JPH01131973A - Method and device for processing radiant image - Google Patents

Abstract

PURPOSE:To obtain a radiant image without having a pseudo image in which diagnostic performance is improved economically and at high speed by finding an unclear mask signal Sus corresponding to an ultra low spatial frequency at each scanning point and performing a specific arithmetic operation assuming an original image signal read out from a phosphor as Sorg and a reproducing image signal as S'. CONSTITUTION:When an accumulative phosphor material is scanned, and radiant image information recorded on the phosphor material is read out, then, converted to an electrical signal and after that, it is reproduced as a visible image, the unclear mask signal Sus corresponding to the ultra low spatial frequency at each scanning point is found and assuming the original image signal read out from the phosphor as Sorg and the reproducing image signal as S', an arithmetic operation shown in equation I is performed, and a frequency component over the ultra low spatial frequency is enhanced. In such a way, it is possible to prevent the pseudo image from being generated at every kind of case and part, which improves the diagnostic performance.

Description

DETAILED DESCRIPTION OF THE INVENTION The present invention relates to an improvement in an image processing method and apparatus in a radiation image information recording and reproducing system used for medical diagnosis, and more specifically, the present invention relates to an improvement in an image processing method and apparatus in a radiation image information recording and reproducing system used for medical diagnosis. Radiation image information recording in which radiation image information is recorded on a phosphor (hereinafter referred to as a "fluorescent material"), and then this radiation image information is read out and reproduced, and this is recorded as a final image on a recording material. This invention relates to improvements to image processing methods and devices in playback systems.

Radiation that has passed through the subject is absorbed by a phosphor to record radiation image information, which is then scanned with a laser beam or the like to excite it and the emitted light is read by a photodetector, and the read radiation image information is recorded. A radiation image information recording and reproducing system is known in which a radiation image is recorded on a recording material such as a photographic film by modulating a light beam with a light beam (see US Patent Section 3.
(No. 859,527) The radiation image information recording and reproducing system using this phosphor is extremely superior in that it can record images over a wide radiation exposure range compared to the conventional radiographic system using silver halide photography. It has high utility value, especially as an X-ray photography system for the human body.

On the other hand, X-rays are harmful to the human body when exposed to large amounts of radiation, so it goes without saying that it is desirable to obtain as much information as possible in just one X-ray photograph, but current X-ray film is not suitable for photographing. It is necessary to have both aptitude for observation and image interpretation, and since the film is designed to satisfy each of these to a certain degree, it cannot be said that the X-ray exposure range is sufficiently wide in terms of suitability for photography, and current X-ray photographic film Regarding the suitability for observation and interpretation of images, there was a problem in that the image quality was not necessarily sufficient for diagnosis.

Furthermore, although the radiation image information recording and reproducing system using a phosphor disclosed in the aforementioned U.S. Pat. It didn't solve the problem.

In order to solve the above-mentioned problems, the present inventors have proposed a radiation image processing method and apparatus that can simultaneously satisfy both suitability for imaging and suitability for observation and interpretation in a radiographic image information recording and reproducing system using a phosphor. (Unexamined Japanese Patent Publication No. 55-887
No. 40) By using this method and device, radiographic images with improved diagnostic performance can be obtained economically and at high speed.

In this method and apparatus, the original image signal read out from the phosphor is SOrg, the unsharp mask signal corresponding to the ultra-low spatial frequency at each scanning point is S org, and the emphasis coefficient is β
, when the reproduced image signal is S', s'-3org+β
Perform the operation expressed as (SOrg −3us),
This method is characterized by emphasizing frequency components higher than the ultra-low spatial frequency. In this method and device, the emphasis coefficient β may be fixed or variable, and if variable, Sorg
.

It may be changed depending on any magnitude of Sus.

Here, the ultra-low spatial frequency means a spatial frequency of approximately 0.5 cycles/m or less.

