JPH07160876A - Radiographic picture processor - Google Patents

Abstract

PURPOSE:To restrict the influence of quantization noise and also process a frequency without permitting unnecessary blurring to occur in a picture by providing plural smoothing filtering means with mutually different modulation transmitting features. CONSTITUTION:A picture signal is read from a picture memory 52 and inputted to a histogram analyzing part 54 and the histogram of picture element data is obtained so that a threshold value for changing-over smoothing filters 56 is obtained based on the histogram. Then, the picture signal is successively read from the picture memory 52 again at every picture element data, inputted to the plural smoothing filtering means 56 with mutually different modulation transmitting functions and also inputted to a comparator 58. The threshold value read from a threshold value storage memory 57 is also inputted to the comparator 58, the inputted threshold value is compared with the value of picture element data and a multiplexer 59 is changed-over based on the comparison result. Then, picture element data smoothing-processed by a smoothing filtering means 56, which corresponds to the picture element data is outputted from the multiplexer 59.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a radiation image processing apparatus for performing frequency processing and smoothing processing on an image formed by radiation transmitted through a subject.

[0002]

2. Description of the Related Art Conventionally, radiographic images such as X-ray images have been widely used for disease diagnosis. Taking an X-ray image as an example, the phosphor layer (fluorescent screen) is irradiated with X-rays that have passed through the subject, whereby the X-rays are converted into visible light, and this visible light is irradiated onto the silver salt film. An X-ray image is obtained by forming a latent image and developing it, and the X-ray image thus obtained is used for disease diagnosis (hereinafter,
Called "S / F method"). When the number of X-ray films thus obtained increases, a large space is required for storage, and when observing changes over time in the same subject (for example, patient) due to illness, X-ray films for comparison. There is a problem that it is difficult to take out the linear film. Therefore, in recent years, X obtained on a silver salt film as described above
A line image is photoelectrically read by a so-called film digitizer to obtain an image signal, and by subjecting this image signal to image processing, various image performance for determining image quality such as sharpness, dynamic range, and graininess, and for disease diagnosis. A system for obtaining a reproduced image with high image quality and high diagnostic performance after improving the diagnostic performance has also been used.

On the other hand, a system using a stimulable phosphor (stimulated phosphor) is beginning to be used in place of the system using the silver salt film. The system using this stimulable phosphor means that a stimulable phosphor panel (including a sheet) in which the stimulable phosphor is formed into a sheet or panel is irradiated with X-rays transmitted through an object. This is a system for accumulating and recording an X-ray image on a panel, then photoelectrically reading this X-ray image to obtain an image signal, performing image processing on the image signal, and then obtaining a reproduced image. The method is described in US Pat. No. 5,859,527. Here, the stimulated phosphor means X-ray, α-ray, β
When irradiated with radiation such as gamma rays and gamma rays, a part of the energy of the radiation is accumulated inside for a while or for a long time, and excitation light such as infrared light, visible light, and ultraviolet light is irradiated during that time. Then, it refers to a phosphor that emits the accumulated energy as stimulated emission light. Depending on the type of the phosphor, the type of radiation that easily accumulates energy, the wavelength of the excitation light that easily emits stimulated emission light, and the emission The wavelengths of stimulated emission light are different.

FIG. 1 is a diagram showing an example of the configuration of a system using a photostimulable phosphor panel. The system shown in FIG. 1 is an example in which a camera and a reader are separately configured. In the camera 10, the subject 12 standing in front of the shooting table 14
The X-rays 13 generated by the X-ray generator 11 are radiated on the illuminating phosphor panel 15, and the X-rays 13 transmitted through the subject 12 are radiated on the photostimulable phosphor panel 15 provided on the imaging table 14, thereby stimulating the photostimulable fluorescence. An X-ray image of the subject 12 is accumulated and recorded on the body panel 15.

After the photographing is performed in this manner, the photostimulable phosphor panel 15 is taken out from the photographing stand 14 and the reader 2
It is set in the panel insertion part 21 of 0. In this case, the stimulated phosphor panel 15 may be housed in a magazine or a cassette. If the photostimulable phosphor panel 15 set in the panel insertion portion 21 is stored in a magazine or cassette, it is taken out from the magazine or cassette and then conveyed along the conveyance path 22.
The reading unit 23 reads the X-ray image stored and recorded in the photostimulable phosphor panel 15 to generate an image signal. The configuration of the reading unit 23 will be described later. The image signal generated by the reading unit 23 is transmitted through the signal transmission path 24.
Is input to the image processing unit 25 via.

