JPS6262373B2 - - 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 radiographic image, reading out and reproducing the radiographic image, and recording it as a final image on a recording material, and an apparatus for carrying out 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 are designed to satisfy both of these requirements to some degree. 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 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 on a recording material. The first object of the present invention is to provide an X-ray image processing method that improves the diagnostic performance of X-ray images. As an image processing method for images on an X-ray photographic film, Japanese Patent Application Laid-Open No. 48-25523 discloses a photographic film having a two-step gradient contrast characteristic having a relatively low contrast gradient portion and a relatively high contrast gradient portion. A technique has been disclosed in which recording is performed using unsharp masking, which emphasizes frequencies in high spatial frequency (hereinafter, "spatial frequency" will simply be referred to as "frequency" in the description of the present invention) region. This technology is an image processing method used to copy large-sized X-ray photographic film onto small-sized photographic film for convenient storage. It compresses the size of the X-ray image and provides the same diagnostic performance as the original photograph. This is to obtain a preserved reduced image. However, the above-mentioned method is carried out for the purpose of copying while preventing deterioration of the system response, and therefore the emphasized frequency is high and noise is likely to be increased. I can't hope for what I did. 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. An object of the present invention is to provide a radiation image processing method that can improve diagnostic performance by emphasizing ultra-low frequency components that are effective for diagnosis and increasing contrast, and an apparatus for implementing the method. That is. Further, the present invention provides a radiation image processing method and apparatus that enhances ultra-low frequency components and at the same time relatively reduces high frequency components in which noise accounts for a large proportion, making it possible to obtain visually easy-to-read images. The purpose is to The present invention scans a phosphor with excitation light, reads out the radiographic image information recorded on it, converts it into an electrical signal, and then reproduces it on a recording material at an ultra-high level at each scanning point. Find the unsharp mask signal Sus corresponding to the frequency, and when the original image signal read from the phosphor is Sorg, the emphasis coefficient is β, and the reproduced image signal is S′, S′ = Sorg + β (Sorg − Sus) This radiation image processing method is characterized in that the frequency components above the ultra-low frequency are emphasized by converting the signal by the following calculation. Here, the ultra-low spatial frequency means a spatial frequency of approximately 0.5 cycles/mm or less. The device of the present invention also includes an excitation light source for scanning a stimulable phosphor to stimulate the radiation image stored therein, and a photodetector for detecting the emitted light and converting it into an electrical signal. In a signal processing device in a radiographic image recording and reproducing system, which includes a detector and a calculation device that processes this electrical signal, the calculation device converts the detected original image signal into Sorg, which corresponds to an ultra-low spatial frequency at each detection point. This radiation image processing apparatus is characterized in that it performs the calculation Sorg+β(Sorg−Sus), where the unsharp mask signal is Sus and the emphasis coefficient is β. Note that the calculations in the above method and apparatus may be performed through any calculation process as long as the same result as this equation is obtained as a result.
Needless to say, the order of the expressions is not limited to this. 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. As this non-sharp mask, one is used whose 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. 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.
It is desirable that the value of the maximum modulation transfer function (referred to as ) is 1.5 to 10 times the value of the modulation transfer function near zero frequency. 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 small diameter light beam, etc., and then use a large diameter light beam that matches the size of the non-sharp mask to signal each scanning point if there are still accumulated images. 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 has a so-called photostimulability property, which, after being irradiated with the first high-energy radiation, re-emits light corresponding to the irradiation amount of the first high-energy radiation by optical stimulation (excitation). A phosphor that shows Here, high-energy radiation includes X-rays, gamma rays, beta rays, alpha rays, neutron rays, and the like.
