JPS6262376B2 - - Google Patents

Description

[Detailed description of the invention]

The present invention relates to an image processing method and apparatus in a radiographic system used for medical diagnosis, and more particularly, the present invention relates to an image processing method and apparatus for a radiographic system used for medical diagnosis, and more specifically, it uses a stimulable phosphor material (hereinafter simply referred to as a "phosphor") as an intermediate medium to process a radiographic image. The present invention relates to an image processing method in a radiographic system for recording a radiation image, reading out and reproducing the radiation image, and recording it as a final image on a recording material, and an apparatus for implementing the method. Such a radiographic system, as previously proposed by the present applicant in Japanese Patent Application No. 53-84741, uses a phosphor to absorb the radiation that has passed through the subject, and then converts this phosphor into a type of There is a method in which the phosphor is excited by the energy of the phosphor to emit the accumulated radiation energy as fluorescence, and this fluorescence is detected and imaged. A radiographic system using this phosphor has extremely high utility value in that it can record images over a wide radiation exposure range compared to a conventional radiographic system using silver halide photography. It has high utility value, especially as an X-ray photography system for the human body. Since X-rays are harmful to the human body if the exposure dose is high, it is desirable to obtain as much information as possible with a single X-ray photograph. 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 have come to be of a form that satisfies both of these to a certain 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 line photographic film for observation and interpretation, there is a problem in that the image quality cannot necessarily be said to be 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 on a recording material. 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. The present inventors conducted research on the frequencies to be emphasized and the diagnostic performance of the obtained X-ray images, and found that the frequencies important for diagnosis differ slightly depending on each part of the human body, but conventional From the sense of 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. It was also found that noise occupies a high proportion in the high frequency range, and that by reducing the emphasis in this high frequency range, the noise becomes less noticeable and becomes easier to see. Therefore, the present inventor proposed a radiation image processing method that can improve diagnostic performance by emphasizing ultra-low frequency components that are effective for diagnosis and increasing contrast. (Patent Application No. 53-163571) This emphasizes the ultra-low frequency components and at the same time
This is a radiation image processing method that relatively reduces high frequency components in which noise accounts for a large proportion, thereby making it possible to obtain visually easy-to-read images. In this method, a phosphor is scanned with excitation light, the radiographic image information recorded on it is read out, converted into an electrical signal, and then reproduced on a recording material. The unsharp mask signal Sus corresponding to the low frequency is determined, the original image signal read from the phosphor is Sorg, the emphasis coefficient is β,
A radiation image processing method characterized in that, when the reproduced image signal is S', the signal is converted by the calculation of S' = Sorg + β (Sorg - Sus), and frequency components above the ultra-low frequency are emphasized. It is. Here, the unsharp mask signal Sus corresponding to the ultra-low frequency is an unsharp image obtained by blurring the original image so that it only includes frequency components lower than the ultra-low frequency components (hereinafter referred to as "unsharp mask").
refers to the signal corresponding to the density of each scanning point. This unsharp mask has a modulation transfer function of 0.5 or more at a spatial frequency of 0.01 cycles/mm, and 0.5
The one that has a spatial frequency of 0.5 or less at a spatial frequency of cycles/mm is used. Here, the ultra-low spatial frequency means a spatial frequency of approximately 0.5 cycles/mm or less. A visible image on a recording material (hereinafter referred to as a "photographic image") created based on the signal emphasized by the above calculation formula.
) is the maximum value of the modulation transfer function of the system that gives
It is desirable that the value be 1.5 to 10 times the value of the modulation transfer function near zero frequency. In addition, since there is a lot of noise in the high frequency range and it is difficult to see, the S′ is characterized by performing smoothing processing such that the modulation transfer function is 0.5 or less in the frequency range of 0.5 to 5 cycles/mm. . 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 along with peripheral data according to the size of the unsharp mask, and calculate the average value (simple average or average value using various weighted averages).
