JPH01145782A - Method and device for processing radiograph - Google Patents

Abstract

PURPOSE:To prevent the generation of a pseudo picture and to improve a diagnostic performance by providing an intensifying coefficient variable means for changing the intensifying coefficient and executing an operation by the use of the changed intensifying coefficient. CONSTITUTION:When an accumulating phosphor material is scanned, medically diagnostic radiograph information recorded thereon is read and converted to an electrical signal and then reproduced as a visual image, a non-sharp mask signal Sus corresponding to an extremely low spatial frequency at respective scanning points is obtained, when an original signal read from the phosphor material is defined to be Sorg, the intensifying coefficient to be beta, and an output signal used for a reproduction to be S', the intensifying coefficient beta is changed so as to be monotonously increased, reach a maximum value and be monotonously decreased according to the increase of the value of the original picture signal Sorg or the non sharp mask signal Sus when the intensifying coefficient betais zero or positive to operate according to S'=Sorg+beta(Sorg-Sus) and intensify a frequency component above the extremely low spatial frequency. Thereby, the diagnostic performance of the radiation picture can be remarkably improved.

Description

DETAILED DESCRIPTION OF THE INVENTION The present invention relates to an image processing method and apparatus in a radiographic image information recording and reproducing system used for medical diagnosis, and more specifically to an image processing method and apparatus using a stimulable phosphor material (hereinafter simply referred to as a "phosphor") as an intermediate medium. An image processing method in a radiographic image information recording and reproducing system that records a radiographic image on a radiographic imager, reads out and reproduces the radiographic image, and records it as a final image on a recording material, and an apparatus that implements the method. It is related to.

Such radiographic image information recording/reproducing method is as proposed by the present applicant in Japanese Patent Publication No. 131-29490.
The radiation transmitted through the object is absorbed by a phosphor, and then this phosphor is excited with a certain kind of energy to emit the accumulated radiation energy as fluorescent light. There are ways to detect and image it.

The radiographic image information recording and reproducing system using this phosphor has extremely high utility value in that it can record images over a wide radiation exposure range compared to conventional radiographic systems using silver halide photography. Therefore, it is particularly useful as an X-ray photography system for the human body.

X-rays are harmful to the human body if the exposure dose is high, so -
It is desirable to obtain as much information as possible in a single X-ray. However, current X-ray photographic films are required to have both suitability for photographing and suitability for observation and interpretation, and as a result, they are designed to meet both requirements to some extent. For this reason, there is a problem in terms of photographic suitability in that the X-ray exposure range is not wide enough (one solution to this problem is the X-ray image recording method using the aforementioned phosphor), and the current X-ray Regarding the suitability of radiographic film for observation and interpretation, there is a problem in that the image quality is not necessarily sufficient for diagnosis.

In view of the above circumstances, the present invention provides an X-ray image recording method using a phosphor, in which a non-sharp mask process is applied when reading out the X-ray image information recorded on the phosphor and reproducing it as a visible image. An object of the present invention is to provide an X-ray image processing method and apparatus that improve diagnostic performance of X-ray images.

As a result of research on the frequencies to be emphasized and the diagnostic performance of the obtained X-ray images, the present inventors found that the frequencies that are important for diagnosis differ slightly depending on each part of the human body, but compared to the conventional sense. In other words, it was found to be in a very low frequency (hereinafter referred to as "ultra-low frequency") region. It has also been found that the conventional method of improving sharpness by emphasizing high frequency components only emphasizes noise components in the case of X-ray image processing, which actually tends to degrade diagnostic performance. In the high frequency region, the proportion of noise is high, and if you reduce the emphasis on this high frequency region,
I also found that the noise was less noticeable and it was easier to see.

Therefore, the present inventor proposed a radiation image processing method that can improve diagnostic performance by emphasizing very low frequency components that are effective for diagnosis and increasing contrast.

