JP2008073515A - X-ray image system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide an X-ray image system capable of outputting a high quality X-ray image even in photography with low dose. <P>SOLUTION: In a photography device 10b, a subject W and an image detector 6 are set as separated from each other, and an X-ray beam of an X-ray energy of 15-30keV is radiated from an X-ray source 2 for phase contrast photography. After that, as data of the X-ray image detected by the image detector 6 are created at a reading device 10d, an image processing device 30 applies processes including a noise reducing process to the created X-ray image. The processed X-ray image is outputted to a film by a film outputting device 50a, or outputted on a display by a diagnostic reading device 50c. <P>COPYRIGHT: (C)2008,JPO&INPIT

Description

  The present invention relates to an X-ray image system that generates and outputs X-ray image data by X-ray imaging.

  Unlike new adults and infants (hereinafter referred to as children), most of the bone marrow is red bone marrow, so the effect of exposure is large. Therefore, it is required to reduce the exposure dose more than adults, and it is necessary to adjust the imaging conditions and reduce the X-ray dose in order to minimize the exposure dose when performing X-ray imaging of children and the like. It is.

In the X-ray image obtained by imaging with a low dose, the signal value decreases as a whole, and therefore noise increases. Conventionally, the obtained X-ray image has been dealt with by performing noise removal processing or noise reduction processing (see, for example, Patent Document 1).
JP 2002-125153 A

  However, if noise reduction processing is performed on an X-ray image taken with a low dose, a certain level of effect is obtained and the image quality is easy to see, but the sharpness is reduced due to the noise removal effect (noise removal). This is because the original signal portion is attenuated). Therefore, the boundary part of a site | part etc. becomes indistinct and the utilization property for a diagnosis becomes low. In order to avoid such a problem, the X-ray dose at the time of imaging cannot be reduced sufficiently, and as a result, it has been difficult to sufficiently reduce the exposure dose.

  An object of the present invention is to provide an X-ray image system capable of outputting a high-quality X-ray image even when photographing at a low dose.

The invention according to claim 1 is an X-ray imaging system,
An X-ray source that irradiates the subject with X-rays and an image detector that detects the X-rays transmitted through the subject; An imaging means for irradiating the subject with X-rays having an X-ray energy of 30 (keV) to perform phase contrast imaging;
Image generating means for generating X-ray image data by the phase contrast imaging;
Image processing means for performing image processing including noise reduction processing on the generated X-ray image;
Output means for outputting a processed image subjected to the image processing;
It is characterized by providing.

The invention according to claim 2 is the X-ray imaging system according to claim 1,
A focal diameter of the X-ray source is 30 to 200 (μm).

The invention described in claim 3 is the X-ray imaging system according to claim 2,
A focal diameter of the X-ray source is 50 to 120 (μm).

  According to a fourth aspect of the present invention, in the X-ray imaging system according to any one of the first to third aspects, an enlargement factor M of the phase contrast imaging is 1.2 ≦ M ≦ 5. And

The invention described in claim 5 is the X-ray imaging system according to claim 4,
The magnification M of the phase contrast imaging is set to 1.5 ≦ M ≦ 3.

The invention according to claim 6 is the X-ray imaging system according to any one of claims 1 to 5,
The image processing includes gradation conversion processing for changing the contrast of the X-ray image,
The image processing means performs gradation conversion processing on the X-ray image and then performs noise reduction processing to generate a processed image.

The invention according to claim 7 is the X-ray imaging system according to any one of claims 1 to 5,
The image processing includes gradation inversion processing for inverting the gradation of the X-ray image,
The image processing means performs the noise reduction processing on the X-ray image, and then performs the gradation inversion processing to generate a processed image.

  According to the first to fifth aspects of the present invention, an X-ray image is obtained by phase contrast imaging, and noise reduction processing is performed on the X-ray image. The X-ray image obtained by phase contrast imaging has the edge of the subject tissue edge-enhanced, so it is easy to extract the edge component by distinguishing it from the noise component in the noise reduction processing, and perform edge enhancement and noise suppression with high accuracy. Can do. That is, since a clear and high-quality processed image can be obtained, an X-ray image having a sufficient image quality for interpretation can be obtained even if the X-ray dose at the time of photographing is low in photographing for children. Therefore, when a child or the like is to be imaged, the exposure dose can be minimized. Furthermore, since the original image itself becomes clear due to the phase contrast effect, the image after noise reduction is still highly clear even if it is slightly blurred compared to the original image due to the influence of noise reduction processing.

  According to the sixth aspect of the present invention, not only edge enhancement and noise suppression can be performed, but also contrast and density range can be adjusted by combining gradation conversion processing and noise reduction processing.

  According to the seventh aspect of the present invention, not only edge enhancement and noise suppression are performed by combining gradation inversion processing and noise reduction processing, but also visibility of structures in the image is improved by gradation inversion. be able to.

  In the present embodiment, an X-ray image with high edge sharpness is obtained by phase contrast imaging, and noise reduction processing is performed on the X-ray image, thereby providing a high-quality X-ray image even under a low dose. An example of an X-ray imaging system capable of performing the above will be described.

First, the configuration will be described.
FIG. 1 shows a system configuration of an X-ray imaging system 1 in the present embodiment.
As shown in FIG. 1, the X-ray imaging system 1 includes imaging devices 10a to 10c, a reading device 10d, a control server 20, an image processing device 30, an image server 40, a film output device 50a, and image interpretation devices 50b and 50c. Has been. Each device is configured to be able to communicate with each other via a network N.

The X-ray image system 1 performs various image processing on the X-ray image data generated by X-ray imaging in the imaging devices 10a to 10c or the reading device 10d to generate processed images. Is output by the film output device 50a or the image interpretation devices 50b and 50c.
FIG. 1 is an example of the system configuration, and the present invention is not limited to this. What is necessary is just to comprise suitably, such as a change of the number of apparatuses, the addition of another apparatus, etc. as needed.

  The imaging devices 10a to 10c perform X-ray imaging and / or X-ray image data generation. The photographing device 10a is a CR (Computed Radiography) device for general photographing, the photographing device 10b is a CR device for photographing children, and the photographing device 10c is a CR device for mammography. The imaging apparatuses 10a to 10c include a type that performs X-ray imaging and generates X-ray image data, and a type that performs only X-ray imaging using a portable image detector. The reading device 10d reads an X-ray image from an image detector used for imaging in the latter type of imaging devices 10a to 10c and generates data thereof.

  The control server 20 is a management device that comprehensively controls a series of processes in the X-ray image system 1 from imaging by the imaging devices 10a to 10c to generation, storage, and output of X-ray images. A general computer device can be applied as the control server 20, and a management operation is performed using a management program.