In the image processing section 25, conventionally, a large number of pixel data representing the X-ray transmission amount of each of a large number of pixels forming an X-ray image is S, a mask area surrounding a predetermined pixel on the X-ray image is M, The weighting coefficient of each pixel inside the mask area M is w, and the pixel data S of each pixel inside the mask area M is weighted by each weighting coefficient w, and the blurred mask data obtained by adding them to each other is S _us. , Pixel data S
Or the unsharp mask data S _us a variable, when the monotone decreasing function is allowed to contain a region that does not change K, the pixel data after the image processing was is Q, wherein Q = S + K · (S -S us) The processing according to is executed while sequentially changing the pixels. As a result, the X-ray image is frequency-processed.

The X-ray image mainly contains quantum noise due to the X-ray quantum, and the quantum noise is emphasized when the above frequency processing is performed. Therefore, the image processing unit 2
In No. 5, in addition to the above-described frequency processing, smoothing processing for suppressing high frequency components of the X-ray image is also performed, whereby high frequency components of increased noise are suppressed. Either of the above frequency processing and smoothing processing may be performed first.

The image signal, which has been subjected to appropriate image processing such as frequency processing and smoothing processing in the image processing section 25, is input to the image display section 27 via the signal transmission path 26,
For example, an X-ray image of the subject 12 is displayed on the CRT display screen. An image recording device such as a laser printer (not shown) may be provided instead of or together with the image display unit 27 for displaying an image, and an X-ray image may be reproduced and recorded on a silver salt film, for example. You may make it develop and obtain the X-ray image as a hard copy.

The photostimulable phosphor panel 15 read by the reading section 23 is erased by the erasing section 29 along the transport path 28.
Be transported to. In the erasing section 29, the stimulable phosphor panel 15 is irradiated with erasing light, so that the energy (afterimage) remaining in the stimulable phosphor panel 15 is erased. Photostimulated phosphor panel 1 from which this afterimage has been erased
The sheet 5 is conveyed to the panel take-out section 31 along the conveying path 30, taken out of the reader 20, set in the photographing device 10, and reused.

FIG. 2 is a diagram showing another system configuration example using the stimulated phosphor panel. In this figure,
The components corresponding to the respective components of the system shown in (1) are assigned the same numbers as those given in FIG. 1, and only the differences will be described. The system shown in FIG. 2 is provided with a standing type image pickup device 40 in which the image pickup stand 14 and the reader 40 of the image pickup device 10 in the system shown in FIG. 1 are integrally configured. The photostimulable phosphor panel 15 includes a photographing unit 31.
Is imaged, is conveyed to the reading unit 23 along the conveyance route 22 for reading, is conveyed to the erasing unit 29 along the conveyance route 28 for erasing, and is further conveyed to the conveyance route 30. Along the line, it is set again in the photographing section 31 and reused for the next photographing.

FIG. 3 is a diagram showing an example of the configuration of the reading unit 23 shown as a block in FIGS. The photostimulable phosphor panel 15 on which the X-ray image is accumulated and recorded is conveyed (sub-scan) in the reading section shown in FIG. Also, during this conveyance (sub scanning), the laser beam 101
A laser beam 102 as excitation light emitted from the laser beam is repeatedly reflected and deflected by a scanner 103 such as a galvanometer mirror or a rotary polygon mirror (polygon mirror).
After passing through the beam shape correction optical system 104 such as an fθ lens, the reflected stimulable phosphor panel 15 is reflected by the reflection mirror 105 and is then irradiated onto the stimulable phosphor panel 15. Repeated scanning (main scanning) is performed in the X direction. From each point of this scanning, stimulated emission light carrying the X-ray image accumulated and recorded in the stimulated phosphor panel 15 is emitted. This stimulated emission light is condensed by a condenser 106 such as an optical fiber array, guided to a photomultiplier tube 108 via an optical filter 107 that cuts the excitation light and transmits the stimulated emission light. It is converted into an electric signal. The stimulated emission light may be directly received without using the condenser 106, for example, by using a CCD photosensor or the like having an optical filter attached to the front surface for transmitting only the stimulated emission light.