Excitation is preferably performed with light in the wavelength range of 600 to 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 selecting an excitation light source that emits light in this wavelength range, or by selecting a pumping light source that emits light in this wavelength range, or by selecting an excitation light source that emits light in this wavelength range, or by selecting an excitation light source that emits light in this wavelength range. It can be obtained by using a combination of an excitation light source and a filter that cuts out light outside the wavelength range of 600 to 700 nm. Excitation light sources that can emit light in the above wavelength range include Kr lasers, various light emitting diodes,
Examples include He-Ne laser and rhodamine B-dye laser. In addition, the wavelength range of tungsten yoso lamps extends from near ultraviolet and visible to infrared, so
It can be used in combination with a filter that transmits light in the 700nm wavelength range. 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. Fluorescent materials that emit light in the wavelength range of 300 to 500 nm include LaOBr:Ce, Tb SrS:Ce, Sm SrS:Ce, Bi BaO・SiO ₂ :Ce BaO・6Al ₂ O ₃ :Eu (0.9Zn , 0.1Cd) S:Ag BaFBr:Eu BaFCl:Eu, etc. Hereinafter, the present invention will be explained in detail based on an X-ray photography system as an embodiment thereof. FIG. 1 shows the process of drawing an X-ray photograph. When X-rays are emitted and irradiated onto a human body, the X-rays that pass through the human body enter the phosphor plate. The phosphor plate stores the energy of the x-ray image at the phosphor trap level. After the X-ray image is taken, the phosphor plate is scanned with excitation light in the wavelength range of 600 to 700 nm to excite the accumulated energy from the traps, causing them to emit light in the wavelength range of 300 to 500 nm. This emitted light is measured with a photodetector, such as a photomultiplier tube or a photodiode, which receives only light in this wavelength range. After reading the X-ray image, the output signal of the photodetector is amplified and then converted to a digital signal by an A/D converter and stored on magnetic tape. The digital signals of each section stored on this magnetic tape are read out to a calculation device, 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 0.5 or more when the spatial frequency is , 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 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 frequency components higher than ultra-low spatial frequencies. This smoothing process ensures that the information necessary for diagnosis is not lost.
Noise can be reduced. This unsharp mask processing will be explained in more detail with reference to FIG. Figure 2 a 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 2b shows a rectangular unsharp mask () whose modulation transfer function is 0.5 or more at 0.01 cycles/mm and 0.5 or less at 0.5 cycles/mm, and a Gaussian distribution unsharp mask (). This shows that. 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 images on the phosphor 6.3
This is equivalent to scanning with a large-sized light beam of mm/6.3 mm. Note that here c is the modulation transfer function
0.5, the value of an arbitrary frequency included in the ultra-low frequency region of 0.5 to 0.01 cycles/mm is shown. 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. 2c is a graph showing the modulation transfer function after calculating (Sorg-Sus). The solid line () in FIG. 2d 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. 2d shows the modulation transfer function when S' in FIG. 2d is subjected to smoothing processing of 5 pixels by 5 pixels. FIG. 3 shows an embodiment in which the emphasis coefficient β is continuously changed according to the original image signal (Sorg) or the unsharp mask signal (Sus). By changing β in this way, it is possible to prevent false images that tend to occur due to frequency emphasis. As an example, when performing the frequency processing on an X-ray image of the stomach using a barium contrast agent while fixing the enhancement coefficient β, the boundary of a wide uniform low-intensity region containing a large amount of contrast agent is detected. 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 has entered, and increasing β in a high-brightness area such as a gastric subdivision. As another example, in the case of frontal chest photography, if β is fixed, noise will increase in low brightness areas of the spine and heart, and in extreme cases, details and white will be missed. (This is visually very noticeable and has a negative impact on diagnostic performance.) Similarly, if β is made smaller in the low-brightness region of the spine and heart region, and β is increased in the high-brightness region of the lung field, the above-mentioned noise can be reduced. It is possible to prevent an increase in white areas and white spots. In both of the above two examples, the emphasis coefficient β
If we fix it to a small value and perform frequency processing, we get
Although it is true that various false images do not occur, the contrast of blood vessels in the gastric subdivisions and lung fields, which make an important contribution to diagnostic performance, does not increase, and diagnostic performance does not improve. In this way, by continuously changing the enhancement coefficient β so that it monotonically increases as the brightness of the image on the phosphor increases, an image with improved diagnostic performance can be obtained while preventing the generation of false images. It will be done. In FIG. 3, the minimum brightness S ₀ and maximum brightness S ₁ are determined from the histogram of the image on the phosphor, and β is changed approximately linearly between them. This change in β may be performed by changing β using an arbitrary curve that increases monotonically (that is, β'≧0) as a basic tone.