However, according to the inventor's subsequent research, when β is fixed, false images (artH'ac
It was found that t) was likely to occur. On the other hand, when β is made variable, for example, when β is monotonically increased (β′
≧0), it is possible to prevent the generation of false images in small areas (low brightness areas) of S org or Sus, but for example, it is possible to prevent the generation of black linear false images on the muscle side at the boundary between bone and muscle. It was difficult to prevent. In other words, in conventional methods, when photographing bones and muscles, the low-brightness side of the boundary between edges is emphasized and falls below the fog density of the final recording medium, resulting in white spots, or conversely, the density of the high-brightness side becomes too high, resulting in black lines. It is difficult to completely prevent the occurrence of false images, such as false images where the barium-filled area has a double contour shape in gastric double contrast imaging, and it is difficult to completely prevent the occurrence of false images. It was difficult to improve the condition sufficiently, and in some cases, it could even lead to misdiagnosis.

SUMMARY OF THE INVENTION In order to solve the above-mentioned problems, it is an object of the present invention to provide a radiation image processing method and apparatus that prevent the generation of false images.

That is, an object of the present invention is to provide a radiation image processing method and apparatus that can economically and rapidly obtain radiation images with improved diagnostic performance and without false images.

In order to achieve the above object, the present inventor has conducted intensive research and found that the above false image is a difference signal I Sorg -3usl
Based on this knowledge, when processing image signals, S' -3or -Sus, F (X) is lX1
When l<lXz l, F'(Xl)≧F'(Xz)≧0, and at least
A certain value Xo (IXs l<lXo l<IX
-F' (Xl) >F' (Xz) with the border ZI) as a monotonically increasing function), and calculate the degree of frequency emphasis where the difference signal l Sorg - Sus l is large. The above objective is achieved by reducing the increase in .

The radiation image processing method and apparatus of the present invention scans a phosphor with excitation light, reads radiation image information recorded therein, converts it into an electrical signal, and then reproduces it as a visible image. At each scanning point, an unsharp mask signal Sus corresponding to the ultra-low spatial frequency is obtained, and the original image signal read out from the phosphor is Sorg, and the reproduced image signal S
', then the above calculation formula (1) % formula % () (However, X-Sorg-X-Sorg (X) is lXl
When l<lXz I, F'(XI)≧F'(Xz)≧0, and at least
A certain value Xo (lX1 l<IX(,1<1X21
), the frequency component above the ultra-low spatial frequency is emphasized by performing an operation expressed by a monotonically increasing function (F'(XI)>F'(Xl)) with F'(XI)>F'(Xl) as the boundary. .

Here, the 1Xol is the difference signal IXI-lslXl-1
Needless to say, it is set within the range of sor.

In other words, the above F (X) is F' when X x 0
It includes a function in which F'(X)<0 when (X)>OlX>O, and a function in which part or all of this curved function is approximated by - or two or more linear functions.

Moreover, this F (X) is at least X −S org
A function of −3 us is sufficient, and it may be a function of S org and/or Sus at the same time. Mathematically, this means that F (X) in the above equation (1) is β(S
org).f(X) or β(SuS).f(X). In such a case, the above F' (X), F'
(X) is 9F (X)/EXS3zF (
Needless to say, it refers to X)/aX2.

The monotonically increasing function F (X) (X-
3org-3us), for example, F (X)-a
・sgn (X) ・IXI' +b (21(
However, α and b are constants and α>0. O<n<1sgn
X=1 (X>0), sgn X −-1(X<
0).

sgn X=0 (X=0) ) F (X) −a ・sin (1) X)
(3) (However, α> 0, l p
<2) or F (X) = 1-e''(X>
A curved monotonically increasing function such as 0)F (X)=-1+e''(X<0) can be considered.

These formulas (31, (41 are both 1Xtl
When <lX21, F' (Xt) >F/ (X
l )>0, and the condition F'(Xr)≧F'(X
l) ≧O is satisfied. Also, both of these functions are F' (X) < 0 (X > 0), F' (X) > 0
(X x 0), and in the region where X is positive, F' (X) gradually becomes smaller (the slope decreases) as ing.

However, as mentioned above, as F (X) of the present invention, the second derivative F' (X) of X may be zero in a part, that is, there may be a straight line part, or it may be zero throughout. Even if there is, it is sufficient as long as the above conditions are satisfied. In this case, F' (X) is zero over the whole, and F' (Xl
) >F' (Xz), for example, each of the functions (2) that is curved as a whole.

+3), (4) may be approximated by a broken line that is a combination of multiple linear functions.