An electric signal obtained by the photomultiplier tube 108 is logarithmically amplified by a logarithmic amplifier 109 and then converted into a digital image signal S by an A / D converter 110. The sampling timing of the A / D converter 110 is controlled by the A / D conversion control unit 113. The digital image signal S is once stored in the frame memory 111 or directly in the storage medium 112 such as a magnetic disk or an optical disk without passing through the frame memory 111. After that, the image signal stored in the storage medium 112 is read out and input to the image processing unit 25 shown in FIGS. Alternatively, the image signal S is temporarily stored in the frame memory 111 and then stored in the storage medium 1.
It is directly input to the image processing unit 25 without passing through 12.

In the system using the photostimulable phosphor, the energy of the radiation applied to the photostimulable phosphor and the amount of the photostimulated luminescent light emitted by the irradiation of the excitation light are proportional over a wide energy range. This ratio can be changed depending on the light amount of the excitation light, so that a radiographic image that is not affected by fluctuations in the radiation exposure amount can be obtained, and imaging mistakes can be reduced. Further, in a system for obtaining an X-ray image of a human body, the exposure dose of the human body in X-ray photography can be reduced.

Further, in both the system using the film reader and the system using the photostimulable phosphor, since a digital image signal is obtained, there is a feature that the space for storage is small and the retrieval is easy. In addition, it has a feature that image processing is possible.

[0015]

As described above, the increase in noise can be suppressed by performing the smoothing process together with the frequency process, but the smoothing process also suppresses the high frequency component of the image signal. At the same time, there is a problem that an image is blurred. The high-frequency noise of the X-ray image is mainly quantum noise due to X-rays, and the smaller the X-ray dose, the greater the noise. In the case of an image reading apparatus using a photostimulable phosphor panel, the image signal intensity increases as the X-ray dose increases.
The magnitude of quantum noise increases as the image signal strength decreases.

For example, in a chest X-ray image, the lung field portion having a large image signal has less noise than the heart / mediastinum portion having a small image signal. Therefore, it is not necessary to smooth the lung field portion so much. Smoothing causes blurring of the image, which is an opposite effect. The present invention is
An object of the present invention is to provide an image processing apparatus that suppresses the influence of quantum noise and performs frequency processing without causing unnecessary blur in an image in consideration of the above problems in the conventional method.

[0017]

In a radiation image processing apparatus of the present invention which achieves the above object, a large number of pixel data representing a radiation transmission amount of each of a large number of pixels forming an image formed by radiation transmitted through an object. S, M is a mask region surrounding a predetermined pixel on the image, w is a weighting factor of each pixel inside the masking region M, and pixel data S of each pixel inside the masking region M is a weighting factor w. The blur mask data obtained by weighting with each other and adding to each other is S _us , the pixel data S or the blur mask data S _us
Where K is a monotonically decreasing function that is allowed to include a region that does not change, and Q is pixel data after image processing, the equation Q = S + K · (S−S _us ) ... (1) A frequency processing unit that executes frequency processing while sequentially changing the predetermined pixels, and a smoothing processing unit that executes smoothing processing that suppresses high-frequency components of the image before or after the frequency processing or simultaneously with the frequency processing. In a radiation image processing apparatus comprising: a plurality of smoothing filtering means having different modulation transfer characteristics, wherein the smoothing processing unit operates the smoothing filtering means according to the value of the pixel data or the blur mask data. It is characterized in that smoothing processing of the pixel data is performed by the smoothing filtering means switched while switching.

Here, the plurality of smoothing filtering means are the blur mask data S in the frequency processing section.
It may be provided independently of the filtering means for obtaining _us , or at least one of the plurality of smoothing filtering means obtains the blur mask data S _us in the frequency processing section. It may also serve as a filtering means for.