The control server 20 performs management using the imaging order information. The imaging order information is imaging instruction information indicating conditions specified by a doctor with respect to imaging contents, image generation conditions, etc., and is not shown HIS (Hospital Information System) or RIS (Radiology Information System).
) And transmitted to the control server 20. The control server 20 analyzes the photographing order information and distributes the photographing order information to the photographing device corresponding to the photographing content among the photographing devices 10a to 10c.

  The control server 20 performs transfer control of the X-ray image generated in the imaging devices 10a to 10c or the reading device 10d to the image processing device 30. At this time, when child imaging is designated by the imaging order information, control information instructing to perform image processing under the image processing conditions for children is generated and transmitted to the image processing apparatus 30 together with the X-ray image.

  The image processing device 30 performs various image processing on the X-ray image and generates a processed image.

  The image server 40 stores and manages X-ray image data in a large-capacity memory. The image server 40 transmits X-ray image data to the film output device 50a, the image interpretation device 50b, and the like in accordance with instructions from the control server 20.

The film output device 50a forms and outputs an X-ray image on the film.
The image interpretation devices 50b and 50c are terminal devices used by a doctor to observe an X-ray image, and display and output the X-ray image on a monitor.

Next, the photographing apparatus 10b and the image processing apparatus 30 according to the present invention will be described in detail.
The imaging device 10b performs phase contrast imaging with a child or the like as the subject W.
As shown in FIG. 2, the photographing apparatus 10 b includes a photographing unit 3 and a main body unit 4 that performs photographing control and the like.
The imaging unit 3 is formed in an arm shape and configured to be movable up and down using the main body unit 4 as a support column. An X-ray source 2 and a holding unit 5 are arranged in opposition to the arm portion of the imaging unit 3. The holding unit 5 holds the image detector 6 and fixes its position. The holding unit 5 is configured to be able to move up and down along the column portion of the photographing unit 3. The photographing distance can be adjusted by raising and lowering the photographing unit 3 and the holding unit 5 and changing their height positions.

The main unit 4 includes a CPU (Central Processing Unit) and a RAM (Random Access Memory).
And a computer including a ROM (Read Only Memory), an operation unit, and the like. The main body 4 controls the raising / lowering of the imaging unit 3 and the raising / lowering of the holding unit 5 in accordance with an operation instruction, and also controls the X-ray irradiation operation by the X-ray source 2.

  The X-ray source 2 generates X-rays and irradiates the subject W. As the X-ray source 2, an X-ray tube having a focal diameter D (μm) suitable for phase contrast imaging for imaging a child or the like is used. The focal diameter D is preferably 30 to 200 (μm), more preferably 50 to 120 (μm).

  The subject W to be imaged is a child or the like. The child or the like takes a picture while being accommodated in the incubator 11 and adjusts the height positions of the X-ray source 2 and the cage 5 according to the height position of the child or the like. As shown in FIG. 2, the incubator 11 is provided with a child-care case on a bed-like support.

  The image detector 6 has a phosphor plate 7 accommodated in a casing called a cassette. The phosphor plate 7 is an X-ray detector that absorbs and stores X-ray energy. The image detector 6 detects the X-rays irradiated from the X-ray source 2 and reached through the subject W by the phosphor plate 7. Thereafter, the image detector 6 is loaded into the reading device 10d and image visualization is performed. The reading device 10d irradiates the phosphor plate 7 with excitation light such as laser light and photoelectrically converts the stimulated light emitted from the phosphor plate 7 into an image signal, thereby generating the image signal.

In the present embodiment, an example of the image detector 6 in which the phosphor plate 7 is accommodated in the cassette will be described, but an FPD (Flat Panel Detector) or the like can also be applied as the image detector 6.
The FPD is a plate in which conversion elements that generate an electric signal according to an incident X-ray dose are arranged in a matrix, and differs from the phosphor plate 7 in that an electric signal is directly generated in the FPD. When the FPD is applied, the electric signal generated in the FPD is A / D converted, and the obtained digital image data is output to the main body unit 4.

Next, phase contrast imaging performed by the imaging unit 3 will be described.
Phase contrast imaging is different from normal magnification imaging in that imaging conditions such as imaging distance and X-ray irradiation conditions are adjusted so that an edge enhancement effect is obtained.
FIG. 3 is a diagram for explaining the outline of phase contrast imaging.
In the case of general imaging, the image detector 6 is installed adjacent to the subject W so that the image detector 6 immediately receives X-rays transmitted through the subject W. Therefore, the X-ray image obtained by general imaging is approximately the same size as the life size (which means the same size as the subject W).

  On the other hand, in phase contrast imaging, a distance is provided between the subject W and the image detector 6 as shown in FIG. That is, the subject W and the image detector 6 are separated from each other. In this case, since the X-rays irradiated from the X-ray source 2 in the shape of a cone beam pass through the subject W and still reach the image detector 6 in the shape of a cone beam, the obtained X-ray image is smaller than the life size. The enlarged size.

In this case, the enlargement ratio M is set such that the distance from the X-ray source 2 to the subject W is R1, the distance from the subject W to the image detector 6 is R2, and the distance from the X-ray source 2 to the image detector 6 is R3 (R3). = R1 + R2), it can be obtained by the following formula (1).
M = R3 / R1 (1)
The enlargement ratio M can be adjusted by changing the ratio of the distances R1 and R2.

In the photographing apparatus 10b, as disclosed in Japanese Patent Application Laid-Open No. 2001-91479 and the like, by setting the R1, R2, R3, and the focal diameter D within predetermined ranges, the edge enhancement effect of the subject W edge can be obtained. The obtained phase contrast imaging is performed.
The imaging conditions described in Japanese Patent Laid-Open No. 2001-91479 are such that the focal diameter D of the X-ray source 2 is D ≧ 30, and the distance R1 is R1 ≧ (D-7) / 200, preferably 0.3 to 1.0. (M), and the distance R2 is R2 ≧ 0.15, preferably 0.3 to 1.0 (m). Under this condition, an edge enhancement effect can be obtained in an X-ray image obtained by imaging. Taking into consideration the overall length of the shooting room, shooting within this range is preferable.

The edge enhancement effect will be described.
In the X-ray image obtained by the above phase contrast imaging, as shown in FIG. 3, the X-ray refracted by passing through the edge of the subject W overlaps and overlaps with the X-ray passed through the subject W. A phenomenon occurs in which the X-ray intensity of the portion increases. For this reason, a phenomenon occurs in which the X-ray intensity is weakened at the inner edge of the subject W. As a result, an edge enhancement function (also referred to as an edge effect or a phase contrast effect) in which an X-ray intensity difference is widened at the border of the subject W works, and a highly visible X-ray image in which the border portion is sharply depicted. Will be obtained.