An example of such a case is (however, a>b
>d >0. c=a IXl 1-blXl' l,
a = b lX2 l+c d lXz l
-a lXr l+b (lXz I-'IX
A polylinear function such as t1)-dlXzl) can be considered.

Note that the F (X) exemplified above are all point symmetrical at the origin, but in the present invention F (X
) is not necessarily limited to those symmetrical about the origin.

According to the invention, F' (X), that is, Sorg and Su
Since the first derivative of the function at the point of s is made smaller as IXI becomes larger, the rate of increase in frequency emphasis at larger IXI can be suppressed, thereby increasing the degree of culm waste wavenumber emphasis where the difference signal is large. On the other hand, where the difference signal is large, the degree of enhancement can be suppressed to prevent the generation of false images.

That is, according to the present invention, frequency enhancement by non-sharp mask processing is normally performed in areas where the difference signal is small in a radiographic image, and in areas where the difference signal is large (for example, the boundary between bone and muscle, the boundary between soft tissue and gas area, and the stomach). (such as blood vessel shadows in boundary angiography between the Ba-filled area and its surroundings), the increase in frequency enhancement is suppressed and the generation of false images is prevented.

In particular, F (X) in the above equation (1) is β
(Sorg)・f (X) or β(Sus)・f
When replacing with (X), β(Sorg) or β
(S us) naturally changes depending on Sorg or Sus, so in addition to the above effect,
Needless to say, this also has the effect of making β variable as disclosed in No. 62373.

In the present invention, the unsharp mask signal Sus corresponding to ultra-low frequencies refers to an unsharp image (
This refers to the signal at each scanning point of the "unsharp mask" (hereinafter referred to as the "unsharp mask"). As this unsharp mask, the modulation transfer function is 0.0
0.5 or more at a spatial frequency of 1 cycle/mm,
and 0 when the spatial frequency is 0.5 cycles/Irun.
．． Those having a value of 5 or less are used. Furthermore, as a non-sharp mask, the modulation transfer function is 0.5 or more at a spatial frequency of 0.02 cycles/mm and 0.5 or less at a spatial frequency of 0.15 cycles/mm. Using a non-sharp mask significantly improves diagnostic performance and is preferred.

If the spatial frequency at which the modulation transfer function is 0.5 is f, then the non-sharp mask used in the present invention has fc of 0.
．． 01~0.5 cycles/mm, preferably 0.02~
This can be said to be within the range of 0.15 cycles/cycle.

In the present invention, the original signal refers to a signal that has been processed by means commonly used in the optical industry;
That is, it goes without saying that the signal includes a signal that has been subjected to nonlinear amplification such as logarithmic amplification for band compression and nonlinear correction.

In addition, as a method for creating a non-sharp mask, (1) the original image signal at each scanning point is stored, and depending on the size of the non-sharp mask, the original image signal is read out along with peripheral data, and the average value (simple average or various A method for determining Sus (average value by weighted average) (in this method,
Sometimes it is created as an analog signal, sometimes it is created as a digital signal after A/D conversion, and sometimes it is created as a digital signal after A/D conversion.
This also includes cases in which the analog signal is de-sharpened with a low-pass filter only in the main scanning direction before D conversion, and digital signal processing is performed in the sub-scanning direction. ) (2) After reading out the original image signal using a light beam with a small diameter, if there is still an accumulated image, read out each scanning point using a light beam with a large diameter that matches the size of the non-sharp mask. (3) A method that takes advantage of the fact that the readout light beam gradually expands in diameter due to scattering in the phosphor layer. A method of creating an original image signal S org using the light emission signal of It can be controlled by changing the size of the aperture that receives the light.).

Among these unsharp mask creation methods, (1
) is the most preferred method.

In order to carry out method (1), the following calculations are ideally required to obtain the unsharp mask signal Sus at each scanning point.

5us-Σ alj Sorg (1,j) Here, 1. j is the coordinate of a circular area centered on each scanning point (the number of pixels in the area is N in the diametrical direction), and a
ij is a weighting coefficient that preferably has an isotropic and smooth change in all directions. However, when such an operation is simply performed, the multiplication for each scanning point is approximately −W N 2 It is necessary to perform the addition N2 times, and if N is large, the calculation takes an extremely long time and is impractical.