Further, the smoothing processing unit is used for the pixel data.
Set one or more thresholds between the minimum and maximum values of
Comparing the pixel data and the threshold with the threshold setting means for setting
Depending on the comparison means and the comparison result in the comparison means,
Switching means for switching the number smoothing filtering means and
It may be configured to have. In this case, the above threshold setting
The threshold value set by the means is set to the minimum value of the pixel data S and
And the maximum value, in order from the minimum value side, Th_i (I = 0,
1, 2, ...) and Th_i ≤ S <Th_{i + 1} Pixel de is
Modulation transmission of smoothing filtering means applied to data S
The function is H_i (U) (i = 0, 1, 2, ...) (u is frequency
And the Nyquist frequency is f _n And when
-F_n <U <f_n Within the range of_i (U) ≦ H_{i + 1} (U) It is preferable that (2) is satisfied.

In the case where the pixel data is pixel data representing each pixel value of an image of a human lung as a subject, the threshold value setting means sets the pixel at the boundary between the lung field and the heart / mediastinum. It is preferable to set the value as one of the thresholds, and it is also a preferable aspect to set the pixel value at the boundary between the heart and the mediastinum as another threshold. In the threshold setting means, for example, a histogram of pixel data is obtained, and the threshold is set based on the shape of the histogram.

[0021]

Since the radiation image processing apparatus of the present invention is provided with a plurality of smoothing filtering means and the smoothing filtering means according to the value of the pixel data performs the smoothing processing on the pixel data, the quantum noise Sufficient smoothing is performed on pixel data with a small conspicuous pixel value, or with a whitish (small amount of radiation transmission) area as a whole, and pixel data with a large pixel value with no conspicuous quantum noise or blackish (radiation as a whole) For a region having a large amount of transmission, the smoothing is suppressed to the extent that the image is not blurred, and as a result, a high quality image can be obtained as a whole.

[0022]

EXAMPLES Examples of the present invention will be described below. FIG. 4 is a block diagram of the first embodiment of the radiation image processing apparatus of the invention, which corresponds to the image processing unit 25 of FIGS. 1 and 2. The image signal generated by the reading unit 23 shown in FIGS. 1 and 2 is input to the radiation image processing apparatus shown in FIG. 5 via the signal transmission path 24, and is stored in the image memory 51 via the data bus 40. It is entered once. The image signal input to the image memory 51 is then read from the image memory 51 and input to the frequency processing unit 55. Frequency processing unit 5
5, the pixel data S representing the pixel value of each pixel of the image
(I, j) (i, j = 1, 2, ...; i and j represent the numbers of the pixels that intersect each other on the image in the respective predetermined directions), and the mask area on the image is M∋ (i = I-i +
Q, j = j to j + R) and the weighting coefficient of the pixel value of each pixel inside the mask region M is k (q, r), the blur mask data _Sus (i, j) of each pixel is

[0023]

[Equation 1]

Then, a monotonically decreasing function which is allowed to include an invariant region, in which the pixel data S or the blur mask data _Sus is used as a variable as the emphasis coefficient K and Q (i, j) = S is adopted. (I, j) + K · {S (i, j) −S _us (i, j)} (4) The frequency processing is performed to obtain the processed pixel data Q (i, j). The image signal that has been subjected to the above frequency processing in the frequency processing unit 55 is once stored in the image memory 5 once.
Stored in 2.

Next, the image signal is read from the image memory 52 and input to the histogram analysis section 54, and the histogram of the pixel data (i, j) is obtained. FIG. 5 is a diagram showing an example of a histogram. Here, a histogram of a chest image of a human body is shown. The horizontal axis represents the pixel value of the pixel data Q (i, j), and the vertical axis represents the appearance frequency of the pixel data of that pixel value.

Returning to FIG. 4, the description will be continued. The histogram analysis unit 54 obtains a histogram as shown in FIG. 5, and obtains a threshold value for switching the smoothing filter 56 described later based on the histogram. As the threshold value, for example, a pixel value at the boundary between the mediastinum and the heart, a pixel value at the boundary between the heart and the lung field, and the like are selected.
The threshold value obtained by the histogram analysis unit 54 is stored in the threshold value storage memory 57.

Next, the image signal is sequentially read again from the image memory 52 for each pixel data, input to the plurality of smoothing filtering means 56 having different modulation transfer functions, and also input to the comparator 58. . The threshold value read from the threshold value storage memory 57 is also input to the comparator 58, the input threshold value is compared with the pixel data value, and the multiplexer 59 is switched based on the comparison result. . The image data to be compared with the threshold value in the comparator 58 is pixel data of the pixel located in the center of the mask of the filter for smoothing, average value of the pixel data of each pixel in the mask, The median value of pixel data of pixels is used.

