JP2015065991A - Scattered radiation correction device, scattered radiation correction method, and x-ray imaging device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a scattered radiation correction device capable of directly leading out an image obtained by reducing a scattered radiation component from an image obtained by X-ray imaging, while having short calculation time.SOLUTION: A scattered radiation correction device according to this invention performs scattered radiation correction on a measurement image obtained by X-ray imaging. The scattered radiation correction device includes: a normalized ideal value calculation unit for calculating a normalized ideal value, which is a normalized luminance value obtained by the whole X-ray, for each pixel in a measurement image; a scattered radiation component calculation unit for calculating a scattered radiation component of the measurement image for each pixel in the measurement image from the normalized ideal value and a normalized luminance value of the measurement image; and correction processing for removing the scattered radiation component calculated by the scattered radiation component calculation processing from a luminance value of each pixel in the measurement image, for each pixel in the measurement image.

Description

  The present invention relates to a scattered radiation correction apparatus, a scattered radiation correction method, and an X-ray imaging apparatus including a scattered radiation correction apparatus that reduce scattered radiation components contained in an image obtained by X-ray imaging by scattered radiation correction.

  An image obtained by X-ray imaging includes a direct line component resulting from a direct line transmitted through the subject and components other than the direct line component. As components other than the direct ray component, there are scattered ray components caused by scattered rays and components caused by light diffusion of the X-ray detection element and electrical noise generated from the system of the X-ray imaging apparatus. The scattered radiation component is more dominant. Examples of scattered radiation include scattered radiation from a subject and scattered radiation from carbon installed in front of a flat panel detector.

  Scattered ray components contained in an image obtained by X-ray imaging greatly reduce the quality of the image obtained by X-ray imaging, and are one of the causes that make it difficult to remove metal artifacts, for example.

  For this reason, various methods of reducing scattered radiation components for reducing scattered radiation components contained in images obtained by X-ray imaging have been proposed.

Japanese Patent No. 4218908

  The scattered radiation component reduction method proposed in Patent Document 1 obtains reconstructed volume data from data including scattered radiation, estimates the distribution of scattered radiation contained in the projected image on the reconstructed volume data, and performs projection. Since the process of subtracting the scattered radiation component from the image and reconstructing it again takes time.

  In view of the above situation, the present invention can directly derive an image obtained by reducing scattered radiation components from an image obtained by X-ray imaging, and has a short calculation time, a scattered radiation correction apparatus, a scattered radiation correction method, And an X-ray imaging apparatus.

  In order to achieve the above object, in the scattered radiation correction apparatus according to the present invention, a scattered radiation correction apparatus that performs scattered radiation correction on a measurement image obtained by X-ray imaging, and a normalized luminance value of all X-rays. A normalized ideal value calculation unit that calculates a normalized ideal value for each pixel of the measurement image, the normalized ideal value for each pixel of the measurement image, and the normalized luminance value of the measurement image From the scattered radiation component calculation unit that calculates the scattered radiation component of the measurement image, and the scattered radiation component calculated by the scattered radiation component calculation processing unit from the luminance value of the pixel of the measurement image for each pixel of the measurement image And a correction process for removing (a first configuration). Thereby, the image which reduced the scattered ray component can be derived | led-out directly from the measurement image obtained by X-ray imaging. In addition, since the reconfiguration twice as in Patent Document 1 is unnecessary, the calculation time can be shortened. In dental CT imaging, since the subject is limited to the human head (particularly the jaw), it is easy to shorten the calculation time.

  In the scattered radiation correction apparatus having the first configuration, the scattered radiation component calculation unit includes, for each pixel of the measurement image, the normalized ideal value, the normalized luminance value of the measurement image, and the X-ray. A configuration (second configuration) may be employed in which the scattered radiation component of the measurement image is calculated from the projection angle of imaging. Thereby, appropriate scattered ray correction can be performed according to the projection angle of X-ray imaging.

  In the scattered radiation correction apparatus having the first or second configuration, a storage unit that stores a relationship between a normalized luminance value of all X-rays and a ratio of a scattered dose included in the total X-ray dose to the total X-ray dose The scattered radiation component calculation unit may calculate a scattered radiation component of the measurement image using the relationship stored in the storage unit (third configuration).

  In the scattered radiation correction apparatus having the third configuration, the storage unit stores a constant value used in the function indicating the relationship in the form of a data table associated with the normalized ideal value (fourth) It is good also as a structure. Thereby, the used capacity of the storage unit can be saved.