In fact, when reading out a normal radiographic image by scanning a phosphor, it is necessary to avoid losing the frequency components of the image, so although there are some differences depending on the image, it is usually ~20 pixels/sampling rate (200 to 50 pixels in terms of pixel size)
Since the non-sharp mask in Rikiki's invention corresponds to extremely low frequencies, it is necessary to perform calculations using a very large number of pixels to create this mask.

For example, in the case of a mask with Gaussian distribution weighting coefficients, if the pixel size is 100μX100μ, fc-0,
For 1 cycle/mm, N will be approximately 50 and f
, -0,02 cycles/mm, N is approximately 2
50, the calculation time becomes enormous.

Furthermore, averaging a circular area means changing the addition range for each scanning line, but having to make such a judgment during calculation greatly complicates the calculation mechanism. It is uneconomical.

In order to solve this problem and perform image processing practically, the method of obtaining a non-sharp mask signal is to use two sides parallel to the main scanning direction and two sides parallel to the sub-scanning direction. A method for obtaining an unsharp mask signal Sus for ultra-low spatial frequencies at each scanning point by simply averaging the original image signals S org at each scanning point within a rectangular area surrounded by -62379), or in the main scanning direction, the analog signal is de-sharpened using a low-pass filter with a certain reduction in transparency, and in the sub-scanning direction, the A/D-converted digital signal is subjected to averaging processing. Method for determining unsharp mask signal Sus for very low spatial frequencies at scanning points (Japanese Patent Publication No. 62-62381 by the same applicant)
No.) is preferred.

In the former case, the weight is uniform in a rectangular area, and therefore, compared to, for example, a mask with Gaussian distribution weights whose weight decays smoothly, the transfer characteristics may oscillate or may be non-sharp depending on the direction. Despite the drawbacks such as different degrees, the present inventor has found that there is no substantial difference in terms of improvement in diagnostic performance from the case of the ideal mask operation described above. . Furthermore, since this method uses a rectangular unsharp mask and obtains the unsharp mask signal Sus by simple averaging of the signals within the mask, it is possible to obtain the unsharp mask signal Sus by an extremely simple method. This enables a significant reduction in calculation time and equipment cost. This is an advantage common to both digital and analog signal processing.

Note that the transfer characteristic of a rectangular unsharp mask with uniform weights is a 5ine function sinπX (sine (x) = −) π The above-mentioned provision that 0.01 to 0.5 cycles/mm, preferably 0.02 to 0.15 cycles/mm is in the range of 0.01 to 0.5 cycles/mm, theoretically in this case It is synonymous with setting the length to GOmm to L, 2mm5, preferably 30mm to 4mm.

Note that even when the shape of the non-sharp mask is rectangular, the length of each side only needs to be within the above range, and for example, a rectangular mask with a large aspect ratio is effective for image processing of linear tomography.

Furthermore, even in the latter method of using a low-pass filter, it is created using a low-pass filter with a spatially asymmetric transfer characteristic in the main scanning direction, and digital averaging is performed in the sub-scanning direction. Although it is based on calculations with heavy weights, there is no real difference in terms of improved diagnostic performance from the ideal mask calculation described above, and moreover, the main scanning direction is a low-pass filter. Therefore, it is possible to significantly reduce the addition operation for digital signals, which takes a long time to calculate.
It has been discovered that significant cost reductions in equipment can be achieved.

Furthermore, in the latter case, if the arithmetic average of the digital signals in the sub-scanning direction is a simple arithmetic average, there is no need to perform multiplication, which simplifies the device and speeds up the calculation, but even with such a method, The inventors have found that there is virtually no difference in diagnostic performance compared to the ideal case.

In the present invention, in addition to the above operations, smoothing processing can also be performed. In general, in the frequency range of ultra-low spatial frequencies or higher, there is a lot of noise and it is often difficult to see, so it is often preferable to perform further smoothing processing to further improve diagnostic performance. The smoothing process is preferably such that the modulation transfer function is 0.5 or more when the spatial frequency is 0.5 cycles/mm, and 0.5 or less when the spatial frequency is wig cycles/mm. What kind of smoothing processing is preferable? For example, when interpreting shadows with a relatively low frequency such as a chest tomographic image, it is preferable to remove as much noise as possible. When it is necessary to trace fine blood vessel shadows that contain high frequency components, excessively strong smoothing processing may be undesirable as it may obscure the shadows that you want to see. According to the research of the present inventor, by performing the smoothing process as described above,
was found to be effective in improving diagnostic performance for almost all X-ray images. Furthermore, this smoothing process is equally effective whether it is applied to S′ after the ultra-low spatial frequency processing of the present invention or to the original image signal S org. It is acknowledged that there is.