FIG. 6 shows the smoothing filtering means 56.
It is the figure which illustrated the modulation transfer function. Horizontal axis u is spatial frequency
Axis and f_n Represents the Nyquist frequency. H_i 
(U), H_{i + 1} (U) is i-th and i + 1-th, respectively
Represents the modulation transfer function of the smoothing filtering means 56 of
And -f_n <U <f_n At H _i (U) ≦ H
_{i + 1} (U).

When the threshold value including the minimum value and the maximum value of the pixel data is set to Th _i (i = 0, 1, 2, ...) In order from the minimum value side, it is input to the comparator 58 shown in FIG. When the pixel data S is Th _i−1 ≦ S <Th _i , the multiplexer 59 selects the i-th smoothing filtering means. By performing such selection, the pixel data with a small pixel value is sufficiently smoothed so that the graininess of the image due to the quantum noise is improved, and the pixel data with a large pixel value is grained. Since the smoothness is not so bad, smoothing is suppressed and blurring of the image is prevented.

In this manner, the multiplexer 59 shown in FIG. 4 outputs the pixel data which has been smoothed by the smoothing filtering means 56 corresponding to the pixel data. The pixel data output from the multiplexer 59 is sequentially transferred to the image memory 53 via the data bus 40.
Entered in. An image signal as a set of pixel data stored in the image memory 53 is subsequently read from the image memory 53 and input to the image display unit 27 shown in FIGS.

FIG. 7 is a block diagram of a second embodiment of the radiation image processing system of the invention. Only differences from the first embodiment shown in FIG. 4 will be described. In the first embodiment shown in FIG. 4, the pixel data sequentially read from the image memory 52 are simultaneously input to the plurality of smoothing filtering means 56 via the data bus 40. However, the second embodiment shown in FIG. In the example, a demultiplexer 60 is provided between the data bus 40 and the plurality of smoothing filtering means 56, and the data bus 40 is connected to the smoothing filtering means 56 according to the comparison result in the comparator 60. . Thereby, the plurality of smoothing filtering means 56
This prevents useless filtering operations from being performed at the same time.

FIG. 8 is a block diagram of a third embodiment of the radiation image processing system of the invention. Only differences from the first embodiment shown in FIG. 4 will be described. In the first embodiment shown in FIG. 4, a plurality of smoothing filtering means 56 composed of hardware are provided, but in the third embodiment shown in FIG. 8, a single filter calculation unit 76 is provided. The filter calculation unit 76 is provided with a filter coefficient according to the comparison result of the comparator 58, and a smoothing filtering process according to the set filter coefficient is performed. That is, in this embodiment, the filter calculation unit 76 in which each filter coefficient is set is considered as each smoothing filtering means in the present invention.

FIG. 9 is a block diagram of a fourth embodiment of the radiation image processing system of the invention. Differences from the third embodiment shown in FIG. 8 will be described. In the third embodiment shown in FIG. 8, the frequency processing unit 55 and the filter calculation unit 76 are provided separately, but in the fourth embodiment shown in FIG. 9, only the filter calculation unit 85 is provided, The filter calculation unit 85 is configured to also have the calculation function of the frequency processing unit 55 shown in FIG.

The above-mentioned equation (3) is for the frequency processing of the equation (4), but the equation (3) is also a kind of smoothing filtering operation. Therefore, like the filter operation unit 76 of the fourth embodiment shown in FIG. 9, one operation unit can be configured to perform both frequency processing and smoothing processing, thereby reducing the circuit scale and the like. Can be achieved.

Although the above embodiment is an example of performing the smoothing process after performing the frequency process based on the equations (3) and (4), the smoothing process may be performed before the frequency process. In that case, the threshold serving as a selection reference when selecting the smoothing filtering means is determined based on the image signal before frequency processing, which is input via the signal transmission path 24.

[0036]

As described above, according to the present invention,
It is possible to perform frequency processing on an image signal read by a radiation image reading device while suppressing the influence of quantum noise and preventing image blur.

[Brief description of drawings]

FIG. 1 is a diagram showing a configuration example of a system using a stimulated phosphor panel.