  In the scattered radiation correction apparatus having any one of the first to fourth configurations, the normalized ideal value calculation unit divides the measurement image into a plurality of regions each of which is an aggregate of a plurality of pixels, A luminance value by all X-rays representing the region (for example, a luminance value by all X-rays at the center position of the region) is obtained for each region, and the normalized ideal value is obtained by referring to the luminance value by all X-rays representing the region. It is good also as a structure (5th structure) to calculate. Thereby, the influence by the error of the luminance value of each pixel can be suppressed.

  In the scattered radiation correction apparatus of the fifth configuration, the normalized ideal value calculation unit corrects the normalized ideal value obtained by simulation with reference to a luminance value by all X-rays representing the region, and It is good also as a structure (sixth structure) which makes the said normalized ideal value after calculation a calculation result. Thereby, the scattered radiation correction adapted to the actual subject can be performed.

  In the scattered radiation correction apparatus having any one of the first to sixth configurations, the normalized ideal value is obtained when bone or metal on an X-ray path that passes through a region of interest of a subject is replaced with skin. It is good also as a structure (7th structure). Thereby, the structure of a human body can be considered.

  In order to achieve the above object, the scattered radiation correction method according to the present invention is a scattered radiation correction method for performing a scattered radiation correction on a measurement image obtained by X-ray imaging, and is based on all normalized X-rays. A normalized ideal value calculating step for calculating a normalized ideal value that is a luminance value for each pixel of the measurement image, and the normalized ideal value and the normalized luminance of the measurement image for each pixel of the measurement image The scattered radiation component calculation step for calculating the scattered radiation component of the measurement image from the value, and the scattering value calculated in the scattered radiation component calculation processing step from the luminance value of the pixel of the measurement image for each pixel of the measurement image A correction processing step for removing the line component.

  In order to achieve the above object, in the X-ray imaging apparatus according to the present invention, an X-ray irradiation unit that irradiates a subject with X-rays, and an X-ray detection unit that detects X-rays transmitted through the subject, A measurement image generation unit that generates a measurement image using a detection result of the X-ray detection unit, a projection image generation unit that generates a projection image by logarithmically converting the measurement image, and a scattered ray with respect to the measurement image The scattered radiation correction apparatus having any one of the first to seventh configurations that performs correction is provided.

  According to the present invention, an image with a reduced scattered radiation component can be directly derived from a measurement image obtained by X-ray imaging. In addition, the calculation time can be shortened.

It is a figure which shows the external appearance of the main-body part of the X-ray imaging apparatus which concerns on one Embodiment of this invention. It is a figure which shows the standard track | orbit of panoramic imaging mode. It is a figure which shows the track | orbit of local CT imaging mode. It is a figure which shows the center of an imaging | photography object site | part at the time of setting the center of an imaging | photography object site | part to the position of the front tooth, and an image reconstruction range in local CT imaging | photography mode. It is a figure which shows the center of an imaging | photography object site | part at the time of setting the center of an imaging | photography object site | part to the position of a left jaw, and an image reconstruction range in local CT imaging | photography mode. It is a figure which shows the center of an imaging | photography object site | part at the time of setting the center of an imaging | photography object site | part to the position of a right 2nd premolar in the local CT imaging | photography mode, and an image reconstruction range. It is a figure which shows the track | orbit of all-tooth CT imaging | photography mode. It is a figure which shows the track | orbit of all jaw CT imaging | photography mode. It is a figure which shows the structure of an image processing apparatus. It is a figure which shows the geometry of calculation of a Monte Carlo simulation. It is a figure which shows the combination of the column-shaped water phantom which assumed the jaw part, the head, and the neck. It is a graph which shows the calculation result of a Monte Carlo simulation. It is a graph which shows the calculation result of a Monte Carlo simulation. It is a graph which shows the calculation result of a Monte Carlo simulation. It is a graph which shows the calculation result of a Monte Carlo simulation.

  Embodiments of the present invention will be described below with reference to the drawings.

  First, the configuration of a main body 1 (hereinafter referred to as “main body 1 of an X-ray imaging apparatus”) of an X-ray imaging apparatus according to an embodiment of the present invention will be described with reference to FIG. 1A and 1B are views showing the appearance of the main body of the X-ray imaging apparatus 1. FIG. 1A is a top view, FIG. 1B is a front view, and FIG. 1C is a side view.