Further, in the present invention, gradation processing may be performed in addition to frequency emphasis processing using a non-sharp mask.

Ultra-low frequency processing has a relatively small effect on diseases where luminance changes slowly over a large area, such as lung cancer and breast cancer.
It is desirable to use the gradation processing disclosed in Japanese Patent Application Laid-open No. 55-88740, Japanese Patent Publication No. 62-53179, Japanese Patent Publication No. 63-26585, etc. In this case, gradation processing may be performed either before or after ultra-low frequency processing.

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. It refers to a so-called photostimulable phosphor that re-emits light corresponding to the amount of radiation irradiated, and those having a photostimulable emission wavelength of 300 to 500r+n+ are particularly preferable, such as rare earth element-activated alkaline earth metals. Fluorohalide phosphor [specifically, Special Publication No. 60-42
It is described in the specification of No. 837 (Ba 1-x-y
, Mg , , Ca y )FX:aEu2+(
However, X is at least one of CQ and Br, X and y are 0xy≦0.6 and xy≠0, and a is to-f1≦a≦5×lrow2 ) described in Japanese Patent Publication No. 59-44333 (Bal
x , MIl, ) FX : y A (where Mn is at least one of Mg, Ca, Cr, Zn and Cd, and X is at least one of CQ, Br and I
A is Eu, Tb, Ce, TIO.

At least one of Dy, Pr, Ho, Nd, Yb and Er, X is O≦X≦0.6, Y is O≦y≦0.
2), etc.]: ZnS:Cu, 'PbSBaO x AL described in Japanese Patent Publication No. 60-9542
03: Eu (however, 0.8≦X≦10) and Mn0.xSiO2:A (however, Mn is Mg, Ca.

Sr, Zn, Cd or Ba, A is Ce.

Tb, Eu, Tm, Pb, T9J, Bi or Mn, and X is 0.5≦X≦2.5); and Tokko Sho 5
LnOX:xA(
However, Ln is at least one of La, Y, Gd, and Lu, X is at least one of C9J and B, A is at least one of Cc and Tb, and X is 0.
<X <0.1)-, and the like. 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, which is a favorable result. (Japanese Patent Publication No. 59-23400).

In the present invention, a laser beam with good directivity is used as excitation light for reading out the radiation image accumulated on the stimulable phosphor plate. As the excitation light source of the laser beam, 500
~800nm s Preferably one that emits light in the range of 600-700nm, such as a He-Ne laser (633nm).
m), Kr laser (847 nm) is preferred, but excitation light sources other than those mentioned above can also be used if a filter that cuts light other than 500 to 800 nm is used in combination.

A radiographic image subjected to image processing according to the present invention is input to a CRT and reproduced as a visible image on the CRT, thereby enabling CRT diagnosis. Further, the radiation image may be displayed on a CRT and observed, and then optically recorded on a recording material such as a silver halide photographic film, a diazo film, or an electrophotographic material.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS The present invention will be described in detail below based on an X-ray image information recording and reproducing system that is an embodiment thereof.