When the X-ray source 2 is regarded as a point source, the X-ray intensity at the edge portion is as shown by the solid line in FIG. E shown in FIG. 4 indicates the half-value width of edge emphasis and can be obtained by the following equation (2). The half-value width E indicates the distance between the peaks and valleys of the edge.

  However, in medical sites and non-destructive inspection facilities, a Coolidge X-ray tube (also referred to as a thermionic X-ray tube) is widely used as the X-ray source 2, and this Coolidge X-ray tube has a large focal diameter. Cannot be considered an ideal point source. In this case, as shown in FIG. 5, a half-width E for edge emphasis is widened and a geometrically unsharp phenomenon occurs in which the intensity is lowered, resulting in an X-ray intensity as indicated by a dotted line in FIG. This geometrical sharpness is called blur.

式中、δ及びｒの定義は式（２）と同じである。
また、ＥＢはボケが無い場合のエッジ強調半値幅Ｅにボケの大きさを示すＢを加え、ＥＢ＝Ｅ＋Ｂで示される。
上述のように、Ｘ線源の焦点径Ｄは小さければ小さいほど、ボケが減少し、得られるエッジ効果が大きくなる。 If the half-value width of edge enhancement when blur occurs is EB, EB can be obtained from the following equation (3).

In the formula, definitions of δ and r are the same as those in the formula (2).
EB is represented by EB = E + B by adding B indicating the size of the blur to the edge emphasis half width E when there is no blur.
As described above, the smaller the focal diameter D of the X-ray source is, the more blur is reduced and the edge effect obtained is increased.

  When a child or the like is an object to be photographed, the photographing apparatus 10b performs photographing based on photographing conditions for children. The imaging conditions for pediatrics are as follows. When performing phase contrast imaging, the focal diameter D of the X-ray source 2 is 30 to 200 (μm), preferably 50 to 120 (μm), and the magnification ratio M is 1.2 ≦ M. ≦ 5, preferably 1.5 ≦ M ≦ 3, distance R1 is R1 ≧ (D−7) / 200, preferably 0.3 to 1.0 (m), distance R2 is R2 ≧ 0.15, preferably Is 0.3 to 1.0 (m). Since children and the like have a small subject size, the enlargement ratio M can be made larger than that of an adult. Further, when the X-ray source 2 uses a W (tungsten) tube, it is preferable that the tube voltage is as low as 20 to 50 (kVp) and the X-ray energy to be irradiated is 15 to 30 (keV). Unlike cartilage, pediatric cartilage is close to soft tissue (muscle, fat, etc.). Therefore, by setting the energy range as described above, it is possible to increase the contrast (signal value difference) between the cartilage and the soft tissue around the cartilage.

  As shown in FIG. 6, only the X-ray source 2 is mounted on the imaging unit 3 of the imaging apparatus 10b, and the incubator 11 is provided with a holding unit that holds the image detector 6 below the subject W. It may be done. In this case, since the distance R2 between the subject W and the image detector 6 is fixed, the magnification rate M is adjusted by the distance R1 between the X-ray source 2 and the subject W. Further, the photographing apparatus 10b may be movable.

Next, the image processing apparatus 30 will be described.
As shown in FIG. 7, the image processing apparatus 30 includes a control unit 31, an operation unit 32, a display unit 33, a communication unit 34, a storage unit 35, an image memory 36, and an image processing unit 37.

The control unit 31 includes a CPU (Central Processing Unit), a RAM (Random Access Memory).
) And the like, and various control programs stored in the storage unit 35 are read out by the CPU and expanded in the RAM, and various calculations and centralized control of each unit are performed according to these programs.

The operation unit 32 includes a mouse, a keyboard, and the like, generates operation signals corresponding to these operations, and outputs them to the control unit 31.
The display unit 33 includes a display and displays various display screens according to the control of the control unit 31.

The communication unit 34 includes a communication interface such as a network interface card, and communicates with an external device on the network N.
The storage unit 35 stores various programs executed by the control unit 31, an image processing program executed by the image processing unit 37, parameters and data necessary for the execution, and the like.
The image memory 36 is a memory for storing an X-ray image to be image processed and a processed image after processing.

The image processing unit 37 executes various image processes according to the image processing program. As image processing, it is possible to perform various image processing such as gradation conversion processing, gradation inversion processing, and noise reduction processing after performing irradiation field recognition processing and region of interest setting as preprocessing.
The image processing unit 37 predetermines image processing conditions for pediatric, general photographing, and photographing applications, and performs image processing according to the image processing conditions specified by the control unit 31.

As the image processing conditions, the type and order of image processing to be performed, processing parameters in each image processing, and the like are determined.
The image processing condition for children is to perform noise reduction processing after tone conversion processing. Alternatively, the gradation inversion process may be performed after the noise reduction process. The processing parameters for children used in each image processing will be described below together with the description of each image processing.

<Tone conversion processing>
The gradation conversion process is a process for adjusting density and contrast at the time of image output. When a doctor diagnoses the presence or absence of a disease in the human body structure by interpretation of the X-ray image, the presence or absence of the disease is determined based on the density and contrast (gradation) of the structure on the X-ray image. Therefore, by adjusting the density and contrast suitable for interpretation, a doctor's disease detection operation can be supported.

  The gradation conversion processing is performed in two stages, (1) normalization processing and (2) conversion processing using a basic LUT (lookup table), so that the desired signal value range and gradation characteristics are finally obtained. Tone conversion is performed. In order to maintain the interpretation ability (diagnostic performance) of doctors from the background that the screen / film method has been used for photography, the digital processing method using an image detector has been adopted. The input signal (read signal) is converted with the target of the gradation characteristics (contrast) cultivated in the above.

  The gradation characteristic obtained by the screen / film system is an S-shaped curve CV as shown in FIG. In the gradation conversion process, an LUT indicating this gradation characteristic is prepared as a basic LUT, and after individual signal adjustment is performed on the target image by the normalization process, signal values are converted using the basic LUT. .

FIG. 9 shows the relationship between the X-ray dose detected by the image detector (in the case of the phosphor plate) and the signal value of the X-ray image finally output according to the X-ray dose.
In the coordinate system of FIG. 9, the first quadrant indicates the reading characteristic, and indicates the relationship between the X-ray dose reaching the image detector and the read signal value (analog signal value). The second quadrant shows normalization characteristics, and shows the relationship between the read signal value and the normalized signal value (digital signal value) after the normalization process is performed. The third quadrant shows tone conversion characteristics and shows the relationship between the normalized signal value and the output density value (digital density signal value) converted by the basic LUT. Here, the output density value is a 12-bit resolution of 0 to 4095.