(Special Publication No. 62-62373) This is a radiation image processing method that emphasizes extremely low frequency components and at the same time relatively reduces high frequency components that account for a large proportion of noise, making it possible to obtain visually easy-to-read images. It's a method.

This method scans the phosphor with excitation light, reads out the radiation image information recorded on it, converts it into an electrical signal, and then reproduces it as a visible image. Find the unsharp mask signal Sus corresponding to the low frequency,
The original image signal read out from the phosphor is
It is characterized by converting the signal by the calculation of S'=Sorg+β(Sorg-5us), where β is the s emphasis coefficient and S' is the reproduced image signal, thereby emphasizing the frequency components above the ultra-low frequency. This is a radiation image processing method.

Here, the unsharp mask signal Sus corresponding to the very low frequency
refers to a signal corresponding to the density of each scanning point of a non-sharp image (hereinafter referred to as "non-sharp mask") obtained by blurring the original image so that it contains only frequency components lower than the ultra-low n wavenumber component. For this unsharp mask, when the modulation transfer function has a spatial frequency of 0.01 cycle/order, 0.5
In the above, the one that is 0.5 or less when the spatial frequency is 0.5 cycles/m+s is used. Here, the ultra-low spatial frequency means a spatial frequency of about 0.5 cycles/mm or less.

The maximum value of the modulation transfer function of the system that provides the visible image created based on the signal emphasized by the above calculation formula is set to be 1°5 to 10 times the value of the modulation transfer function near zero frequency. It is desirable to do so.

In addition, there is a lot of noise in the high frequency region, making it difficult to see the 0,
The modulation transfer function is 0 in the frequency range of 5 to 5 cycles/am.
, 5 or less is performed on the S1. This smoothing process averages out noise components, resulting in an image that is easy to view.

The non-sharp mask can be created by the following various methods.

The first method is to store the original image signal at each scanning point, read it out together with peripheral data according to the size of the unsharp mask, and calculate the average value (simple average or average value based on various weighted averages) of Sus. This is a method to find.

The second method is to read out the original image signal using a light beam with a small diameter, and then, if the accumulated image remains, use a light beam with a large diameter that matches the size of the non-sharp mask to read out the signal at each scanning point. In this method, the signal is averaged together with the surrounding signals and read out.

The third method takes advantage of the fact that the beam diameter of the readout light beam gradually expands due to scattering in the phosphor layer.The original image signal Sorg is created using the light emission signal from the incident side of the light beam, and the light beam A non-sharp mask signal Sus is generated by light emission on the side through which the light is transmitted. In this case, the size of the unsharp mask can be controlled by changing the degree of light scattering of the phosphor layer and by changing the size of the aperture that receives this light.

In the present invention, a phosphor refers to a phosphor that is irradiated with the first light or high-energy radiation and then is stimulated (excited) thermally, mechanically, chemically, electrically, etc. A phosphor that exhibits so-called photostimulability, which re-emits light corresponding to the amount of radiation irradiated.

Here, light includes visible light, ultraviolet light, and infrared light among electromagnetic radiation, and high-energy radiation includes X-rays, gamma rays, beta rays, alpha rays, neutron rays, and the like. Excitation is 500~
It is desirable to use light in the wavelength range of 800 nIIl, preferably 600 to 700 nIIl, and the excitation light in this wavelength range can be obtained by selecting an excitation light source that emits light in this wavelength range, or by using excitation light having a peak in the above wavelength range. It can be obtained by using a combination of a light source and a filter that cuts light outside the wavelength range of 500 to 800 nI11.

Excitation light sources that can emit light in the above wavelength range include lasers (647nn+), various light emitting diodes, He-Ne lasers (B33na+), and rhodamine B-dye lasers. ranges from near ultraviolet to visible to infrared,
It can be used in combination with a filter that transmits light in the wavelength range of 500 to 800 nm.