This is a method to find Sus. 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 created by emitting light on the transmitted side. 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 the light. In the present invention, a phosphor refers to a phosphor that is irradiated with the first light or high-energy radiation and then stimulated (excited) optically, thermally, mechanically, chemically, or electrically. A phosphor that exhibits so-called photostimulability, which re-emits light corresponding to the amount of radiation irradiated. Light here includes visible light, ultraviolet light, and infrared light among electromagnetic radiation, and high-energy radiation refers to
rays, gamma rays, beta rays, alpha rays, neutron rays, etc. Excitation is from 500 to 800 nm, preferably from 600 to
It is preferable to use light in the wavelength range of 700 nm, 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 an excitation light source that has a peak in the above wavelength range, 500～
This can be achieved by using it in combination with a filter that cuts out light outside the 800 nm wavelength range. Excitation light sources that can emit light in the above wavelength range include Kr laser (647 nm), various light emitting diodes, He-Ne laser (633 nm), and Rhodamine B.
There are dye lasers, etc. Furthermore, since the wavelength range of tungsten yoso lamps extends from near ultraviolet and visible to infrared, they 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 10 ⁴ :
Since it is normal that the ratio is about 1 to ¹⁰⁶ :1,
When excitation light enters the photodetector, the S/N ratio is extremely reduced. By setting the light emission at the short wavelength side and the excitation light at the long wavelength side and separating the two as much as possible so that the excitation light does not enter the photodetector, the above-mentioned decrease in the S/N ratio can be prevented. For this purpose, it is desirable to use a phosphor whose emitted light is in the wavelength range of 300 to 500 nm. 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, described in Japanese Patent Application No. 1984-84742 (Ba _1-xy , Mg _x , Ca _y ) FX: aEu ²⁺ (X is Cl
and Br, x and y are 0<x+y≦0.6 and xy≠0, and a
is 10 ^-6 ≦ a ≦ 5 × 10 ^-2 ), -Patent Application 1973-
It is described in the specification of No. 84744 (Ba _1-x , M〓
_x ) FX: yA (However, M〓 is Mg, Ca, Sr, Zn and
at least one of Cd, X is at least one of Cl, Br, and A is Eu, Tb, Ce,
At least one of Tm, Dy, Pr, Ho, Nd, Yb and Er, x is 0≦x≦0.6, y is 0≦y≦
ZnS: Cu, Pb, BaO x Al ₂ O ₃ : Eu described in Japanese Patent Application No. 53-84740)
(However, 0.8≦x≦10) and M〓O・xSiO ₂ :A
(However, M〓 is Mg, Ca, Sr, Zn, Cd or Ba, and A is Ce, Tb, Eu, Tm, Pb, Tl, Bi or
Mn and x is 0.5≦x≦2.5); and LnOX described in Japanese Patent Application No. 53-84743:
xA (However, Ln is at least one of La, Y, Gd and Lu, X is at least one of Cl and Br, A is at least one of Ce and Tb, x is 0<x<0.1 ); and so on. 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. Furthermore, if the phosphor layer of a stimulable phosphor plate made using this stimulable phosphor is colored with a pigment or dye, the sharpness of the final image will be improved and favorable results will be obtained. can get. (Special application 1971-71604
No. 53-163571 states that when performing calculations for non-sharp mask processing, by reducing β in low brightness areas and increasing β in high brightness areas, it is possible to eliminate false images that are likely to occur due to frequency emphasis. It has been disclosed that this can be prevented. As an example, if an original photograph of the stomach (mergen) using barium contrast agent is subjected to the frequency processing with the enhancement coefficient β fixed,
The border of a wide uniform low-intensity area containing a large amount of contrast agent 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 has entered, and increasing β in a high-brightness area such as a gastric subdivision. False images inevitably have a very large impact on diagnosis. For example, if a false image occurs in a certain area, one 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 small in the low-luminance region of the spine and heart region, and β is made large 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 Application No. 53-163571, the method of changing the emphasis coefficient β is limited to a monotonous increase (β'≧0) that increases as the luminance value increases. As a result of further research, the present inventor found that when the emphasis coefficient β is positive or 0 (β≧0), it is better to include cases where the differential with respect to luminance is negative (β′<0), which makes the diagnosis more effective. We found that performance improved. In other words, it has been discovered that it is preferable to change the emphasis coefficient to 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 according to the brightness so that the low-brightness part is emphasized more than the high-brightness part, an image that is easy to see without the high-brightness part 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 way to change the emphasis coefficient β is to emphasize low brightness, where β decreases as the brightness increases (Figure 1C).
Then, it is possible to emphasize medium brightness (Fig. 1D) by increasing β at low brightness areas and decreasing β at high brightness areas. In low-luminance enhancement, β is kept constant at a large value in low-luminance parts, β is kept constant at a small value in high-luminance parts, and β is continuously decreased during this period2.
Two types are possible: a stepped type (FIG. 1C, a) and a curved type, in which β is decreased so that the absolute value of negative β' continuously increases (FIG. 1C, b). Hereinafter, the present invention will be explained in detail with reference to the drawings. Figures 1A to 1D are diagrams showing various ways of changing β in response to changes in luminance; the flat type (β = constant) in Figure 1A and the monotonically increasing type (β'≧0) in Figure 1B. )
is included in the method of Japanese Patent Application No. 53-163571. 1st C
The low brightness enhancement shown in the figure and the medium brightness emphasis shown in FIG. 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 improves diagnostic performance. will be significantly improved. Medium brightness enhancement in Figure 1D is suitable when low brightness areas 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 hinder diagnosis, so excluding these areas, only medium-brightness areas that are the target of diagnosis will be used. 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. The symbols A, B, C, and D in the table represent the emphasis of the types in Figures 1A, 1B, 1C, and 1D, respectively.

【table】

[Table] From Table 1, emphasis type C according to the method of the present invention,
Type D was rated higher than Type A in all cases;
It can be seen that in most cases, the evaluation is higher than type B (monotonic increase). 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. As shown in FIG. 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 of excitation light in the direction of arrow B with a scanning mirror 3a. The light that is stimulated by the scanning of the excitation light is collected by a light condenser 4a made of a light-guiding sheet material.
The light is detected by a photodetector 4 such as a photodetector placed at the output end of the sensor and converted into an electrical signal. This electrical signal is amplified by an amplifier 5 and then sent to an A/D converter 6.
The signal is converted into a digital signal and stored on the magnetic tape 7. The digital signals of each section stored on the magnetic tape 7 are read out to an arithmetic unit 8, such as a minicomputer, and after determining Sus, the above-mentioned calculation of S'=Sorg+β(Sorg-Sus) is performed. The above Sus has a modulation transfer function of 0.01 cycles/mm
A value must be specified that is greater than or equal to 0.5 when the spatial frequency is , and less than or equal to 0.5 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 depending on the part of the human body or each case, and these values may be stored in the memory of the computing device. Smoothing processing is performed on the S' to reduce high frequency components higher than very low spatial frequencies. This smoothing process makes it possible to reduce noise without damaging information necessary for diagnosis. This smoothing process will be explained in more detail with reference to FIG. Figure 3a shows the accumulated image on the phosphor at 10 pixels/mm.
This shows the frequency response when sampled at . This curve shows that when a rectangular aperture is used as the photodetector aperture,
It is known that when a Gaussian distribution aperture is used for a sinc curve, the curve becomes a Gaussian distribution curve. Figure 3b shows that the modulation transfer function is 0.01 cycles/mm.
This shows a rectangular unsharp mask () that is 0.5 or more when the spatial frequency is , and 0.5 or less when the spatial frequency is 0.5 cycles/mm, and a Gaussian distribution unsharp mask (). . In this example (), when the image on the phosphor is sampled at 10 pixels/mm, a simple average of approximately 63 pixels x 63 pixels (this is called "Non-sharp mask size N = 63") is taken. This is the case when a non-sharp mask is created. This means that the pixel on the phosphor is 6.3mm.