  In the second quadrant, the straight line indicating the normalization characteristic can adjust the output value range (size between SH and SL) by changing the slope thereof, and can also change the contrast of the entire image. This slope is defined as a G value. Further, by changing the intercept of the straight line indicating the gradation conversion characteristic, the height of the entire output value range (SH-SL movement) can be adjusted, and thereby the density of the entire image can be changed. Let this intercept be an S value.

For example, when the normalization is performed using the straight line h2 and the straight line h3 shown in FIG. 9, the normalized signal value corresponding to the subject W corresponds to the straight line region of the basic LUT by increasing the G value. Contrast will be improved.
That is, the density range and contrast of the output image can be adjusted by controlling the gradient G value and intercept S value of the straight line indicating the gradation conversion characteristics as gradation conversion parameters.

The G value is determined by the following equation (4) for obtaining the gradient of the gradation characteristic curve CV in the screen / film system shown in FIG.
G = (J2-J1) / (logE2-logE1) (4)
here,
J1 = 0.25 + Fog, J2 = 2.0 + Fog, Fog = 0.2,
E1 and E2 are incident X-ray doses corresponding to J2 and J1, respectively.
In the case where each part of the human body such as the chest and breast is to be observed, the G value is generally about 2.5 to 5.0 in many cases.

Moreover, S value is calculated | required by following formula (5).
S = QR × K1 / K2 (5)
here,
QR is the quantization domain value,
K1 is the reached X-ray dose that results in a signal value 1535 (QR = 200, output density 1.2), and K2 is the actual reached X-ray dose of the pixel that has an output density of 1.2 in the image after gradation conversion.
The value of K1 is uniquely determined by the setting of the quantization region QR value before photographing.

<Tone reversal processing>
The gradation inversion process is a process for inverting the density corresponding to the gradation level. For example, in the case of the minimum gradation 0 (display density: black) to the maximum gradation 4095 (display density: white), the display is reversed when the minimum gradation 0 and the display density is white and the maximum gradation 4095 is displayed. Density: Assign to black.

<Noise reduction processing>
The noise reduction process is a process for reducing noise components included in the image signal of the X-ray image. Here, an example will be described in which the degree of noise reduction is increased in a low density part where noise is conspicuous, and the degree of noise suppression is reduced in a high density part where noise is not noticeable.

In the noise reduction process, multiresolution decomposition processing, edge / noise information acquisition processing, smoothing processing, edge enhancement / noise suppression processing, and restoration processing are sequentially executed.
First, multi-resolution decomposition processing, edge / noise information acquisition processing, smoothing processing, and edge enhancement / noise suppression processing will be described with reference to FIG. FIG. 10 is a diagram showing the relationship between each process and the image signal.

《Multi-resolution decomposition processing》
In the multi-resolution decomposition process, the original image signal Sin is subjected to multi-resolution conversion to obtain a plurality of non-sharp image signals gk ′ (k = 1 to L (L is an integer of 1 or more, the same applies hereinafter)) in a plurality of spatial frequency bands.
A low-pass filter is used for multi-resolution conversion. The original image signal Sin is subjected to filter processing 121 by a low-pass filter, and the processed original image signal Sin is sampled pixel by pixel (down-sampling), thereby generating an unsharp image signal g1. The unsharp image signal g1 has a size that is ¼ that of the original image.

  Next, the interpolation process 122 supplements pixels having a value of “0” at every sampling interval of the unsharp image signal g1. This interpolation is performed by inserting a pixel column and a pixel row having a value of “0” for each column and row of pixels constituting the unsharp image signal g1. It should be noted that the non-sharp image signal interpolated in this way is in a state where the change of the signal value is not smooth because the pixels in the case of “0” are inserted every other pixel. Then, after such interpolation, filtering is again performed by a low-pass filter to obtain a non-sharp image signal g1 ′. The unsharp image signal g1 ′ has a smooth signal value change as compared with the unsharp image signal just after the interpolation.

  This non-sharp image g1 'is made the same size as the original image signal by interpolating 0 and filtering every other pixel after making the image 1/4, and is higher than half the spatial frequency of the original image signal. The frequency is removed.

  Further, the non-sharp image signal g1 is filtered 121 in order to obtain a non-sharp image signal of a frequency band one step lower. As a result, the unsharp image signal g1 is further converted into an unsharp image signal g2 having a size of 1/4 (original 1/16) sampled every other pixel. The non-sharp image signal g2 is subjected to an interpolation process 122 to generate a non-sharp image signal g2 ′.

  By repeating such processing sequentially, it is possible to obtain the unsharp image signal gk and the unsharp image signal gk ′ in the first to nth frequency bands (resolutions) having different frequency characteristics.

《Edge noise information acquisition processing》
The edge noise information acquisition process acquires edge information and noise information in the original image signal Sin based on an unsharp image signal in any one of a plurality of frequency bands acquired by the multi-resolution decomposition process. It is. Here, an example of acquiring based on an unsharp image signal in the second frequency band will be described, but the reference frequency band is not the highest frequency band with a pixel size of about 0.5 to 1.0 (mm) pitch. A frequency band is preferred. For example, when the downsampling rate is 1/2 with an image having a pixel size of 0.175 pitch, the third frequency band is most preferable. This is a condition where the noise / contrast ratio is highest at a resolution of 0.5 to 1.0 mm with respect to a pseudo minute edge image obtained by adding Gaussian noise to an edge signal of about 30 step / 1 to 2 mm, and the edge is easily recognized. It is because it becomes.

First, an edge information acquisition method will be described.
The edge information includes edge component information Ev and edge direction information Ed. First, a method for acquiring the edge component information Ev will be described.
As shown in FIG. 10, a difference image signal M2 between the unsharp image signal g1 and the unsharp image signal g2 ′ is first obtained by the difference processing 126. The edge component information Ev is acquired from the difference image signal M2 by the comparison process 127.
In the comparison process 127, the target pixel is set in the difference image signal M2. When the signal value of the target pixel is compared with the threshold values Bh and B1 and the signal value is equal to or greater than the threshold value Bh, it is determined that the target pixel constitutes a positive edge component, and the signal value is equal to or less than the threshold value Bl Is determined that the target pixel constitutes a negative edge component. The target pixel has threshold values Bh and B1 are threshold values prepared in advance for discriminating positive edge components and negative edge components, respectively.

On the other hand, when the signal value is lower than the threshold value Bh but higher than the threshold value Bl, the adjacent pixel of the target pixel is referred to and it is determined whether the adjacent pixel is a positive or negative edge component. When the adjacent pixel has a positive or negative edge component, it is determined that the target pixel is an edge inflection point. The edge inflection point is a point in the vicinity of a signal value 0 that becomes a positive / negative transition of the edge component, and is a point that serves as a reference for the edge.
If the adjacent pixel does not have a positive or negative edge component, it is determined that the target pixel is a non-edge component.