FIG. 1 shows an example of an apparatus that performs image processing according to the present invention in the process of creating a reproduced image. When X-rays are emitted and irradiated onto a human body, the X-rays that pass through the human body enter the phosphor plate. This phosphor plate stores the energy of the x-ray image at the phosphor trap level. A stimulable phosphor sheet 1 on which a radiation image has been accumulated and recorded by this X-ray photography 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. Main scanning is performed by scanning a mirror 3a with laser light from a laser light source 3 of excitation light having a wavelength of 500 to 800 nI11.
This is done by scanning in the direction of arrow B. By scanning this excitation light, stimulated luminescence in the wavelength range of 300 to 500r++n is generated, and this stimulated luminescence is collected by a light condenser 4a made of a light-guiding sheet material.
It is detected by a photodetector 4 such as a photomultiplier placed at the output end of a and 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 sent to a calculation section 7. In the calculation unit 7, the calculation device 8a for obtaining the unsharp mask signal Sus
Then, in the difference signal calculation device 8b, Sorg
-8us is obtained, and then F (X)
After that, the above-mentioned calculation formula (1), S' -3or
This calculation is performed in the calculation device 8d which calculates g +F(X), and the digital signal S′ obtained after the calculation is
is converted into an analog signal by the D/A converter 9, and then sent to the amplifier [
After being amplified by 0, the signal is sent to the CRT 15 and generated as a visible image on the CRT, making CRT diagnosis possible. Furthermore, this image data> is inputted to the recording light source 11 as required. The light generated from the recording light source 11 is transmitted through the lens 12.
A recording material 13, for example a photographic film, mounted on a printing drum 14 is irradiated through the photoreceptor. In this case, a radiation image is reproduced on a photographic film, and diagnosis can be made by observing this image.

The image processing described above may be performed online by directly using the output of the photodetector 4 as in the embodiment described above, or may be performed offline based on data previously recorded on a magnetic tape or the like.

The non-sharp mask processing uses the non-sharp mask signal Sus and the original image signal Sorg obtained by the photodetector, S' = Sorg + F (X) (where F (X) is the definition of the above calculation formula (1) by)
This is performed by the operation expressed as .

This unsharp mask signal Sus is obtained by the method described later, and the modulation transfer function is 0.01 cycle/tn.
Use a modulation transfer function that is 0.5 or more at a spatial frequency of m and 0.5 or less at a spatial frequency of 0.5 cycles/mm, or preferably a modulation transfer function of 0.0.
0.5 or more at a spatial frequency of 2 cycles/mm,
And it must be specified to be 0.5 or less at a spatial frequency of 0.15 cycles/tnm. Furthermore, when calculating the above equation, the function F (X) must be specified. This function 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 functions may be stored in the memory of the computing device.

Hereinafter, an embodiment in which image processing is performed by specifically determining F (X) will be described in detail.

As mentioned above, F (X) can be changed in various ways, and an appropriate one can be selected from among them. However, without expressing F (X) in advance as a general formula, it is possible to F (X) may be determined by finding F (X) from the value. That is, for example, X and F (X
) may be created on a disk or memory, and the value of F(X) may be output according to the input value of X using this conversion table.

In the description of the embodiments below, F (X) is used as F (X
)−αJ8(6)
A table-looking method using an X4F(X) conversion table will be explained as a representative example.

Figure 2 shows a flowchart for calculating F (X) - α, /TTT. In this case, the original image signal (S org
) is determined (21), and based on this S org, the unsharp mask signal (S
(22). Next, this S org
and Sus to calculate XX-Sorg-5us to find X (23). If X is positive or 0, then
Let X be X and α be α24.25), and when X is negative, replace
g +F (X), that is, SOrg+α5 is calculated (27) to find S'.

In order to perform the calculations shown in the flowchart of FIG. 2, input/output devices 31 . Controller 32. An arithmetic unit 33 and a memory 34 are used, and this arithmetic unit 33 must have the functions of calculating the square root Cr) and performing addition, subtraction, multiplication, and division.

Figure 4 shows X-F(X) (for example, F (X) -ctff
) shows an example of using the conversion table of (41,
42) , XX Sorg - Sus is calculated to find X (43). The data is converted from the X obtained here by referring to the conversion table (X-F(X)), and F (X) corresponding to X is obtained (44). Next, using F (X) obtained in this way, the calculation of Sorg +F (X) is performed (45) to obtain S'.

In order to perform the calculations shown in the flowchart of FIG. 4, input/output devices 51. Controller 52. In addition to using the arithmetic unit 53 and memory 55, a table memory 54 is used.

In this case, the arithmetic unit 53 need only be capable of addition and subtraction since it is not necessary to calculate F(X)-αF ball].

In addition, the signal S which has been further frequency-emphasized as described above is
By applying smoothing processing to reduce high frequency components to ', it is possible to reduce noise without damaging the information necessary for diagnosis.