  A main body 1 of an X-ray imaging apparatus is a main body of an X-ray imaging apparatus for dental use or otolaryngology, and is erected in a vertical direction from a base 2 placed on a floor surface and the base 2. The lower pole 3, the upper pole 4 connected to the lower pole 3 so as to be slidable in the vertical direction, the fixed arm 5 fixed to the upper end of the upper pole 4, and the swivel connected to the fixed arm 5 so as to be rotatable An arm 6 and a head holding portion 7 that is fixed to the central portion of the upper pole 4 and holds the head of a human body including a subject (for example, a tooth) are provided. In the embodiment, the fixed arm 5 is fixed to the upper pole 4. For example, the fixed arm 5 is directly or directly on the wall or ceiling of the room where the main body 1 of the X-ray imaging apparatus is installed. The aspect attached through the adjustment mechanism which can adjust distance may be sufficient.

  The swivel arm 6 has an X-ray irradiation unit 8 that irradiates the subject with X-rays and an X-ray detection unit 9 that detects X-rays transmitted through the subject. In the present embodiment, as the X-ray detection unit 9, a flat panel detector is used in which conversion elements that generate electrical signals in accordance with irradiated X-rays are two-dimensionally arranged. Carbon is installed on the front of the flat panel detector.

  The imaging mode of the main body 1 of the X-ray imaging apparatus is not particularly limited, and examples include a panoramic imaging mode and a CT imaging mode. In the panoramic imaging mode, the swivel axis of the swivel arm 6 is perpendicular to the swivel axis (X direction, X direction, so that the X-ray irradiation unit 8 and the X-ray detection unit 9 draw a predetermined locus along the shape of the dental arch. The tomography is performed while moving the swivel arm 6 about the swivel axis. In the CT imaging mode, tomographic imaging of the target imaging region (image reconstruction range) is performed while rotating the swivel arm 6 around the target imaging region (image reconstruction range) of the head.

  Here, the panorama shooting mode will be described in more detail with reference to FIG. FIG. 2 shows a standard (adult) trajectory in the panoramic shooting mode. In the standard (adult) trajectory of the panoramic imaging mode, the X-ray irradiation unit 8 and the X-ray detection unit 9 draw a predetermined trajectory along the shape of the virtual dental arch 201 and the trajectory of the X-ray beam is an envelope. The turning arm 6 on which the X-ray irradiating unit 8 and the X-ray detecting unit 9 are arranged is moved from the imaging start position P1 to the imaging end position P2 along the trajectory shown in FIG. The turning angle of the turning arm 6 between the photographing start position P1 and the photographing end position P2 is about 220 degrees. The positions of the swing arm 6 shown in FIG. 2 excluding the shooting end position P2 are the shooting positions in the left half of the shooting area of the subject. X-rays emitted from the X-ray focal point 8A of the X-ray irradiation unit 8 are focused by an X-ray diaphragm 8B provided in the X-ray irradiation unit 8, and the X-ray beam width W on the X-ray detection unit 9 is reduced. Adjusted.

  In addition to the standard (adult) trajectory described above, the panoramic imaging mode preferably has a pediatric trajectory, an orthogonal trajectory, a temporomandibular joint trajectory, a maxillary sinus photographing trajectory, and the like. The trajectory for children is different from the standard trajectory in that the shape of the virtual dental arch 201 is reduced. The orthogonal trajectory is different from the standard trajectory in that the X-ray beam at each imaging position passes between the teeth of the patient dental arch 203. The temporomandibular joint imaging trajectory (side surface) is a swivel arm so that the X-ray irradiation unit 8 and the X-ray detection unit 9 draw a predetermined trajectory along the shape of both end portions of the virtual dental arch 201 (part where temporomandibular joint imaging is possible). The point 6 is moved is different from the standard trajectory. The temporomandibular joint imaging trajectory (front) differs from the standard trajectory in that the X-ray irradiation unit 8 and the X-ray detection unit 9 move the swivel arm 6 so as to draw a predetermined trajectory along the shape of the virtual line 204. . The maxillary sinus imaging trajectory is different from the standard trajectory in that the X-ray irradiation unit 8 and the X-ray detection unit 9 move the turning arm 6 so as to draw a predetermined trajectory along the shape of the virtual line 205.

  Subsequently, the CT imaging mode will be described in more detail with reference to FIGS. 3 to 8, the same parts as those in FIG. 2 are denoted by the same reference numerals.

  The local CT imaging mode is a CT imaging mode in which a specific region narrower than the entire upper and lower tooth regions in the tooth jaw region is imaged. The image reconstruction range in the local CT imaging mode is, for example, a cylindrical space region having a diameter of 51 mm and a height of 55 mm. FIG. 3 shows the trajectory in the local CT imaging mode. In the local CT imaging mode, as shown in FIG. 3, the swivel arm 6 is placed so that the center of the X-ray detection unit 9 is on the extension line of the line connecting the X-ray irradiation unit 8 and the swivel axis center 206 of the swivel arm 6. Shooting is performed at a plurality of shooting positions while turning. In the local CT imaging mode, the pivot axis center 206 of the pivot arm 6 is normally at a fixed position as shown in FIG. In FIG. 3, four shooting positions are illustrated, but this is only an example, and the shooting positions are not limited to the illustrated positions.