When the component of the pixel of interest is determined, the determination result (either a positive edge component, a negative edge component, an edge inflection point, or a non-edge component) is associated with pixel position information as edge component information Ev, and a storage unit 35.
The above process is repeated for all pixels.

  FIG. 11A is a diagram schematically showing an image signal for a certain column of the difference image signal M2. FIG. 11B shows an edge component (positive edge component, negative edge component, edge inflection point) of the difference image signal M2 based on the edge component information Ev obtained for the difference image signal M2 of FIG. , Non-edge components) are shown schematically.

Next, a method for acquiring the edge direction information Ed will be described.
The edge direction information Ed is acquired by subjecting the unsharp image signal g2 ′ to a filter process 128 using a Sobel filter, a Prewitt filter, or the like.

Next, a method for obtaining the noise information Ns will be described.
If the difference image signal between the unsharp image signal gn-1 ′ and the unsharp image signal gn ′ is Mn (in the first frequency band, the difference image signal between the original image signal Sin and the unsharp image signal g1 ′), The noise information Ns can be acquired from, for example, a local variance value, entropy value information, and pixel position information of the difference image signal Mn in each frequency band.
The entropy value Se of the image can be obtained by the following equation (6) using the probability P (z) that a certain pixel takes the pixel value z and the gradation number Z of the image (“Digital” published by Wiley-Interscience) Image Processing 3rd Edition ").

  When the entropy value is lower than a predetermined value in the intermediate frequency band (second to nth frequency bands) and the entropy is large in the highest frequency band (first frequency band), the noise component is dominant at the pixel position. A pixel at the pixel position can be used as a noise component. Therefore, information indicating that the pixel is a noise component is stored in the storage unit 35 as noise information Ns in association with the information on the pixel position.

  The obtained edge information Ev, Ed, and noise information Ns are used in processes 1231 and 1232. In processes 1231 and 1232, based on the edge information Ev, Ed and / or noise information Ns, the original image signal Sin and the unsharp image signal gk (k = 1 to (L−1)) in each frequency band are smoothed, and the density Dependency correction processing is performed. The smoothing process is divided into an edge smoothing process and a noise smoothing process. That is, the process 1231 performs a density dependence correction process after the edge smoothing process, and the process 1232 performs a density dependence correction process after the noise smoothing process. A processed signal G01 is output from the process 1231, and a processed signal G02 is output from the process 1232.

<Smoothing process>
The filter used for smoothing is desirably a low-pass filter used in multi-resolution decomposition processing or a low-pass filter having a frequency response close to that. FIG. 12A and FIG. 12B show examples of the filter. The filter example shown in FIG. 12A is an example of a two-dimensional smoothing filter used when down-sampling of multi-resolution decomposition is performed by a Laplacian pyramid method in which binomial filtering is performed 8 times. In this smoothing filter, smoothing filter coefficients (5 taps) are set two-dimensionally. When the number of taps is 5 taps, the filter coefficient F (x) is expressed by functions shown in the following equations (7) and (8).

  On the other hand, the filter example shown in FIG. 12B is an example of a one-dimensional smoothing filter used when down-sampling of multi-resolution decomposition is performed by a Laplacian pyramid method in which binomial filtering is performed 8 times.

First, the edge smoothing process will be described.
In the edge smoothing process, the edge component information Ev of each pixel is acquired from the storage unit 35, and based on the edge component information Ev, the original image signal Sin and the unsharp image signal gk (k = 1 to ( L-1)) is determined for each edge component pixel. Then, the original image signal Sin and the unsharp image signal gk in each frequency band are subjected to one-dimensional smoothing only in the edge gradient direction by a smoothing filter (for example, see FIG. 12B) for the edge component pixels. Apply processing. As a result, an image signal mainly composed of a noise component and a frequency component in a frequency band one step lower can be obtained. That is, the edge smoothing process is to obtain an image signal in which the noise component is enhanced by reducing the signal strength of the edge component.

Next, the noise smoothing process will be described.
In the noise smoothing process, each of the original image signal Sin and the unsharp image signal gk (k = 1 to (L−1)) of each frequency band based on the edge component information Ev of each pixel acquired from the storage unit 35. The non-edge component and edge component pixels are discriminated. Then, for each image of the original image signal Sin and the unsharp image signal gk in each frequency band, a two-dimensional smoothing process is performed by using a smoothing filter (see, for example, FIG. 12A) for pixels of non-edge components. I do. For the edge component pixels, a one-dimensional smoothing process is performed only in a direction other than the edge gradient direction or in a direction perpendicular to the edge gradient direction by a smoothing filter (see, for example, FIG. 12B). As a result, only an image signal mainly composed of an edge component and a frequency component in a frequency band one step lower is obtained.

<Density-dependent correction processing>
In the density-dependent correction process, the image signal after the smoothing process is corrected in advance in order to suppress the generation of artifacts and noise generated in the noise suppression process 1241 and the edge enhancement process 1242 in the subsequent stage. The degree of correction is controlled by the density (image signal value) of the image signal. Specifically, the correction component value is obtained by the correction function R shown in FIG. 13, and the correction component value is subtracted from the image signal value obtained after the smoothing process.
The correction function R is to obtain a correction component value from the image signal value and contrast. The contrast is a signal value of a difference image signal between adjacent non-sharp image signals gk in adjacent frequency bands before density-dependent correction processing (for the original image signal Sin, a difference image signal from the unsharp image signal g1 in the highest frequency band). It is. The correction function R is prepared for each of the edge smoothing processing signal, the noise smoothing processing signal, the edge smoothing processing, and the noise smoothing processing, and is defined for each frequency band. Has been.

  The reason why such correction is performed is that, in the noise suppression processing 1241 and the edge enhancement processing 1242, if the level of processing is increased, the image quality may be adversely affected due to artifacts or noise amplification. The edge enhancement processing 1242 has a problem of partial image quality degradation due to the occurrence of artifacts, and the noise suppression processing 1241 has a problem of amplification of noise components. Such artifacts and noises are particularly noticeable in low density areas. Therefore, it is necessary to perform density-dependent correction processing so that the degree of edge enhancement and noise suppression processing is reduced in the low density portion.

  The difference image signal is an image signal mainly composed of a noise component if the edge smoothing process is performed, and an image signal mainly composed of an edge component if the noise smoothing process is performed. As shown in FIG. 13, the correction function R is designed to have a low density and the correction component value increases as the value of the differential image signal approaches 0, so that it is added to the original image signal Sin at the low density portion. It is possible to control so that the signal component to be reduced is small, that is, the degree of edge enhancement and noise suppression is small.