Furthermore, in addition to frequency enhancement processing using a non-sharp mask,
Gradation processing can also be used together. When gradation processing is performed before ultra-low frequency processing, A/D conversion is performed after gradation processing is performed using a nonlinear analog circuit. If it is performed after A/D conversion, digital processing can also be performed using a minicomputer. After ultra-low frequency processing, digital processing or D/A conversion converter analog processing is performed.

Furthermore, when reproducing and recording images on a CRT or photographic film, a reduced image can be obtained by recording at a higher sampling frequency than during input scanning. For example, in the input system, 10 pixels/
mm, and if the output system scans at 20 scans/mm, the image will be reduced to 1/2. Images reduced to 1/2 to 1/3 in this way are close to the frequency components considered necessary for diagnosis or the frequency range with the highest visibility, so the contrast appears to be higher visually and is very easy to see. Become.

Note that in the above embodiment, the original image signal So
rg includes cases where it means a signal after performing band compression such as logarithmic transformation and nonlinear correction.

Practically speaking, since the output of the photodetector is subjected to signal processing, it is desirable to perform band compression such as logarithmic transformation. It goes without saying that, in principle, it is also possible to use the output of the photodetector as it is as Sorg for subsequent processing.

In addition, theoretically, when calculating this mask, the average energy should be calculated, but according to the inventor's experiments, when calculating this unsharp mask signal, a value corresponding to the logarithmically compressed density is used. Even when we calculated the average value, the results did not change. This is practically advantageous in terms of processing.

Examples will be given below to make the effects of the present invention more clear.

Examples A total of 50 cases of typical sites listed in Table 1 were recorded directly on conventional radiographic film and read out from the phosphor according to the present invention, and are shown in Figure 6 as F (X). Curves A, B, D and polygonal line C were selected and compared with an image created by performing ultra-low frequency processing according to the above-mentioned formula (1) to examine the improvement in diagnostic performance for the main parts of the human body.

The solid line A in Fig. 6 is F (X) −0,4・sgn (
X)l X I 1''2, i.e. the above function (2)
α-0,4, n −1/2 , b −0, and is a continuous curve-type function. Gradient (F' (
X)) becomes smaller as IXI increases,
F' (X) is negative when X is positive, and positive when X is negative.

Broken line B is F (X)-1-e-”” (X>0), F
(X)=-1+e""(X<0), that is, the coefficient of X in the function (4) is set to 1.4, and this is also a continuous curve-type function. Gradient (F'(X))
becomes smaller as IXI increases, and F'
When (X) is positive, it is negative, and when X is negative, it is positive.

The dashed line C is F (X) -sgn (X) (n
l X l +c.

nst) indicates a linear function in which n of IXI decreases as IXI increases, and in the above function (5), a -1, b -0,75, c = 0.
This corresponds to 5. That is, it is a polylinear function expressed as 0.3≦IXI. In this function, the gradient F'
(X) decreases stepwise as IXI increases, and F' (X) is 0.

The dotted line indicates a conversion table that is convex upward in the region of X>0 and convex downward in the region of X×0. This means that X>0. This is an example where the shape is not symmetrical in the region of X<0. In this case, calculations were performed using the table-looking method.

Since it is virtually impossible to confirm whether or not there has been an improvement in diagnostic performance using physical evaluation values (e.g., sharpness, contrast granularity, etc.) of ordinary photography, Based on subjective evaluation by

The evaluation results are shown in Table 1.

Table 1 Case 1 Site Review Image Title 2 The skull is white and does not come off, and there are no black line-like false images of the facial muscles, making it easy to see and diagnosing muscle tumors.

Bone and muscle: No false images were generated in both the bone and muscle areas, and accurate diagnosis was made for both.

Angiography No two-color images are generated, making it possible to diagnose both thin and large contrast-enhanced blood vessels.

Double contrast angiography of the stomach: No images were generated in the margins and areas filled with a large amount of contrast agent, and the overall diagnosis was good.

Simple abdomen: The gas area of the intestines is not emphasized more than necessary, making it easier to diagnose the entire abdomen.

Note that curves A, B, and D in FIG. 6 are used as F (X).

Although some differences in the degree of improvement in diagnostic performance were observed for individual images depending on which of the polygonal lines C was selected, on average no substantial difference was observed for each case.