The ratio of excitation energy to emission energy is 104:1-1
Since the ratio is normally about 0.6:1, when excitation light enters the photodetector, the S/N ratio is extremely reduced. The above-mentioned decrease in the S/N ratio can be prevented by setting the emitted light at the short wavelength side and setting the excitation light at the long wavelength side and separating them as much as possible so that the excitation light does not enter the photodetector.

For this purpose, it is desirable to use a phosphor whose emitted light is in the wavelength range of 300 to 500 r+m.

Examples of phosphors that emit light in the wavelength range of 300 to 500 nm include rare earth element-activated alkaline earth metal fluorohalide phosphors [specifically, Japanese Patent Publication No. 60-42
It is described in the specification of No. 837 (Ba l-ty
+ Mgx, Ca)') F X : a E u
” (However, X is at least one of C exemption and B
, and X and y are Q < xy≦0.6 and x
y≠0, and a is 10-6≦a≦5XlO-2); FX:yA (However, Ml is Mg5Ca, Sr.

At least one of 2 mouths and Cd, X is C, B
At least one of r and ■, A is Eu5Tbs
Ces Tl11% Dy5Prs HolNd%Y
at least one of b and Er, X is 0≦X≦o
, e, y is 0≦y≦0.2), etc.];
ZnS: Cu, P described in specification No.-9542
b, BaC1xAQ, 203:Eu (however, 0,8≦X
≦10) and M”O・x Si 02 :A (however, Ml is Mgs ca, 5rSZns Cd or Ba
and A is Cc STb s Eu STm 5Pb
ST51. , B1 or Mn, and X is 0.5≦X≦
2.5); and LnOX:xA described in Japanese Patent Publication No. 59-44339 (however, Ln is La.

at least one of YSGd and Lu, X is C9
J and B", A is at least one of Ce and Tb, X is 0 < X < Q, L
)-, etc. Among these, rare earth element-activated alkaline earth metal fluorohalide phosphors are preferred, and among these, barium fluorohalides shown as specific examples are particularly preferred because of their excellent stimulable luminescence.

In addition, when the phosphor layer of a stimulable phosphor plate made using this stimulable phosphor is made of stone using pigment or dye,
The sharpness of the final image is improved and a desirable result is obtained. (Japanese Patent Publication No. 59-23400) The above-mentioned Japanese Patent Publication No. 62-82373 states that when performing calculations for non-sharp mask processing, if β is made smaller in the low luminance region and β is increased in the high luminance region, frequency emphasis can be achieved. It is disclosed that it is possible to prevent false images that are likely to occur.

- As an example, if we perform the frequency processing on an original photograph of the stomach using barium contrast agent with a fixed enhancement coefficient β, the boundaries of wide low-intensity areas containing a large amount of contrast agent will be The image is emphasized more than necessary, resulting in a double-contoured false image. Instead, the occurrence of the double contour can be prevented by making the enhancement coefficient β variable, that is, decreasing β in a low-brightness area where a large amount of contrast agent is present, and increasing β in a high-brightness area such as a gastric subdivision. False images have a huge psychological impact on diagnosis. For example, if a false image occurs in a certain area, the user may suspect that false images also occur in other easily diagnosed areas. As another example, in the case of frontal chest photography, if β is fixed, noise will increase in the low-brightness region of the spine and heart, and in extreme cases, details will be washed out. This is visually very noticeable and has a negative impact on diagnostic performance. Similarly, if β is made smaller in the low-luminance region of the spine and the heart, and β is made larger in the high-luminance region of the lung field, the increase in noise and white areas described above can be prevented.

In this way, by continuously changing the emphasis coefficient β according to the luminance, it is possible to improve diagnostic performance while preventing the generation of false images. However, in the above-mentioned Japanese Patent Publication No. 62-62373, the emphasis coefficient β is changed monotonically (βl≧0
) is limited to.

As a result of further research, the inventor found that when the emphasis coefficient β is positive or 0 (β≧0), including the case where the differential with respect to luminance is negative (β1×O) improves diagnostic performance. I found that it improves.

In other words, it has been discovered that it is preferable to make changes that include a portion where the emphasis coefficient β is decreased as the luminance increases. For example, in the case of intestinal gas areas in abdominal contrast imaging, high-brightness areas that are not necessary for diagnosis are emphasized, making them too conspicuous and greatly hindering viewing of the area to be diagnosed.

In this case as well, if the emphasis coefficient β is changed so that the low-brightness portion is emphasized more than the high-brightness portion depending on the brightness, an image that is easy to see without the high-brightness portion being emphasized can be obtained.

The present invention provides a method in which diagnostic performance is further improved by changing the emphasis coefficient β, and an apparatus for implementing the method.

The emphasis coefficient β can be changed using low-luminance enhancement (Figure 1C), in which β decreases as the luminance increases, and medium-luminance enhancement (Figure 1D), in which β is increased at low luminance and decreased at high luminance. ) is possible. In low-luminance enhancement, β is kept constant at a large value in low-luminance areas, and β is kept constant in high-luminance areas.
There is a two-stage type in which β is kept constant at a small value and β is continuously decreased during this period (Fig. 1C a), and a curved type in which β is continuously decreased so that the absolute value of negative β' becomes larger (Fig. 1C). 1C diagram b
) are possible.

Hereinafter, the present invention will be explained in detail with reference to the drawings.

Figures 1A to ID are diagrams showing various ways of changing β in response to changes in luminance, including the flat type (constant par) in Figure 1A and the monotonically increasing type (β'≧0) in Figure 1B. Special Public Service 1982-623
Included in the method of '73. The σ low luminance enhancement shown in FIG. 1C and the medium luminance enhancement shown in FIG. 1D are based on the method of the present invention, and both have a two-stage type (curve a) and a curved type (curve b).

The low brightness enhancement shown in FIG. 1C is suitable for cases where diagnosis of low brightness areas is particularly important and the low brightness areas do not occupy a large portion of the entire image. For example, this applies to angiography and lymphangiography, and in these radiographic images, it is desirable to greatly improve the sharpness of the desired area even if the noise increases slightly, so this low brightness enhancement greatly improves diagnostic performance. improve.

Mid-brightness enhancement in Figure 1D is suitable when low-brightness and high-brightness areas occupy a considerable portion of the entire image, and these areas are not diagnostically important, while medium-brightness areas are particularly diagnostically important. ing. For example, cholecystography and hepatography correspond to this case, and if noise or gas areas are emphasized in these radiographic images, it will interfere with diagnosis. It is desirable to emphasize.

Table 1 shows the results of experiments in which the above-mentioned various β changes were carried out and the images obtained from the experiments were evaluated by a radiologist. Symbols A, B5C5D in the table are 1A, IB, and IC, respectively.
Represents the type emphasis of the ID diagram.

From Table 1, it can be seen that type C2D of enhancement by the method of the present invention is evaluated higher than type A in all cases, and higher than type B (monotonically increasing) in most cases.

Next, specific embodiments of the method of the present invention will be described in detail with reference to the drawings.

As shown in FIG. 2, a stimulable phosphor sheet 1 on which a radiation image has been accumulated and recorded by radiography is fed by a roller 2. This roller 2 feeds the sheet 1 in the direction of arrow A and performs sub-scanning for image reading from the sheet 1. The main scanning is performed by scanning a laser beam from a laser light source 3 for excitation light in the direction of arrow B with a scanning mirror 3a. The stimulated light emitted by the scanning of the excitation light is collected by a light collector 4a made of a light-guiding sheet material and detected by a photodetector 4 such as a photomultiplier placed at the output end of the light collector 4a. converted into an electrical signal. This electrical signal is amplified by an amplifier 5, converted to a digital signal by an A/D converter 6, and stored on a magnetic tape 7.

The digital signals of each part stored on this magnetic tape 7 are
Su
After calculating s, the above-mentioned S'=Sorg+β(Sorg
-5 us) is performed.