FIG. 14 is a diagram showing the response of the correction function R to the contrast. As shown in FIG. 14, the correction function R has a higher degree of correction in the high contrast portion than in the low contrast portion in any frequency band.
On the other hand, FIG. 15 is a diagram showing the response of the correction function R to the density.
FIG. 15A is a diagram showing a response to the density of the correction function R corresponding to the noise smoothing process. As shown in FIG. 15A, it can be seen that the degree of correction is greater in the low density part than in the high density part, and the degree of edge enhancement processing for the high density part is greater.
On the other hand, FIG. 15B is a diagram showing a response to the density of the correction function R corresponding to the edge smoothing process. As shown in FIG. 15B, it can be seen that the degree of correction is higher in the high density part than in the low density part, and the degree of noise suppression processing is higher in the low density part.
As described above, by the density-dependent correction process, noise can be reduced by increasing the degree of the noise suppression process and reducing the degree of the edge enhancement process in the low density part where noise is conspicuous. On the other hand, by controlling the noise suppression process to be small and the edge enhancement process to be increased in a high density portion where noise is not conspicuous, it is possible to reduce the occurrence of noise and artifacts due to the noise suppression process.

<Edge enhancement / noise suppression processing>
In the noise suppression process 1241, the processed signal G01 obtained by the process 1231 and the unsharp image signal g1 ′ are subtracted. Subtraction is performed between corresponding pixels of each image signal. Then, the difference image signal mainly composed of noise components obtained by subtraction is multiplied by a noise suppression coefficient βN1 to obtain a difference image signal N0 with reduced noise components. The noise suppression coefficient is prepared for each frequency band and is represented by βNk (−1 <βNk <0).

Similarly, in the edge enhancement process 1242, the processed signal G02 obtained by the process 1232 and the unsharp image signal g1 ′ are subtracted. Then, the difference image signal mainly composed of edge components obtained by subtraction is multiplied by the edge enhancement coefficient βE1 to obtain a difference image signal E0 in which the edge components are enhanced. The edge enhancement coefficient is prepared for each frequency band and is represented by βEk (0 <βEk <1).
Then, the difference image signals N0 and E0 are added by the addition process 125 to obtain the difference image signal B0.
By performing the above processing in each frequency band, a differential image signal Bk-1 (k = 1 to L) is obtained. The obtained difference image signal Bk−1 is stored in the storage unit 35.

  FIG. 16 shows the relationship between the processing response and density (characteristic shown by the solid line) when multiplied by the noise suppression coefficient βNk in the noise suppression processing 1241, and the processing response when multiplied by the edge enhancement coefficient βEk in the edge enhancement processing 1242. It is a figure which shows the relationship (characteristic shown with a dotted line) with a density | concentration. As shown in FIG. 16, the density-dependent correction processing is individually performed on the noise component signal (the signal after the edge smoothing process) and the edge component signal (the signal after the noise smoothing process). The degree of noise suppression processing is large, and the degree of edge enhancement processing is small. Thereby, noise can be made inconspicuous in the low density portion. Further, in a high density portion where noise is not conspicuous, the degree of noise suppression processing is small and the degree of edge enhancement processing is large. Thereby, in the high density portion, it is possible to further enhance the edge while reducing the degree of noise suppression processing and reducing artifacts caused by the noise suppression processing.

《Restore processing》
The restoration process is a process of adding the difference image signal bk-1 to the original image signal Sin and inversely transforming it as shown in FIG. In this embodiment, it is assumed that inverse transformation (restoration) is performed from a plurality of detail images by the Laplacian pyramid method. Although this Laplacian pyramid method enables rapid processing, other methods can also be used.

First, the differential image signal bL-1 in the lowest frequency band is interpolated to interpolate between the pixels to obtain an image signal bL-1 ′ having a quadruple magnitude. Next, the interpolated image signal bL-1 ′ and the non-sharp image signal bL-2 having a high one-stage frequency band are added together with corresponding pixels, and the added image signal (bL-1 ′ + bL-2) is added. obtain.
Next, this added image signal (bL-1 ′ + bL-2) is subjected to an interpolation process, and interpolation between pixels is performed to obtain an image bL-2 ′ having a size four times as large. Next, the interpolated image bL-2 ′ and the unsharp image signal bL-3 having a high one-step frequency band are added together to obtain the added image signal (bL-2 ′ + bL-3). .

  The above processing is repeated for the difference image signal bk having a higher frequency. Then, finally, the processed image signal Sout is obtained by adding the sum of the image signal b1 'and the highest resolution difference image signal b0 to the original image signal Sin.

The image processing conditions for children are as follows.
When performing tone conversion processing for children, the G value is set to a value larger than that for general imaging, and the noise suppression coefficient βNk of the noise suppression processing is set to a value larger than that for general imaging in the noise reduction processing. Increase the degree of noise suppression. Alternatively, the G value is set to a value smaller than the G value for general photographing, and the edge enhancement coefficient βEk in the edge enhancement processing is set to a value larger than that for general photographing in the noise suppression processing to increase the degree of edge enhancement.

In general, the G value at the time of gradation conversion processing of a child's hip joint image is often given in a range of about 2.0 to 2.2. The coefficients βNk and βEk related to noise suppression and edge enhancement vary depending on the X-ray dose of the image itself and the purpose of diagnosis, but βNk is about 0.2 to 0.6 and βEk is about 0 to 0.4.
For example, when image processing is performed on an X-ray image with an edge component SE = 100 and a noise component SN = 100 under image processing conditions of G value 2.0, βNk = 0.4, βEk = 0.2, The edge component SE and noise component SN in the region of interest after image processing are SE = 240 and SN = 120, respectively, and the ratio SE / SN is SE / SN = 240/120 = 2.0.

In order to keep SN constant and improve SE, for example, the G value may be set to 3.0, βNk = 0.6, and βEk = 0.2. That is, the G value is increased and βNk is increased.
On the other hand, in order to keep SE constant and improve SN, for example, the G value may be 1.5, βNk = 0.4, and βEk = 0.6. That is, the G value is decreased and βEk is increased.

  When a child or the like is to be imaged, phase contrast imaging is performed. In the case of an X-ray image obtained by phase contrast imaging, the edge of a human tissue or the like originally has an edge-enhanced image quality due to the phase contrast effect. Therefore, the edge component and the noise component can be extracted with high accuracy in the noise reduction processing, and the edge of the edge portion is further emphasized by performing the edge enhancement processing 1242 in the noise reduction processing. In addition, since imaging is performed in a short time using low energy X-rays, the signal value is low and noise is likely to occur compared to general imaging. However, the noise suppression processing 1241 in the noise reduction processing can reduce noise at a low signal value (low density portion) as much as possible, and a clearer image quality can be obtained by synergy with the edge enhancement. It becomes possible.