As is clear from Table 1, according to the present invention, various cases 7
The occurrence of false images was prevented, and diagnostic performance was improved.

As described above, the method of the present invention can greatly improve diagnostic performance in a radiation image processing method using a stimulable phosphor, and has a remarkable practical effect.

[Brief explanation of the drawing]

FIG. 1 is a schematic block diagram showing an example of a radiation image recording and reproducing system implementing the radiation image processing method of the present invention, FIG. 2 is a flowchart showing an example of the method of the invention, and FIG. 3 is shown in FIG. A block diagram showing an example of the configuration of a calculation unit used to implement the method, FIG. 4 is a flowchart showing the other side of the method of the present invention, and FIG. FIG. 6 is a block diagram showing an example of the configuration of an arithmetic unit that calculates the arithmetic expression S' −3org
It is a graph which shows six examples of the function F (X) of X). 1...Storage phosphor sheet 3...Laser light source 4.
...Photodetector 4a...Concentrator 5.10.
...Amplifier 7...Arithmetic section 15...CRT Fig. 4 Fig. 5 5゜ Fig. 6

Claims (11)

[Claims] (1) After scanning the stimulable phosphor material and reading out the radiation image information recorded on the phosphor material and converting it into an electrical signal, at each scanning point when reproducing it as a visible image. The unsharp mask signal Sus corresponding to the ultra-low spatial frequency of is determined, and the original image signal read out from the phosphor is
org, and when the reproduced image signal is S', the calculation formula S'=Sorg+F(X) (However, X=Sorg-Sus, F(X) is |X_1|
<|X_2|, then F'(X_1)≧F'(X_2)≧0, and at least a certain value of X
2 |) A radiographic image characterized by performing an operation expressed by a monotonically increasing function such that F'(X_1)>F'(X_2) as a boundary, and emphasizing frequency components above the ultra-low spatial frequency. Processing method. (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 0.5 or less at a spatial frequency of 0.5 cycles/mm. The radiation image processing method according to claim 1 or 2, characterized in that a sharp mask is used. (4) As a non-sharp mask, the modulation transfer function is 0.5 or more at a spatial frequency of 0.02 cycles/mm and 0.5 or less at a spatial frequency of 0.15 cycles/mm. The radiation image processing method according to claim 1 or 2, characterized in that a sharp mask is used. (5) The monotonically increasing function F(X) is a curved function that satisfies the following conditions: F″(X)<0(X>0) F″(X)>0(X<0) A radiation image processing method according to any one of claims 1 to 4. (6) The monotonically increasing function F(X) is mainly F″(X)
<0(X>0) F″(X)>0(X<0) It consists of a curved part that satisfies the condition, and partially F″
The radiation image processing method according to any one of claims 1 to 4, characterized in that the method includes a linear portion where (X)=0. (7) The monotonically increasing function F(X) approximates a curved function that satisfies the following conditions: F″(X)<0(X>0) F″(X)>0(X<0) A radiation image processing method according to any one of claims 1 to 4, characterized in that the method is comprised of a combination of linear functions. (8) The function F (X) is F (
6. The radiation image processing method according to claim 5, wherein the radiation image processing method is a curve-type function expressed as follows. (9) A patent claim characterized in that the function F(X) is a curved function expressed as F(X)=α・sin(pX) (where |pX|<π/2, α>0) The radiation image processing method according to item 5. (10) The function F(X) is a curved function expressed as F(X)=1-e^-x(X>0) F(X)=-1+e^x(X<0) A radiation image processing method according to claim 5, characterized in that: (11) An excitation light source that scans the stimulable phosphor to stimulate the radiation image stored and recorded therein, a photodetector that detects this emission and converts it into an electrical signal, and In a signal processing device in a radiographic image recording and reproducing system, which is equipped with a calculation device that processes electrical signals, the calculation device sends the detected original image signal to Sorg, and a non-sharp mask corresponding to the ultra-low spatial frequency at each detection point. S signal
When us, S'=Sorg+F(X) (However, X-Sorg-Sus, F(X) is |X_1|
<|X_2|, then F'(X_1)≧F'(X_2)≧0, and at least a certain value of X
A monotonically increasing function with F'(X_1) >F'(X_2) at
A radiation image processing device characterized in that it performs the calculation expressed by: