality obtained by changing the predetermined range the number of unsharp mask signal _{S us.k (k = 1,2, ...} , n; n is an integer indicating the number of unsharp mask signal), medium in the original image signal or the original image signal The image signal subjected to processing S _b1,
S _b2 , one or a plurality of attenuation coefficients respectively corresponding to the one or a plurality of blur mask signals, β _k (k = 1,
2, ..., N), where S ′ is the image signal after the arithmetic processing, at least 1 of the attenuation coefficients β _k (k = 1, 2, ..., N)
The attenuation coefficient β ₁ (l is an integer within 1 to n) is always in the range of 0 ≦ β ₁ <1 and is a variable that changes in each radiographic image, and the original image signal or the original image signal It is a function of the image signal that has undergone intermediate processing on the image signal, and using this attenuation coefficient β _l A radiation image processing apparatus, characterized in that the calculation is performed in accordance with