Next, the operation of the X-ray image system 1 will be described.
FIG. 18 is a flowchart showing a flow from imaging to output of an X-ray image in the X-ray image system 1.
When performing imaging for a child, the control server 20 transmits imaging order information to the imaging apparatus 10b that can perform imaging under the imaging conditions for children (step S11).
In the imaging device 10b, phase contrast imaging is performed according to imaging conditions for children according to the imaging order information (step S12). That is, the focal diameter D of X-rays irradiated from the X-ray source 2 is set to 30 to 200 (μm), the tube voltage is set to 50 kVp or less, the enlargement ratio M is set to 1.2 ≦ M ≦ 5, R1 ≧ (D-7) / 200, the distance R2 is in the range of R2 ≧ 0.15. Also, the shooting time is within 0.02 s.

  The photographed image detector 6 is loaded into the reading device 10d by the photographer. The reading device 10d executes a reading process and generates X-ray image data. Then, an image processing condition or the like is written in the header of the data of the generated X-ray image and used as incidental information related to image generation. Thereafter, the data of the X-ray image is transmitted from the reading device 10 d to the image processing device 30 under the control of the control server 20. The control server 20 identifies X-ray images as follows. First, the photographer inputs an identification number assigned to the image detector 6 used for photographing in advance, and the control server 20 stores the identification number in association with the photographing order information. When the X-ray image is read from the image detector 6 in the reading device 10d, the identification number of the image detector 6 is also read, and information on this identification number is attached to the X-ray image. The control server 20 can identify the X-ray image by referring to the identification number attached to the X-ray image to determine the imaging order information of the X-ray image based on the identification number.

The image processing apparatus 30 performs image processing on the X-ray image data under the image processing conditions for children under the control of the control server 20. The image processing is performed in the order of noise reduction processing after tone conversion processing or tone inversion processing after noise reduction processing (step S13). In the gradation conversion process, the G value is increased, and in the noise reduction process, the degree of noise suppression is increased. Alternatively, the G value is lowered, noise suppression is performed in the noise reduction processing, and edge enhancement with a high degree of enhancement is performed.
When each image processing is completed, the image processing apparatus 30 writes the image processing conditions in the header of the X-ray image data and uses it as incidental information related to the image processing. Thereafter, the X-ray image data is transmitted to the image server 40 under the control of the control server 20. In the image server 40, the X-ray image data is stored (step S14).

  The control server 20 receives a transmission request from the image interpretation devices 50b and 50c for the X-ray image data stored in the image server 40, and instructs to transfer the X-ray image data requested for transmission to the image interpretation devices 50a and 50b. Control information is transmitted to the image server 40. The image server 40 transmits X-ray image data designated in accordance with the control information to the image interpretation devices 50a and 50b.

  The image interpretation devices 50b and 50c display and output X-ray image data. When an instruction to output to the film is made by the doctor, the X-ray image data is transmitted to the film output device 50a. The film output device 50a performs film output of the received X-ray image data (step S15).

Imaging (Comparative Examples 1 and 2 and Examples 1-3) was performed under the following experimental conditions in an X-ray image system, and visual evaluation was performed on an output image obtained by outputting the obtained X-ray image on a film.
<Experimental conditions>
Subject: A hip phantom of a simulated phantom (Kyoto Science Co., Ltd.) imitating a child was photographed.
Imaging device: The imaging device used was a prototype manufactured by Konica Minolta.
X-ray source: In Comparative Examples 1 and 2, a conventionally used medical W anode X-ray tube (Varian A-192) was used. The focal diameter D is 1200 (μm). On the other hand, in Example 1-3, a W anode X-ray tube manufactured by Konica Minolta was used. The focal diameter D is 100 (μm).
Image detector: The phosphor plate used was the company's Regius plate.
Reading device: Regius model 190 manufactured by the same company was used. The reading pitch is 87.5 μm.
Output device: DRYPRO model 793 manufactured by the company

<Shooting conditions>
X-ray energy: X-ray energy was adjusted as shown in Table 1 by controlling the tube voltage.
Tube current: 100 mA
Shooting time: Variable in the range of 0.005 to 0.02 (s).
Shooting distance: In Comparative Example 1, the enlargement ratio M = 1, the shooting distance R1 was 0.5 m, and R2 was 0 m. In Examples 1 to 5 and Comparative Example 2, the enlargement ratio M = 2, and the shooting distances R1 and R2 were each 0.5 m.
<Image processing conditions>
Comparative Examples 1 and 2 and Example 1-3: No image processing Example 4, 5: G value 1.5 gradation conversion processing, βNk = 0.4, βEk = 0.6 noise reduction processing gave.

<Evaluation criteria>
The image quality of the hip joint portion of the X-ray image output and formed on the film was evaluated. Image quality evaluation comprehensively evaluates sharpness and graininess.
A: Visible very clearly.
○: Visible clearly.
Δ: Visible.
X: Not visible.
In accordance with the above evaluation criteria, seven image evaluators observed images on the film and evaluated the image quality.

なお、表１においてｍＡｓ値は管電流の値に撮影時間を乗算して求めた値で、Ｘ線量と比例する。すなわち、ｍＡｓ値が大きいほど、Ｘ線量は増加する。 <Evaluation results>
Table 1 shows the evaluation results.

In Table 1, the mAs value is a value obtained by multiplying the tube current value by the imaging time, and is proportional to the X-ray dose. That is, the larger the mAs value, the higher the X-ray dose.

Table 1 shows that the X-ray images according to Examples 1 to 5 have good image quality in terms of sharpness and graininess.
Comparing Comparative Example 1 and Example 2, the image quality of Example 2 is better with the same X-ray dose. This is an effect of the phase contrast imaging method.
Comparing Example 1 and Example 3, in Example 3, the imaging time is shortened, and although the X-ray dose is reduced to about 25% of the X-ray dose of Example 1, there is no problem in practical use. Is obtained. Therefore, according to the present invention, it is apparent that the same image quality as before can be realized even if the X-ray dose is reduced.
When Example 3 and Example 4 are compared, the image quality is improved in Example 4 although the X-ray dose is the same. This is the effect of the noise reduction processing performed in the fourth embodiment. Therefore, even if the X-ray dose is reduced, the image quality can be improved by performing the noise reduction processing according to the present invention.
When Example 2 and Comparative Example 2 are compared, the image quality is degraded in Comparative Example 2 although the X-ray dose is the same. This is because in Comparative Example 2, the X-ray energy was set higher than the upper limit (30 keV) of the range of the present invention. In order to obtain the effect of the phase contrast imaging method, it is necessary to adjust the X-ray energy to 30 KeV or less. .
Note that it is not preferable to set the value lower than the lower limit (15 keV) of the present invention because the exposure amount of the subject increases.