  In the local CT imaging mode, since the X-ray beam width W on the X-ray detection unit 9 is narrower than the all-tooth CT imaging mode and the full jaw CT imaging mode described later, even if the size of the X-ray detection unit 9 is small. It can be implemented.

  In the local CT imaging mode, the position of the turning axis center 206 of the turning arm 6 is changed according to where the center of the region to be imaged (region of interest) is set. Usually, as shown in FIG. Further, the position adjustment is performed so that the center of the imaging target region (region of interest) and the position of the turning axis center 206 of the turning arm 6 coincide with each other. The center of the region to be imaged (region of interest) in the local CT imaging mode can be arbitrarily set. In addition to the position setting shown in FIG. 3, for example, as shown in FIG. 4, the center 208 of the region to be imaged (region of interest) can be set to the position of the front tooth on the virtual dental arch 201, as shown in FIG. As shown in FIG. 6, the center 208 of the imaging target region (region of interest) can be set to the position of the left jaw on the virtual dental arch 201. As shown in FIG. The position of the right second premolar on the virtual dental arch 201 can also be set, and various other position settings are possible.

  The all-tooth CT imaging mode is a CT imaging mode in which the entire upper and lower tooth regions are to be imaged. The image reconstruction range in the all-tooth CT imaging mode is, for example, a cylindrical space region having a diameter of 97 mm and a height of 100 mm. FIG. 7 shows the trajectory in the full-tooth CT imaging mode. In the all-tooth CT imaging mode, as shown in FIG. 7, the swivel arm 6 is such that the center of the X-ray detection unit 9 is on the extended line of the line connecting the X-ray irradiation unit 8 and the swivel axis center 206 of the swivel arm 6. Shooting is performed at a plurality of shooting positions while turning. In the all-tooth CT imaging mode, the pivot axis center 206 of the pivot arm 6 is normally at a fixed position as shown in FIG. In FIG. 7, four shooting positions are illustrated, but this is only an example, and the shooting positions are not limited to the illustrated positions.

  Compared with the above-described local CT imaging mode, the full-tooth CT imaging mode has a wider imaging target and the X-ray beam width W on the X-ray detection unit 9 is wider, and therefore corresponds to the wide X-ray beam width W. The size of the X-ray detection unit 9 is required.

  The all jaw CT imaging mode is a CT imaging mode in which the entire range of the tooth jaw region is an imaging target. The image reconstruction range in the all jaw CT imaging mode is, for example, a cylindrical space region having a diameter of 161 mm and a height of 100 mm. FIG. 8 shows the trajectory in the full jaw CT imaging mode. In the full jaw CT imaging mode, as shown in FIG. 8, the swivel arm 6 is such that the center of the X-ray detection unit 9 is displaced from the extended line of the line connecting the X-ray irradiation unit 8 and the swivel axis center 206 of the swivel arm 6. Shooting is performed at a plurality of shooting positions while turning. In the all jaw CT imaging mode, the pivot axis center 206 of the pivot arm 6 is normally at a fixed position as shown in FIG. In FIG. 8, four shooting positions are illustrated, but this is merely an example, and the shooting positions are not limited to the illustrated positions.

  In the all jaw CT imaging mode, imaging is performed by shifting the center of the X-ray detection unit 9 from the extended line of the line connecting the X-ray irradiation unit 8 and the turning axis center 206 of the turning arm 6. The image reconstruction range 207 can be expanded as compared with the CT imaging mode. Accordingly, the entire range of the tooth and jaw region can be taken as an imaging target while suppressing an increase in the size of the X-ray detection unit 9.

  In the all-chin CT imaging mode, the X-ray detection unit 9 is sized up so that the X-ray beam width W on the X-ray detection unit 9 is larger than that shown in FIG. By using a cylindrical space region having a diameter of 230 mm and a height of 164 mm, it is possible to capture not only the entire region of the tooth and jaw region but also the entire region of the head and neck region.

  In the above-described panoramic imaging mode and CT imaging mode, the patient's dental arch 203 is present at the position assumed at the time of imaging (the position illustrated in FIGS. 2, 3, 7, and 8). You can shoot as you intended. As a method for easily realizing the alignment of the patient dental arch 203 to the assumed position, for example, a method using a light beam can be cited. Examples of the light beam include a midline light beam indicating the position of the midline of the head, a horizontal light beam indicating the position of the line connecting the lower edge of the orbit and the ear canal, and a tomogram indicating the position of the canine (reference position for tomography). There are reference line light beams and the like, and an output unit of these light beams may be provided in the X-ray imaging apparatus, and the patient may finely adjust the position of the head with reference to these light beams.

  Further, a cephalometric unit (not shown) installed at a position away from the swivel arm 6 may be used to perform photographing in the cephalometric photographing mode. The cephalometric unit includes a cephalometric X-ray detection unit for detecting X-rays transmitted through the subject and imaging the subject, and a head fixing unit for fixing the head. Cephalometric imaging is used for orthodontic diagnosis and the like, and imaging is performed using a head-specific X-ray imaging method (cephalometric imaging method). In cephalometric imaging, for example, the ear rods of the head fixing part are inserted and fixed in the left and right outer ear hole parts of the head, and X-rays are irradiated from the X-ray irradiation part 8 provided on the turning arm 6 to obtain the subject. X-rays that have passed through are detected by a Cefaro X-ray detector.

  An X-ray imaging apparatus according to an embodiment of the present invention includes an image processing apparatus 10 shown in FIG. 9 in addition to the main body 1 of the X-ray imaging apparatus.

  The image processing apparatus 10 includes a CPU 101 that controls the entire image processing apparatus 10 according to programs stored in the ROM 102 and the HDD 107, a ROM 102 that records fixed programs and data, a RAM 103 that provides a working memory, and an X-ray imaging. A communication interface unit 104 for communicating with a control unit (not shown) that is stored in the main unit 1 of the apparatus and controls each unit of the main unit 1 of the X-ray imaging apparatus, and temporarily stores image data. VRAM 105, display unit 106 for displaying an image based on image data stored in VRAM 105, an imaging control program for controlling X-ray imaging operation in cooperation with the control unit and CPU 101, and a reconstructed image. Image reconstruction processing program for generating, scattered radiation correction processing program for performing scattered radiation correction processing Various programs, set values of various parameters used in executing various programs, and includes a HDD107 for storing various data such as the reconstructed image data, a keyboard, an input unit 108 such as a pointing device.

  In the image processing apparatus 10, the communication method between the image processing apparatus 10 and the control unit may be wired communication, wireless communication, or communication combining wired and wireless. An example of the image processing apparatus 10 is a personal computer. In addition to image processing, the image processing apparatus 10 also performs remote operation and image display of the main body 1 of the X-ray imaging apparatus. Each program stored in the HDD 107 may be preinstalled in the image processing apparatus 10, distributed in a form stored in a storage medium such as an optical disk, and installed in the image processing apparatus 10. And installed in the image processing apparatus 10.

  When the scattered radiation correction processing program is executed, the image processing apparatus 10 functions as a scattered radiation correction apparatus. The scattered radiation correction process is executed by interrupting the image reconstruction process.

<< Theory underlying scattered radiation correction processing >>
Before explaining the content of the scattered radiation correction process, the theory which is the premise of the scattered radiation correction process will be described. The theory is a theory for calculating a luminance value by a scattered ray based on a luminance value of a pixel of a measurement image, and is constructed independently by the present inventor. The present inventor constructed the theory using the results of Monte Carlo simulation.

  Here, the geometry of the Monte Carlo simulation calculation is shown in FIG. FIG. 10A is a top view, and FIG. 10B is a side view. The subject 11 is disposed between the X-ray focal point 8A and the X-ray detector 9A. The X-ray detector 9A was a flat panel detector equipped with a scintillator and the like. A carbon 9B is installed in front of the X-ray detector 9A.

  In the calculation of the Monte Carlo simulation, the X-ray spectrum generated at the X-ray focal point 8A is emitted from the X-ray tube based on the specifications of the X-ray tube used in the actual imaging, and is X-rayed at the position of the X-ray detector 9A. A uniform X-ray flux collimated to be a sensitive region of the detector 9A was obtained.

  For the subject, a cylindrical water phantom 12A having a diameter of 15 cm assuming the jaw, a cylindrical water phantom 12B having a diameter of 18 cm assuming the head, and a column having a diameter of 13 cm assuming the neck are used in order to approach the living body. The water phantom 12C is combined as shown in the side view shown in FIG. 11 (a) and the top view shown in FIG. 11 (b), and a cylindrical cortical bone having a thickness of 2 mm is installed inside the water phantom 12A. A cylindrical cortical bone having a thickness of 1 mm was placed inside the phantom, and an elliptic cylindrical cervical vertebra portion 12D having a major axis of 4 cm, a minor axis of 3 cm, and a thickness of 3 mm was installed inside the water phantoms 12A and 12C. However, the cylindrical cortical bone and the cervical spine portion 12D were cut in the portion surrounded by the alternate long and short dash line shown in FIG.