The Sus has a modulation transfer function of 0.01 cycles/+n
A value must be specified that is 0.5 or more when the spatial frequency is +e and 0.5 or less when the spatial frequency is 0.5 cycles/mm. Furthermore, when calculating the above equation, it is necessary to specify the emphasis coefficient β. These values may be specified individually from the outside, or several types may be determined for each part of the human body or for each case, and these values may be stored in the memory of the computing device.

Smoothing processing is performed on the S' to reduce frequency components higher than ultra-low spatial frequencies. This smoothing process can reduce noise without damaging information necessary for diagnosis.

This smoothing process will be explained in more detail with reference to FIG.

FIG. 3(a) shows the frequency response when the accumulated image on the phosphor is sampled at 10 pixels/■. It is known that this curve becomes a 5ine curve when a rectangular aperture is used as the aperture of the photodetector, and becomes a Gaussian distribution curve when a Gaussian distribution aperture is used.

Figure 3(b) shows a modulation transfer function such that the modulation transfer function is 0.5 or more when the spatial frequency is o, 6t cycles/mm, and 0.5 or less when the spatial frequency is 0.5 cycles/mm. A rectangular unsharp mask (1) and a Gaussian distribution unsharp mask (II) are shown.

In this example (1), when the image on the phosphor is sampled at a rate of 10 pixels, approximately 63 pixels x 63 pixels (this can be referred to as "
This is a case where an unsharp mask is created by taking a simple average of the unsharp mask size N-83J). This is equivalent to scanning the pixels on the phosphor with a large light beam of 6.3 mm x 8.3 mm. Note that here f'
c is 085 to 0.0 when the modulation transfer function is 0.5.
The value of an arbitrary frequency included in the extremely low frequency region of 01 cycles/llll11 is shown.

On the other hand, the Gaussian unsharp mask (n) is basically the same as the rectangular unsharp mask (I) except that a Gaussian distribution weight is applied when averaging the pixels. The non-sharp masks (1) and (If) differ mainly in the shape on the high frequency side, but
The difference in the effect of ultra-low frequency processing due to this difference is extremely small.

FIG. 3(C) is a graph showing the modulation transfer function after calculating (Sorg-Sus).

The solid line (1) in FIG. 3(d) shows the calculation result S'. Here, β is set to "3". As a result of the above calculation, the maximum value (B) of the modulation transfer function of the enhanced image signal is approximately 4.0% of the modulation transfer function (A) near zero frequency.
It is 6 times more.

The dotted line (II) in FIG. 3(d) shows the modulation transfer function when S' in FIG. 3(d) is subjected to a smoothing process of 5 pixels by 5 pixels.

In such unsharp mask processing, the emphasis coefficient β is
Diagnostic performance can be improved as described above by changing the brightness of the line image.

Details of examples of how to change β (FIGS. 1C and 1D) are shown in FIGS. 4 and 5. Figure 4 corresponds to the case of the two-stage type (curve a) of low luminance emphasis (Figure 1C), where β is the maximum value βll1aX to the minimum value βmin between luminances A and B.
has decreased to In other words, the emphasis coefficient is increased (βmax) in the low luminance region (from Smin to A).
However, in the high brightness region (from B to 5Ilax), it is small (βmin). Brightness A is the minimum brightness (Sa+
In), maximum brightness (Smax) and minimum brightness (Smln
) and 0.2 to 0.5 times the difference (ΔS) [
Sa+in + (0,2~0.5)ΔS] is good, and the brightness B is the same <0.7~1 times the size [Dmln
+(0,7-1)ΔD] is good.

Figure 5 shows the case of medium brightness: A (Figure 1D), and in the case of the two-stage type indicated by the solid line a, β is the first
From the minimum value (βmini) to the maximum value (βl1laX)
and decreases between C and D from the maximum value (βll1ax) to the second minimum value (βmin2). That is, the low brightness area (from S min to A) and the high brightness area (
D to 5IIIax), the emphasis coefficient is small (βminl, βmin 2) L, and in the medium brightness region (B to C) it is large (βmax). Here, the first minimum value (βmini) and the second minimum value (βmin 2) may be equal.