This is equivalent to scanning with a large-sized light beam of ×6.3 mm. Note that here fc is the modulation transfer function of 0.5
It shows the value of an arbitrary frequency included in the ultra-low frequency region of 0.5 to 0.01 cycles/mm when . On the other hand, with the Gaussian unsharp mask (), when averaging pixels, basically the rectangular unsharp mask ()
is the same as The non-sharp masks () and () 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. 3c is a graph showing the modulation transfer function after calculating (Sorg-Sus). The solid line () in FIG. 3d 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 photographic image signal is approximately 4.6 times the modulation transfer function (A) near zero frequency. The dotted line () in FIG. 3d shows the modulation transfer function when S' in FIG. 3d is subjected to smoothing processing of 5 pixels by 5 pixels. In such non-sharp mask processing, the diagnostic performance can be improved as described above by changing the emphasis coefficient β according to the brightness of the X-ray image. Examples of how to change β (Figure 1C, Figure 1D)
The details are shown in FIGS. 4 and 5. Figure 4 corresponds to the two-stage type (curve a) of low luminance emphasis (Figure 1C), where β is the maximum value β between luminance A and B.
It decreases from max to the minimum value β min. In other words, the enhancement coefficient is increased (β max) in the low luminance region (from S min to A), and is decreased (β max) in the high luminance region (from B to S max).
min). Brightness A is the minimum brightness (S min)
The maximum brightness (S max) and the minimum brightness (S
A good size is [Smin+(0.2-0.5)ΔS], which is the sum of 0.2 to 0.5 times the difference (△S) from
1) △D] is good. Figure 5 shows the case of medium brightness emphasis (Figure 1D), and in the case of the two-stage type indicated by the solid line a, β is between the first minimum value (β min1) and the maximum value between brightness A and B. (β max) and decreases between C and D from the maximum value (β max) to a second minimum value (β min2). That is, in the low brightness area (from S min to A) and the high brightness area (from D to S max), the emphasis coefficient is set small (β min 1, β
min 2), and is increased (β max) in the medium brightness region (from B to C). Here the first
(β min1) and the second minimum value (β
min2) may be equal to In the case of the mountain shape indicated by the dashed-dotted line b, β increases between brightness A and E, and between E and D
is decreasing between Brightness A is the minimum brightness (S
min), maximum brightness (S max) and minimum brightness (S
The brightness B is the sum of 0 to 0.2 times the difference (△S) from the average brightness (=S min + S max/2 or statistical average value) to The magnitude obtained by subtracting 0 to 0.2 times the difference (△S) [-(0 to 0.2) △S], the brightness E is the average brightness (), and the brightness C is the average brightness minus the difference (△S) from 0 to 0.
The brightness D is the difference (ΔS) from the maximum brightness (S max).
The size obtained by subtracting 0 to 0.2 times [S max - (0 to
0.2) △S] are respectively desirable. Here, the emphasis coefficient β can be arbitrarily selected within 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. Multiply the difference between Sorg and Sus (Sorg − Sus) by the gain (β) set by the value of Sorg,
By adding to Sorg, the desired S′=Sorg+
β(Sorg−Sus) is obtained. At this time, β
The values of max and β min are stored in the memory of the arithmetic unit 8 or inputted from the outside to be applied 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 monotonically decreasing region (A-B 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 in the arithmetic unit 8 or input from the outside. In addition, in the above calculation, the maximum brightness (S
max) and minimum brightness (S min) both correspond to the maximum and minimum in the target substantial image, and there may be brightness higher or lower than these in parts other than the image. sell. Note that, depending on the case, the maximum and minimum of the entire screen 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 is approximately the same when β is changed using the original image signal of the image on the phosphor and when β is changed using a non-sharp mask signal. It was the same. (In this specification, both cases are represented by the brightness of the original image.) Gradation processing may be performed simultaneously with the frequency enhancement described above. Very low frequency processing has relatively little effect on diseases where brightness changes slowly over a large area, such as lung cancer or 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.