  In the calculation of the Monte Carlo simulation, the X-ray attenuation and scattered rays caused by the metal are examined with the portion surrounded by the one-dot chain line shown in FIG. This is to estimate the effect of only the metal that does not depend on bone while maintaining the attenuation effect of scattered radiation from the surroundings due to the presence of bone.

  First, Monte Carlo simulation calculation is performed using a combination of the water phantoms shown in FIG. 11, in which a metal is not placed in a portion surrounded by a one-dot chain line shown in FIG. The calculation result is shown by a solid line in FIG. The horizontal axis of the graph shown in FIG. 12 is the ratio of the luminance value of all X-rays (direct rays and scattered rays) to the average luminance value in the white image, and the vertical axis of the graph shown in FIG. 12 is that of the X-ray detector 9A. The ratio of the scattered dose contained in the total X-ray dose to the total X-ray dose detected at the pixel.

  The ratio of the luminance value of all X-rays to the average luminance value in the white image increases as the thickness of the subject in the X-ray irradiation direction decreases. From the calculation result indicated by the solid line in FIG. 12, it can be seen that the ratio of the scattered dose to the total X-ray dose decreases as the thickness of the subject in the X-ray irradiation direction decreases. The calculation result shown by the solid line in FIG. 12 can be approximated by, for example, y = A / (x−C) + B. Assuming that A = 0.049556, B = 0.046501, and C = −0.16274, an approximate curve indicated by a broken line in FIG. 12 is obtained.

  Next, each calculation result of the Monte Carlo simulation in the following cases (1) to (3) is compared with the graphs shown in FIGS.

  The horizontal axis of the graph shown in FIG. 13 is the pixel number indicating the pixel position in the x-axis direction shown in FIG. 11, and the vertical axis of the graph shown in FIG. 13 is the ratio of the luminance value of all X-rays to the average luminance value in the white image. It is. Note that the larger the pixel number, the thicker the subject through which X-rays incident on the corresponding pixel of the X-ray detector 9A are transmitted. A curve C1 in FIG. 13 is a calculation result obtained in the case of (1) below. Similarly, a curve C2 in FIG. 13 is a calculation result obtained in the following case (2), and a curve C3 in FIG. 13 is a calculation result obtained in the following case (3).

  Further, the horizontal axis of the graph shown in FIG. 14 is a pixel number indicating the pixel position in the x-axis direction shown in FIG. 11, and the vertical axis of the graph shown in FIG. 14 is detected by a certain pixel of the X-ray detector 9A. The ratio of the scattered dose contained in the total X-ray dose to the total X-ray dose. A curve C11 in FIG. 14 is a calculation result obtained in the case of (1) below. Similarly, a curve C12 in FIG. 14 is a calculation result obtained in the following case (2), and a curve C13 in FIG. 14 is a calculation result obtained in the following case (3).

(1) A combination of the water phantoms shown in FIG. 11 and a subject in which no metal is placed in a portion surrounded by a dashed line shown in FIG. 11 (2) A combination of the water phantoms shown in FIG. 11, titanium having a width of 5 mm, a length of 8 cm, and a thickness of 3 mm was placed in a portion surrounded by a one-dot chain line shown in FIG. 11 so that the width direction is the z axis, the length direction is the y axis, and the thickness direction is the x axis. When the object is the subject and the X-ray irradiation direction is aligned with the x-axis (3) A combination of the water phantoms shown in FIG. 11, the portion surrounded by the alternate long and short dash line shown in FIG. When a 1 mm thick AuAgPd alloy is installed with the width direction aligned with the z-axis, the length direction aligned with the y-axis, and the thickness direction aligned with the x-axis, the X-ray irradiation direction aligned with the x-axis

  From FIG. 13 and FIG. 14, even though the luminance values of all X-rays detected by the pixels of the X-ray detector 9 </ b> A (corresponding to the luminance values of the pixels of the measurement image) are similar, It can be seen that the scattered dose varies depending on the type and thickness.

  For example, the condition that the average luminance value of the white image is 30,000 and the luminance value of all X-rays detected by the pixels of the X-ray detector 9A is 600, that is, the luminance value of all X-rays with respect to the average luminance value of the white image The pixel number that satisfies the condition that the ratio is 0.02 is 201 in the case of (2), and is 5 in the case of (3) (see FIG. 13). When the ratio of the scattered dose contained in the total X-ray dose detected with respect to the total X-ray dose detected by the pixels corresponding to these pixel numbers is obtained from FIG. 14, it is 0.491 and 0.987, respectively. To 295 and 592, respectively. Therefore, even if the luminance values of the pixels of the measurement image are the same 600, depending on what kind of material the X-ray has transmitted, that is, depending on the attenuation coefficient of the material, The scattered radiation component is different.