- In the case of the dashed dotted line, β increases between brightness A and E, and decreases between brightness E and D. Brightness A is the minimum brightness (S
min), the maximum brightness (Smax) and the minimum brightness (Smi
n) plus 0 to 0.2 times the difference (ΔS) [5
rAin + (0-0.2)ΔS], brightness B is the average brightness or statistical average value) to the difference (ΔS) from 0 to 0.
The size obtained by subtracting 2 times [5-(0 to 0.2)ΔS], the brightness E is the average brightness (S), and the brightness C is the difference (ΔS) from the average brightness.
) [S+(Q~0.2)
ΔS], the brightness is calculated from the maximum brightness (Smax) by the difference (Δ
S) minus 0-0.2 times [Smax - (0
~0.2) ΔS is desirable.

Here, the emphasis coefficient β can be arbitrarily selected in the range of about 0 to 10.

The above calculations are performed by the calculation device 8 shown in FIG. 2, and a block diagram of this calculation process is shown in FIG. Sorg
By multiplying the difference between and Sus (Sorg-Sus) by the gain (β) set by the value of Sorg and adding it to S org, the desired S'-3Org+β(So
rg-Sus) is obtained.

At this time, the values of βmax and βn+in are either stored in the memory of the arithmetic unit 8 or inputted from the outside to be given to the gain setting circuit 20.

In addition, the monotonically increasing region of β in Figs. 4 and 5 (C in Fig. 5)
-D, E-D) and a monotonically decreasing region (A-8 in Figure 4,
The monotonically increasing/decreasing function β-f (D) between A-B and A-E in FIG. 5 must also be stored in the memory of the arithmetic unit 8 or input from the outside.

In addition, in the above calculation, the maximum brightness (Sa+ax) and the minimum brightness (Smin) both correspond to the maximum and minimum in the target substantial image, and parts other than the image are larger than this or There may also be a small brightness. Note that depending on the case, the maximum and minimum values among all the screens may be simply taken.

Further, the circuit for the above calculation process is not limited to the one shown in FIG. 6, but any circuit may be used as long as it ultimately gives the same result as the calculation of the above equation.

By changing the emphasis coefficient β according to the luminance using the method described above, the diagnostic performance of radiographic images can be significantly improved as described above.

In addition, in experiments conducted by the present inventors, the effect of changing β with the original image signal of the image on the phosphor and the case with changing β with a non-sharp mask signal was different from the difference between lines, etc. Ta. (In this specification, both cases are included and explained using the brightness of the original image as a representative.) 10 Ultra-low frequency processing that may perform gradation processing at the same time as the frequency enhancement described above However, the effect is relatively small on diseases where brightness changes slowly over a large area, such as lung cancer and breast cancer. For these cases, diagnostic performance can be improved by increasing the contrast in combination with gradation processing. This gradation processing may be performed either before or after the ultra-low frequency processing. Before ultra-low frequency processing, gradation processing is performed using a nonlinear analog circuit, and then A/D conversion is performed. If it is performed after A/D conversion, digital processing can also be performed using a minicomputer.

Also, after ultra-low frequency processing, do digital processing or D
/A conversion converter analog processing. The data, which has undergone frequency enhancement and gradation processing as necessary, is recorded on a magnetic tape 7 (see FIG. 2). The data on the magnetic tape 7 is sequentially read out, converted into an analog signal by the D/A converter 9, and amplified by the amplifier IO.
The image is sent to the computer and reproduced as a visible image on a CRT, making CRT diagnosis possible. Further, this image data is input to the recording light source 11 as necessary. Light generated from a recording light source 11 passes through a lens 12 and is irradiated onto a recording material 13, such as a photographic film, mounted on a printing drum 14.