Further, digital processing is performed after ultra-low frequency processing, or analog processing is performed after D/A conversion. These frequency-emphasized and, if necessary, gradation-processed data are recorded on a magnetic tape 7 (see FIG. 2). The data on this magnetic tape 7 is read out sequentially and converted into an analog signal by a D/A converter 9.
After being amplified by an amplifier 10, the signal is input to a recording light source 11. Light generated from this 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. A radiographic image is reproduced on this photographic film, and a diagnosis is made by observing this image. When reproducing and recording images on photographic film, a reduced photographic image can be obtained by recording at a higher sampling frequency than during input scanning. For example, 10 pixels/mm for the input system and 20 pixels/mm for the output system.
If you scan at pixels/mm, you will get a photographic image reduced in size to 1/2. As will be described later, a photographic image reduced to 1/2 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 setting the phosphor on a rotating drum, by plane two-dimensional scanning, or by electronic scanning such as a flying spot scanner. Further, the calculation of the unsharp mask 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, image data may be temporarily stored on a magnetic tape to create an integral histogram, and the position of the bending point of 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 luminance position where the bending point is desired may be inputted using a tractopen or the like. In the above embodiment, the final image is recorded on a photographic film, but diazo film, electrophotographic material, etc. can also be used as the recording material. Furthermore, instead of recording on a recording material, it may be displayed on a CRT for observation. Furthermore, this may be optically recorded on the recording material. Furthermore, the present invention provides an X-ray photographic system using a photographic film with an average gamma of 0.3 to 1.5 capable of recording radiation energy as an original recording medium (the present applicant previously filed a patent application
This was proposed in No. 53-28533. ) can also be applied.

[Brief explanation of the drawing]

Figures 1A, 1B, 1C, and 1D are graphs showing various examples of changing the emphasis coefficient β in non-sharp mask processing according to image brightness, and Figure 2 is a conceptual diagram showing the system of a system implementing the method of the present invention. Figure, 3rd
The figure is a graph showing the modulation transfer function at each stage of signal processing in the unsharp mask processing used in the present invention.
FIGS. 4 and 5 are graphs showing the relationship between emphasis coefficient and luminance, showing a method of changing the emphasis coefficient β in the present invention, and FIG. 6 is a block diagram showing an example of an arithmetic unit used in the present invention. 1... Fluorescent sheet, 3... Laser light source, 4...
...Photodetector, 11... Recording light source, 13... Recording material.

Claims (1)

[Claims] 1. When scanning a stimulable phosphor material to read out radiation image information recorded therein and converting it into an electrical signal, and then reproducing it as a visible image on the recording material, The unsharp mask signal Sus corresponding to the ultra-low spatial frequency at each scanning point is determined, the original image signal read out from the phosphor is Sorg, the emphasis coefficient is β, and the output signal used for reproduction is S′. In the method of emphasizing the frequency components above the ultra-low spatial frequency by performing the calculation S' = Sorg + β (Sorg - Sus), the emphasis coefficient β is set to (β≧0)
A radiation image processing method characterized in that the differential (β') with respect to the luminance of the original image or the unsharp mask is changed according to the magnitude of the signal so that it becomes negative (including the part where β'<0). 2. The patent claim is characterized in that a non-sharp mask is used as the sharp mask, the modulation transfer function being 0.5 or more at a spatial frequency of 0.01 cycles/mm and 0.5 or less at a spatial frequency of 0.5 cycles/mm. The radiation image processing method according to range 1. 3. The enhancement coefficient β has a maximum value in a low brightness region below a first brightness value (A) that is larger than the minimum value (S min) of brightness of the original image or non-sharp mask. (β max), and set the minimum value (β min) in a high luminance region equal to or higher than the second luminance value (B) which is greater than the first luminance value (A) but not greater than the maximum luminance value (S max). and the first brightness value (A) and the second brightness value (A).