  11 is a combination of the water phantoms shown in FIG. 11, and the average luminance value in the white image obtained by performing the Monte Carlo simulation using the object in which the metal is not installed in the portion surrounded by the one-dot chain line shown in FIG. The ratio of the luminance values due to all X-rays differs depending on the thickness of the subject in the X-ray irradiation direction (see curve C1 in FIG. 13). Hereinafter, the luminance value obtained by the pixel of the X-ray detector 9A when bone or metal on the X-ray path that passes through the target region of the subject is replaced with the skin is referred to as an ideal value, and the average luminance value in the white image is obtained. The ratio of ideal values is called the ideal value ratio.

  While changing the type and thickness of the metal with the ideal value ratio fixed, the ratio of the luminance value of all X-rays to the average luminance value in the white image and the total X detected by a certain pixel of the X-ray detector 9A One graph shown in FIG. 15 is obtained by obtaining the relationship with the ratio of the scattered dose contained in the total X-ray dose to the dose. For example, each point of the white circle and white triangle in FIG. 15 is obtained from the white circle and white triangle in FIGS. 13 and 14 with the ideal value ratio fixed at 0.1. 15 is obtained by approximating the curve of y = A / (x−C) + B and obtaining the values of A, B, and C, respectively.

  Then, each graph shown in FIG. 15 is obtained by changing the ideal value ratio. A curve C21 in FIG. 15 is a graph with an ideal value ratio of 1, a curve C22 in FIG. 15 is a graph with an ideal value ratio of 0.3, and a curve C23 in FIG. 15 has an ideal value ratio of 0.1. A curve C24 in FIG. 15 is a graph with an ideal value ratio of 0.03.

  Here, each curve in FIG. 15 corresponds to a case where the projection angle of X-ray imaging is fixed to a certain value. The relationship between the ratio of the luminance value of all X-rays to the average luminance value in the white image and the ratio of the scattered dose included in the total X-ray dose to the total X-ray dose detected by a certain pixel of the X-ray detector 9A. The curve indicating varies depending on the projection angle of X-ray imaging. This is because, in the case of X-ray imaging at a projection angle such that the dentition is located on the side closer to the X-ray focal point 8A, the metal is X-ray focal point within the portion surrounded by the one-dot chain line shown in FIG. When the X-ray imaging is performed at a projection angle such that the dentition is located on the side close to 8A and the dentition is located on the side far from the X-ray focal point 8A, the inside of the portion surrounded by the alternate long and short dash line shown in FIG. The metal installation position in the subject differs depending on the projection angle of X-ray photography, such as placing metal on the far side of the X-ray focal point 8A, and the difference in the metal installation position detects a direct line that has passed through the metal. This is because it affects the scattered dose detected by the pixels of the X-ray detector 9A.

As described above, the method for calculating the luminance value based on the scattered radiation based on the luminance value of the pixel of the measurement image is summarized as follows.
[1] The ideal value ratio is calculated for all pixels of the measurement image.
[2] The scattered ray components of all pixels of the measurement image are calculated from the ideal value ratio, the ratio of the luminance value of the measurement image to the average luminance value in the white image, and the projection angle of the X-ray imaging.

  By the way, in CT imaging, X-rays are cut in the vicinity of the X-ray focal point 8A so that the X-ray irradiation range matches the sensitive region of the X-ray detector 9A in order to suppress excessive exposure. Therefore, the pixel located at the end of the X-ray detector 9A is less affected by scattered rays from the subject than the pixel located at the center of the X-ray detector 9A. As a result, the scattered radiation component is reduced at the pixel located at the end.

  This decreasing tendency is stronger as the pixel at the end of the X-ray detector 9A, but continuously decreases with increasing distance from the end of the X-ray detector 9A, and the X-ray detection is 120 pixels or more away from the end of the X-ray detector 9A. In the center of the vessel 9A, it becomes almost zero. However, in a pixel where the ratio of the luminance value of all X-rays to the average luminance value in the white image is high, the scattered light from the subject is originally small because the thickness of the subject in the X-ray irradiation direction is thin. For this reason, the smaller the ideal value ratio, the smaller the decrease in the scattered radiation component due to the pixel being located at the end of the X-ray detector 9A. Therefore, when the pixel is located toward the end of the X-ray detector 9A and the ideal value ratio is low, it is desirable to obtain the scattered radiation component in consideration of the decrease in the scattered radiation component described above.