In this case, the radiographic image is reproduced on photographic film,
Diagnosis can also be made by observing this image. Above C
When reproducing and recording images on RT or photographic film, a reduced photographic image can be obtained by recording at a higher sampling frequency than during input scanning. For example, in the input system, 10 pixels/m11
In a single output system, scanning at 20 pixels/mm results in a photographic image reduced to 1/2. As will be described later, a photographic image reduced to 172 to 1/3 will appear to have a higher contrast and will be much easier to see.

The present invention is not limited to the embodiments described above, and various configuration changes are possible.

The image on the phosphor can be read out by a method of setting the phosphor on a rotating drum, a method of two-dimensional scanning in a plane, or an electronic scanning method such as a flying spot scanner. Further, the unsharp mask calculation can be performed by digitally processing only the sub-scanning direction by unsharpening the analog signal using a low-pass filter in the main scanning direction before A/D conversion. Further, the above calculation may be performed offline after all data is stored on the magnetic tape, or may be partially stored in a core memory and sequentially processed online. For example, the image data may be temporarily stored on a disk or magnetic tape to create an integral histogram, and the position of the bending point of the β variable may be determined from this integral histogram.

Alternatively, for example, the data may be output and displayed on a CRT monitor, and the brightness position where the bending point is desired may be inputted using a light pen or the like.

In the embodiments described above, the images reproduced on the CRT are recorded on photographic film as necessary, but diazo film, electrophotographic materials, etc. can also be used as the recording material. Furthermore, the present invention provides an X-ray photographic system using, as an original recording medium, a photographic film with an average gamma of 0.3 to 1.5 capable of recording radiation energy (which the present applicant previously disclosed in Japanese Patent Publication No. 81-5193). ) can also be applied.

[Brief explanation of the drawing]

T, IA, IB, IC, ID are graphs showing various examples of changing the emphasis coefficient β in non-sharp mask processing according to image brightness, and Fig. 2 shows the system of a system implementing the method of the present invention. A schematic diagram, FIG. 3 is a graph showing the modulation transfer function of each stage of signal processing in the unsharp mask processing used in the present invention, and FIGS. 4 and 5 show the method of changing the emphasis coefficient β in the present invention. A graph showing the relationship between emphasis coefficient and brightness, and FIG. 6 is a block diagram showing an example of an arithmetic device used in the present invention. 1... Fluorescent sheet 3... Laser light source 4
... Photodetector 11 ... Recording light source 13.
...Recording materials 15... CRT Figure 1A Old figure 1C Figure 1D Figure 2 7fmm) Fig. 6 At the same time, there is a request for examination of the application and there are procedures.

Claims (1)