S ₀ and S ₁ are determined depending on the type of X-ray image to be processed, and for example, the minimum and maximum brightness may be the brightness values when the integral histogram is 0 to 10% and 90 to 100%, respectively. In addition, in the experiments conducted by the present inventors,
β by the original image signal of the image on the phosphor
The effect was approximately the same when β was changed using a non-sharp mask signal and when β was changed using a non-sharp mask signal. Gradation processing may be performed simultaneously with the frequency enhancement described above. Ultra-low frequency processing has relatively little effect on diseases where the concentration 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. The data, which has undergone frequency enhancement and gradation processing if necessary, is recorded on a magnetic tape. The data on this magnetic tape is sequentially read out, converted into an analog signal by a D/A converter, amplified by an amplifier, and then input to a recording light source. Light generated from this recording light source passes through a lens and is irradiated onto a recording material, such as a photographic film.
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, if the input system scans at 10 pixels/mm and the output system scans at 20 pixels/mm, the photographic image will be reduced 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, 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. Furthermore, 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. In the above embodiment, the reproduced image is recorded on 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. For more than 100 cases, we compared images recorded directly on conventional X-ray photographic film with photographic images created by reading out from the phosphor and subjecting it to frequency processing using the method of the present invention. We investigated the improvement in diagnostic performance. At this time, copy photos are made with various frequencies to be emphasized and emphasis coefficient β changed,
We investigated the relationship between frequency and diagnostic performance. This improvement in diagnostic performance is based on the physical evaluation values of ordinary photography (e.g. sharpness, contrast,
It is difficult to substantiate this with graininess, etc.). Therefore, we asked four radiology interpretation experts (radiologists) to make observations, and statistically processed their subjective evaluations to evaluate diagnostic performance. The evaluation criteria are as follows. +2: Lesions that could not be seen with conventional X-ray film methods can now be seen,
Lesions that are extremely difficult to diagnose are now easier to see, and diagnostic performance has clearly improved. +1: Lesions that are difficult to diagnose using the conventional X-ray film method are now easier to see, improving diagnostic performance. 0: Compared to the conventional X-ray photographic film method,
Although it is becoming easier to see, there is no particular improvement in diagnostic performance. -1: There were areas where diagnostic performance improved, but there were also areas where it was difficult to diagnose. -2: There was no area where the diagnostic performance was improved, and there were areas where it was difficult to diagnose. Emphasized frequencies in Figure 4 a and b (c in Figure 2 b)
The results of the relationship between and evaluation are shown. Figures a and b are typical examples of frontal chest photography and bone photography, respectively. The thin solid line () is the result of performing the aforementioned ultra-low frequency processing with the emphasis coefficient β fixed at β=3. As is clear from comparing a and b, the areas with high evaluation values (areas with improved diagnostic performance) are closer to the lower frequency side than the bones in the frontal chest image. As will be seen, the frequencies to be emphasized differ depending on the case and location. The broken line ( ) is an example in which the emphasis coefficient β is continuously changed according to the original image signal.
For both a and b, the evaluations on the low frequency side and high frequency side are both high. This is because the former prevents white spots from occurring in the heart and bone regions (including the spine), and the latter prevents the increase in noise, which is the problem in areas that are difficult to diagnose in the -1 section of the evaluation criteria mentioned above. This is because the occurrence was prevented and the score moved from -1 to +1 or +2. In this chest example, the integral histogram is 10
% brightness is S ₀ (this almost corresponds to the maximum brightness of the spine), the brightness of 50% is S ₁ (corresponds to the lowest brightness of the lung field), β at brightness S ₀ is 0, and brightness S ₁
β is set to 3, and the value is changed linearly between these values. The dashed-dotted line () is the one that has been subjected to gradation processing in addition to the above processing, and the chest X-ray image a has been processed to lower the contrast of the heart and increase the contrast of the lung field. Image b was processed to increase the overall contrast by 1.5 times. The thick solid line () is the evaluation result of the image further reduced to 1/2 to 1/3 and presented. In the case of gradation processing, as mentioned above, lung cancer,
Contrast has increased and diagnostic performance has improved for diseases that change slowly over a large area, such as in king meat species. Furthermore, through reduction processing, the ultra-low frequencies that are important for diagnosis approach the optimal frequency (1 to 2 cycles/mm) of the modulation transfer function for human vision, so the contrast appears to be higher and the diagnostic performance is improved. Improved. At the same time as emphasizing ultra-low frequency components, applying smoothing processing to reduce the modulation transfer function to 0.5 or less in the frequency range of 0.5 to 5 cycles/mm will remove noise (graininess) on the photographic image and improve diagnostic performance. Improved. FIG. 5 is a diagram showing the effective range of the degree of enhancement for chest photographs. In this case, the frequency range to be emphasized is fixed, that is, c is fixed at c=0.1, and photographic images are created and evaluated by varying the emphasis coefficient β. Curve a in Figure 5 is the result when β is constant regardless of the original image signal, and curve b is the result at the maximum B/A value when it is continuously varied depending on the original image signal. be. When β of curve a is constant, when B/A is 6 to 7 or more, false images become noticeable and the evaluation becomes 0 or less, but if β is made variable, false images are removed and 1.5≦B/A≦10 The evaluation was 0 or more in the range of . Improvements in diagnostic performance were also observed in various other cases within the range of approximately 1.5≦B/A≦10. Table 1 shows the range of c in which the evaluation was 0 or more, that is, the diagnostic performance was improved when similar ultra-low frequency processing was applied to other parts and cases. (This frequency is only on the original photo.)