In Examples 3 and 4, a sufficient image quality is obtained in a short time of 0.005 s. Children are difficult to keep stationary during filming. Therefore, if an X-ray image with good visibility can be obtained even in such a short time, it is a very preferable configuration from the viewpoint of ease of imaging work and exposure to children. In addition, since the blur caused by the movement of the child that occurs during the photographing can be suppressed, the sharpness of the image is in a better direction than the conventional method of photographing for a long time. Furthermore, in Examples 1 to 3, the overall image quality is better than that of Comparative Examples 1 and 2. It is considered that one of the factors is that by setting the focal diameter to 100 μm and a small focal diameter, the spread of the blur B in the phase contrast effect is suppressed and the edge enhancement effect is enhanced.
In Examples 1 to 5, the same evaluation results as above were obtained when the enlargement ratio M was varied in the range of 1.2 to 5.

  As described above, according to the present embodiment, when a child or the like is to be imaged, an X-ray image is obtained by phase contrast imaging, and noise reduction processing is performed on this. The X-ray image obtained by phase contrast imaging has the edge of the subject tissue edge-enhanced, so it is easy to extract the edge component by distinguishing it from the noise component in the noise reduction processing, and perform edge enhancement and noise suppression with high accuracy. Can do. That is, since a clear and high-quality processed image can be obtained, an X-ray image with sufficient image quality for interpretation can be obtained even if the X-ray dose at the time of imaging is small in imaging for children. In addition, if priority is given to image quality, X-ray images can be obtained with sharper and higher image quality than normal imaging by performing imaging without adjusting the X-ray dose.

  In the noise reduction processing, it is possible to separately adjust the degree of edge enhancement and noise suppression by the edge enhancement coefficient βEk and the noise suppression coefficient βNk. Therefore, processing with a high degree of freedom is possible. Furthermore, by performing density-dependent correction processing, it is possible to suppress the occurrence of noise and artifacts associated with edge enhancement and noise suppression.

  Further, by performing gradation conversion processing and gradation inversion processing together with noise reduction processing, the image quality of the X-ray image can be effectively improved. For example, it is possible to adjust the noise component and the edge component by adjusting the G value in the gradation conversion process and the coefficients βEk and βNk in the noise reduction process. In addition, since the X-ray dose is low in pediatric imaging, the signal value of the entire X-ray image becomes small, and it is difficult to distinguish between noise and image signals. Thus, the visibility of noise can be further reduced.

In addition, the said embodiment is a suitable example to which this invention is applied, It is not limited to this.
For example, the sensitivity of the image detector may be improved in order to reduce the X-ray dose when photographing for children. It is conceivable to increase the photomultiplier sensitivity when using a phosphor plate for the image detector, and to increase the gain of the amplifier when using an FPD.

  Further, as the phosphor plate of the image detector, a granularity type that emphasizes graininess and a sharpness type that emphasizes sharpness may be selected. In the granularity type, the thickness of the phosphor plate is greatly adjusted. Since the amount of X-ray energy that can be stored is increased by increasing the thickness of the phosphor layer, the signal value increases as a whole, and a low-noise X-ray image can be obtained. On the other hand, the sharpness type is obtained by adjusting the thickness of the phosphor plate to be small. By reducing the thickness of the phosphor layer, it is possible to reduce reflection of incident X-rays in the phosphor layer, reduce blur due to reflection, and obtain a sharp X-ray image.

It is a figure which shows the system configuration | structure of the X-ray imaging system in this embodiment. It is a figure which shows an example of an X-ray imaging apparatus. It is a figure explaining a phase contrast imaging and a phase contrast effect. It is a figure which shows the relationship between the edge strength in a phase contrast effect, and a blur. It is a figure explaining the case where a blur arises in a phase contrast effect. It is a figure which shows the other example of an X-ray imaging apparatus. It is a figure which shows the internal structure of an image processing apparatus. It is a figure which shows the gradation conversion characteristic made into the target in a gradation conversion process. It is a figure which shows the flow of the conversion of the signal value at the time of a gradation conversion process. It is a figure which shows a part of flow of a noise reduction process. (A) is a figure which shows typically the image signal for one row with a difference image signal. (B) is a figure which shows the classification | category of the edge component about the image signal of (a), and a non-edge component. (A) is a figure which shows an example of a two-dimensional smoothing filter coefficient. (B) is a figure which shows an example of a one-dimensional smoothing filter coefficient. It is a figure which shows an example of the correction function used in a density | concentration dependence correction process. It is a figure which shows the response with respect to the contrast of the correction function of FIG. (A) is a figure which shows the response with respect to the density | concentration of the correction function corresponding to a noise smoothing process. (B) is a figure which shows the response with respect to the density | concentration of the correction function corresponding to an edge smoothing process. It is a figure which shows the processing response with respect to the density | concentration of edge emphasis processing, and the processing response with respect to the density | concentration of noise suppression processing. It is a figure which shows a part of flow of a noise reduction process. It is a figure explaining the flow of the whole operation | movement in a X-ray imaging system.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 X-ray imaging system 10b Imaging device 2 X-ray source 3 Imaging part 6 Image detector 7 Phosphor plate W Subject 10d Reading device 20 Control server 30 Image processing unit 37 Image processing unit 40 Image server 50a Film output device 50b, 50c Interpretation apparatus

Claims (7)

An X-ray source that irradiates the subject with X-rays and an image detector that detects the X-rays transmitted through the subject; An imaging means for irradiating the subject with X-rays having an X-ray energy of 30 (keV) to perform phase contrast imaging;
Image generating means for generating X-ray image data by the phase contrast imaging;
Image processing means for performing image processing including noise reduction processing on the generated X-ray image;
Output means for outputting a processed image subjected to the image processing;
An X-ray imaging system comprising:   The X-ray imaging system according to claim 1, wherein a focal diameter of the X-ray source is 30 to 200 (μm).   The X-ray imaging system according to claim 2, wherein a focal diameter of the X-ray source is 50 to 120 (μm).   The X-ray imaging system according to any one of claims 1 to 3, wherein an enlargement factor M of the phase contrast imaging is 1.2 ≦ M ≦ 5.   The X-ray imaging system according to claim 4, wherein an enlargement factor M of the phase contrast imaging is 1.5 ≦ M ≦ 3. The image processing includes gradation conversion processing for changing the contrast of the X-ray image,
The said image processing means produces | generates a processed image by performing the said noise reduction process after performing a gradation conversion process to the said X-ray image, The Claim 1 characterized by the above-mentioned. X-ray imaging system. The image processing includes gradation inversion processing for inverting the gradation of the X-ray image,
The said image processing means produces | generates a processed image by performing the said gradation inversion process after performing the said noise reduction process to the said X-ray image. X-ray imaging system.

Images