[Claims] 1) After scanning the stimulable phosphor material and reading out the radiation image information recorded therein and converting it into an electrical signal,
When reproducing as a visible image, the unsharp mask signal Sus corresponding to the ultra-low spatial frequency at each scanning point is obtained, the original image signal read from the phosphor is set as Sorg, the emphasis coefficient is set as β, and the reproduction is performed using In the method of emphasizing the frequency components above the ultra-low spatial frequency by performing the calculation S' = Sorg + β (Sorg - Sus), where the output signal to be used is S', the emphasis coefficient β is When the coefficient β is zero or positive (β≧0), the differential (β′) with respect to the luminance of the original image or unsharp mask becomes negative (including the part where β′<0) according to the magnitude of the signal. A radiation image processing method characterized by changing the radiation image. 2) The radiation image processing method according to claim 1, wherein the visible image is reproduced on a CRT. 3) As a non-sharp mask, the modulation transfer function is 0.5 or more at a spatial frequency of 0.01 cycles/mm, and 0
．． 3. The radiation image processing method according to claim 1, wherein a non-sharp mask having a sharpness of 0.5 or less at a spatial frequency of 5 cycles/mm is used. 4) A first brightness value (
A) The maximum value (βmax) is taken in the following low luminance region, and the maximum luminance value (βmax) is greater than the first luminance value (A).
The minimum value (βmin) is taken in a high brightness region equal to or higher than the second brightness value (B) which is not larger than the first brightness value (Smax), and
The radiation image processing according to any one of claims 1 to 3, characterized in that the luminance value (A) and the second luminance value (B) monotonically decrease. Method. 5) The first brightness value (A) is equal to the minimum brightness value (Sm
in) plus 0.2 to 0.5 times the difference (ΔS) between the maximum brightness value (Smax) and the minimum brightness value (Smin). The radiation image processing method described in Section 1. 6) The second brightness value (B) is equal to the minimum brightness value (Sm
in) plus 0.7 to 1 times the difference (ΔS) between the maximum brightness value (Smax) and the minimum brightness value (Smin). 5
The radiation image processing method described in Section 1. 7) The enhancement coefficient β has a first minimum value (βmin1) in a low brightness region equal to or larger than a first brightness value (A) that is equal to or larger than the minimum value (Smin) of brightness of the original image or the unsharp mask. A second minimum value (βmin2) is taken in a high brightness region equal to or higher than a second brightness value (D) that is equal to or smaller than the maximum brightness value (Smax), and the first brightness value (A) and the second Claims 1 to 3 are characterized in that between the two brightness values (D), the brightness increases from the low brightness side to the high brightness side to the maximum value (βmax), and then decreases. The radiation image processing method according to any one of the above. 8) The first minimum value (βmin1) and the second minimum value (
8. The radiation image processing method according to claim 7, wherein βmin2) are equal to each other. 9) The first brightness value (A) is the minimum brightness value (Sm
in) plus 0 to 0.2 times the difference (ΔS) between the maximum brightness value (Smax) and the minimum brightness value (Smin), and the second brightness value (D) is Maximum brightness value (
9. The radiation image processing method according to claim 7, wherein the value is a value obtained by subtracting 0 to 0.2 times the difference (ΔS) from Smax). 10) The emphasis coefficient β is the first minimum value (βmin1) between the third brightness value (B) larger than the first brightness value (A) and the first brightness value (A). to the maximum value (βmax), and this third luminance value (B)
and a fourth brightness value (C) smaller than the second brightness value (D).
) between the fourth brightness value (C) and the second brightness value (D), and between the fourth brightness value (C) and the second brightness value (D) from the maximum value (βmax) to the second minimum value (βmin2)
Claim 7 characterized in that monotonically decreases to
9. The radiation image processing method according to any one of paragraphs 9 to 9. 11) The third brightness value (B) is the average value of image brightness (S
), the maximum brightness value (Smax) and the minimum brightness value (
The fourth luminance value (C) is the average value (S) minus 0 to 0.2 times the difference (ΔS) from the average value (Smin). 11. The radiation image processing method according to claim 10, wherein the value is a value obtained by adding times. 12) The emphasis coefficient β is set to the first minimum value (βm) between the first brightness value (A) and the average value (E) of pixel brightness.
in1) to the maximum value (βmax), and between this average value (E) and the second luminance value (D), from the maximum value (βmax) to the second minimum value (βmin2
) The radiation image processing method according to any one of claims 7 to 9, characterized in that the radiation image processing method monotonically decreases to ). 13) Scanning the stimulable phosphor with excitation light to stimulate the radiation image stored and recorded on it, detect this emission with a photodetector, convert it into an electrical signal, and then calculate this electrical signal. In a signal processing device in a radiographic image recording and reproducing system that reproduces a visible image based on the processed signal, the arithmetic device converts the detected original image signal into Sorg. When the unsharp mask signal corresponding to the spatial frequency is Sus, the emphasis coefficient is β, and the reproduced image signal is S', the following calculation is performed. Alternatively, a radiation image processing apparatus characterized in that, when the magnitude of the unsharp mask signal Sus is within a predetermined range, an emphasis coefficient variable means is provided that decreases the emphasis coefficient β as the magnitude increases.