[Table] As can be seen from this table, frequencies important for diagnosis are distributed in a very low frequency region, approximately 0.01≦
c≦0.5 cycles/mm. The improvement of diagnostic performance by combining ultra-low frequency enhancement with other processing (change of emphasis coefficient β, gradation processing, reduction, smoothing processing) was carried out on the various cases mentioned above, and in both cases the diagnosis The results show that the performance is further improved. Since the present invention having the above configuration emphasizes the frequency response from the ultra-low frequency region,
Frequency regions important for diagnosis are greatly emphasized. Therefore, contrast is improved and diagnostic performance is improved. Furthermore, by changing the degree of enhancement according to brightness, shape, etc., it is possible to prevent false images from occurring and from making it difficult to see diseases important for diagnosis. Furthermore, since high-frequency components are not emphasized, noise components are reduced, resulting in a smoother image. As a result, an easily visible photographic image can be obtained. All of these image processing techniques are ultimately designed to bring the modulation transfer function closer to the optimum frequency for human vision, making their effects even more pronounced. Reduction is particularly effective.

[Brief explanation of the drawing]

FIG. 1 is a flowchart showing the method of the present invention;
Figure 2 is a graph showing the steps of frequency emphasis, Figure 3 is a graph showing an example of a combination of emphasis coefficient and density, Figure 4 is a graph showing frequencies to be emphasized and evaluation of their diagnostic performance, and Figure 5. is a graph showing evaluation of enhancement coefficients and diagnostic performance.

Claims (1)

[Scope of 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 it as a visible image, at each scanning point Determine the unsharp mask signal Sus corresponding to the ultra-low spatial frequency of (Sorg-Sus) A radiation image processing method characterized in that frequency components above the ultra-low spatial frequency are emphasized by performing the calculation. 2 As a non-sharp mask, the modulation transfer function is 0.5 or more at a spatial frequency of 0.01 cycles/mm and
The radiation image processing method according to claim 1, characterized in that a non-sharp mask having a sharpness of 0.5 or less is used at a spatial frequency of 0.5 cycles/mm. 3. The radiation image processing method according to claim 1 or 2, characterized in that the emphasis coefficient β is changed monotonically in accordance with an increase in the value of the original image signal or the non-sharp mask signal. 4. The method according to claim 3, wherein the maximum modulation transfer function of the photographic image emphasized by the arithmetic expression is 1.5 to 10 times the modulation transfer function near the zero spatial frequency. Radiographic image processing method. 5. A radiation image processing method according to claims 1 to 4, characterized in that the reproduced photographic image is smaller than the image stored on the phosphor. 6. An excitation light source for scanning a stimulable phosphor to stimulate the radiation image stored and recorded therein, a photodetector for detecting this emission and converting it into an electrical signal, and this electrical signal. In a signal processing device in a radiation image recording and reproducing system, the signal processing device is equipped with a calculation device that processes the detected original image signal as Sorg, and a non-sharp mask signal corresponding to the ultra-low spatial frequency at each detection point. 1. A radiation image processing device characterized in that, where S is Sus, β is an emphasis coefficient, and S is a reproduced image signal, the following calculation is performed: S′=Sorg+β(Sorg−Sus). 7 The arithmetic device receives the original image signal.
7. The radiation image processing apparatus according to claim 6, further comprising emphasis coefficient variable means for monotonically increasing the emphasis coefficient β in accordance with the magnitude of Sorg or the unsharp mask signal Sus.