FIG. 2 is a diagram showing another system configuration example using a stimulated phosphor panel.

FIG. 3 is a diagram showing a configuration example of a reading unit 23 shown as a block in FIGS. 1 and 2.

FIG. 4 is a block diagram of a first embodiment of a radiation image processing apparatus of the invention.

FIG. 5 is a diagram showing an example of a histogram.

FIG. 6 is a diagram illustrating a modulation transfer function of a smoothing filtering unit.

FIG. 7 is a block diagram of a second embodiment of the radiation image processing apparatus of the invention.

FIG. 8 is a block diagram of a radiation image processing apparatus according to a third embodiment of the present invention.

FIG. 9 is a block diagram of a fourth embodiment of the radiation image processing apparatus of the invention.

[Explanation of symbols]

 40 data bus 51, 52, 53 image memory 54 histogram analysis unit 55 frequency processing unit 56 smoothing filtering unit 57 threshold value storage memory 58 comparator 59 multiplexer 60 demultiplexer 76, 85 filter operation unit

Continuation of the front page (51) Int.Cl. ⁶ Identification code Office reference number FI Technical display location G06T 1/00 H04N 5/32 9287-5L G06F 15/62 390 A

Claims (8)

[Claims] 1. A large number of pixel data representing a radiation transmission amount of each of a large number of pixels forming an image formed by radiation transmitted through an object, S, a mask region surrounding a predetermined pixel on the image, and M. The weighting coefficient of each pixel inside the mask area M is w, and the pixel data S of each pixel inside the mask area M is weighted by each weighting coefficient w, and the blurred mask data obtained by adding them to each other is S _us , When K is a monotonically decreasing function that is allowed to include a region that does not change and has pixel data S or blur mask data S _us as a variable, and Q is pixel data after image processing, the equation Q = S + K · (S− S _us frequency processing according to), and frequency processing unit for executing while sequentially changing the predetermined pixel, the high frequency component at the same time the image with longitudinal or the frequency processing of the frequency processing A radiation image processing apparatus comprising: a smoothing processing unit that executes a suppressing smoothing process; and a plurality of smoothing filtering means having different modulation transfer characteristics, wherein the smoothing processing unit includes the pixel data or the blur. A radiation image processing apparatus, wherein the smoothing filtering means is switched while the smoothing filtering means is switched according to the value of the mask data to perform the smoothing processing of the pixel data. 2. The blur mask data S in the frequency processing unit is provided by the plurality of smoothing filtering means.
The radiation image processing apparatus according to claim 1, wherein the radiation image processing apparatus is provided independently of a filtering unit for obtaining _us . 3. The at least one of the plurality of smoothing filtering means also functions as a filtering means for obtaining the blur mask data S _us in the frequency processing section. The described radiation image processing apparatus. 4. The comparison unit for comparing the pixel data with the threshold, wherein the smoothing processing unit sets one or a plurality of thresholds between the minimum value and the maximum value of the pixel data. The radiation image processing apparatus according to claim 1, further comprising: a means and a switching means for switching the plurality of smoothing filtering means according to a comparison result of the comparing means. 5. The threshold value set by the threshold value setting means is set to Th _i (i = 0, 1, 2, ...) In order from the minimum value side including the minimum value and the maximum value of the pixel data S, T
The modulation transfer function of the smoothing filtering means applied to the pixel data S with h _i ≦ S <Th _{i + 1} is H _i (u)
(I = 0, 1, 2, ...) (u represents frequency),
When the Nyquist frequency is f _n , -f _n <u <f _n
The radiation image processing apparatus according to claim 4, wherein H _i (u) ≦ H _{i + 1} (u) is satisfied within the range. 6. When the pixel data is pixel data representing each pixel value of an image of a human lung as an object, the threshold value setting means sets a pixel at a boundary between a lung field and a heart / mediastinum. The radiation image processing apparatus according to claim 4, wherein the value is set as one of the threshold values. 7. The radiation image processing apparatus according to claim 6, wherein the threshold setting means sets a pixel value at a boundary between the heart and the mediastinum as another threshold. 8. The radiation image processing apparatus according to claim 4, wherein the threshold value setting means obtains a histogram of the pixel data and sets the threshold value based on the shape of the histogram.