3. The radiation image processing method according to claim 1, wherein the radiation image processing method monotonically decreases between the brightness values (B). 4 The first brightness value (A) is the minimum brightness value (S
min) is 0.2 to 0.5 of the difference (ΔS) between the maximum brightness value (S max) and the minimum brightness value (S min).
4. The radiation image processing method according to claim 3, wherein the value is a value obtained by adding times. 5 The second brightness value (B) is the minimum brightness value (S
min) is 0.7 to 1 of the difference (ΔS) between the maximum brightness value (S max) and the minimum brightness value (S min).
5. The radiation image processing method according to claim 3, wherein the value is a value obtained by adding times. 6 The enhancement coefficient β has a first minimum value (β min1) in a low brightness region below a first brightness value (A) that is equal to or greater than the minimum value (S min) of brightness of the original image or unsharp mask. A second minimum value (β min2) is taken in a high brightness region equal to or greater than a second brightness value (D) that is equal to or smaller than the maximum brightness value (S max), and the first brightness value (A ) and the second
Claims 1 or 2, characterized in that between the brightness values (D), the brightness increases from the low brightness side to the high brightness side until it reaches a maximum value (β max), and then decreases. The radiographic image processing method described. 7. The radiation image processing method according to claim 6, wherein the first minimum value (β min1) and the second minimum value (β min2) are equal. 8 The first brightness value (A) is the minimum brightness value (S
min) is 0 to 0.2 of the difference (ΔS) between the maximum brightness value (S max) and the minimum brightness value (S min).
A patent characterized in that the second brightness value (D) is a value obtained by subtracting 0 to 0.2 times the difference (ΔS) from the maximum brightness value (S max). A radiation image processing method according to claim 6 or 7. 9 The emphasis coefficient β is the first minimum value (β min1) between the third brightness value (B) larger than the first brightness value (A) and the first brightness value (A). to the maximum value (β max), and between this third brightness value (B) and a fourth brightness value (C) smaller than the second brightness value (D), the maximum value (β β max),
The fourth brightness value (C) and the second brightness value (D) are characterized by a monotonous decrease from the maximum value (β max) to the second minimum value (β min2). A radiation image processing method according to any one of claims 6 to 8. 10 The third brightness value (B) is 0 to 0.2 of the difference (ΔS) between the maximum brightness value (S max) and the minimum brightness value (S min) from the average value () of the image brightness.
The fourth brightness value (C) is a value obtained by adding 0 to 0.2 times the difference (ΔS) to the average value (). The radiation image processing method according to item 9. 11 The enhancement coefficient β monotonically increases from the first minimum value (β min1) to the maximum value (β max) between the first brightness value (A) and the average value (E) of image brightness. A patent claim characterized in that the difference between the average value (E) and the second brightness value (D) decreases monotonically from the maximum value (β max) to the second minimum value (β min2). The radiation image processing method according to any one of the ranges 6 to 8. 12 The stimulable phosphor is scanned with excitation light to cause the radiation image stored and recorded therein to be stimulated to emit light, and this emitted light is detected by a photodetector and converted into an electrical signal, and then this electrical signal is sent to a calculation device. In a signal processing device in a radiographic image recording and reproducing system that records a visible image on a recording material based on the processed signal, the arithmetic device converts the detected original image signal into Sorg and at each detection point. The unsharp mask signal corresponding to the very low spatial frequency of is Sus, the emphasis coefficient is β,
When the reproduced image signal is S', the following calculation is performed: S' = Sorg + β (Sorg - Sus), and furthermore, when the magnitude of the original photographic signal Sorg or the unsharp mask signal Sus is within a predetermined range, A radiation image processing apparatus characterized by comprising an emphasis coefficient variable means that decreases the emphasis coefficient β as the size increases.