<< Contents of scattered radiation correction process >>
The scattered radiation correction processing is divided into ideal value ratio calculation processing, scattered radiation component calculation processing, and correction processing, and is executed in the order of ideal value ratio calculation processing, scattered radiation component calculation processing, and correction processing.

<Ideal value ratio calculation processing>
In the ideal value ratio calculation process, for example, the image processing apparatus 10 collects measurement images of 512 × 512 pixels for each 16 × 16 pixels, divides them into a total of 1024 regions of 32 in the horizontal direction and 32 in the vertical direction, After calculating the ideal value ratio of each region, the ideal value ratio of each pixel is calculated. The reason for such a calculation procedure is that although the luminance value of the pixel gradually changes according to the thickness of the subject in the X-ray irradiation direction, the error of the luminance value of each pixel is determined by examining each pixel. This is because it is very difficult to determine whether the decrease or increase in the luminance value is due to the change in the thickness of the subject in the X-ray irradiation direction or due to an error. The pixel size and the number of areas to be divided are merely examples, and are not limited to the values shown above.

  In addition, the image processing apparatus 10 first calculates 32 curves (horizontal regions in the horizontal direction) obtained by moving the regions in the horizontal direction with respect to the ideal value ratio of the center positions of 32 regions in the horizontal direction × 32 regions in the vertical direction. 11 is a combination of the measurement image and the water phantom shown in FIG. 11, which is a curve obtained by setting y, the ratio of the luminance value of all the X-rays at the center position to the average luminance value in the white image, as shown in FIG. Thirty-two curves (vertical lengths) obtained by calculating each from the data obtained by performing Monte Carlo simulation calculation on a subject in which no metal is placed in the portion surrounded by the alternate long and short dash line, and then moving the region in the vertical direction 11 is a combination of the measured image and the water phantom shown in FIG. 11 those not including the metal on the subject in a portion surrounded by a one-dot chain line shown in (a) performing the ideal value ratio calculation processing in step of calculating from each of the data obtained by performing the calculation of the Monte Carlo simulation. In this embodiment, the ideal value ratio of each center position is adopted as the ideal value ratio representing each region, but the ideal value ratio representing each region is not limited to this, for example, In each region, an average value of ideal value ratios of the entire regions may be obtained, and each average value may be adopted as an ideal value ratio representing each region.

  When the region is moved in the direction from high to low in the horizontal direction for the measurement image, the ratio of the luminance value of all X-rays at the center position to the average luminance value in the white image enters the bone portion. In the area of only the skin up to this point, it basically changes according to the curve of y = A / (x−C) + B, and when entering the bone part from the area of only the skin, it deviates greatly from the curve. Further, the curve (the range of x is limited to the part of the skin only until entering the bone part) does not necessarily match the curve C1 in FIG.

  Therefore, the image processing apparatus 10 changes the curve C1 in FIG. 13 to x so that the curve (the range of x is limited to only the portion of the skin until entering the bone portion) and the curve C1 in FIG. The ideal value ratio of the center position of the region is calculated based on the curve C1 in FIG. 13 after the expansion / contraction process or reduction process in the axial direction and the expansion / contraction process or reduction process in the x-axis direction.

  However, when X-rays are irradiated from the vicinity of the back of the subject and when irradiated from the vicinity of the front, the range of only the skin becomes extremely narrow, so the luminance value is high in the horizontal direction for the measurement image. The accuracy of calculating the ideal value ratio obtained when the region is moved in the direction from the lower side to the lower side is deteriorated.

  On the other hand, in the longitudinal direction, there is a skin-only part in a wide range such as the jaw and neck. Therefore, at the projection angle corresponding to the case where X-rays are irradiated from the vicinity of the back of the subject or the vicinity of the front, there is only a skin portion in the calculation of the ideal value ratio when the region is moved in the horizontal direction. When the value falls below the range, the calculation is switched to the ideal value ratio calculation when the region is moved in the vertical direction, and the ideal value ratio is calculated when the region is moved in the vertical direction for, for example, five columns from the end. Like that. The ideal value ratio is calculated from the skin-only part by assuming that the value obtained from the calculation result of the ideal value ratio when the region is moved in the vertical direction is the ideal value ratio obtained from the skin-only part. Since it is possible to calculate the ideal value ratio when the region is moved in the horizontal direction, which is difficult to perform, the calculation of the ideal value ratio when the region is moved in the horizontal direction is resumed.

  Since the calculation of the ideal value ratio when the region is moved in the horizontal direction is performed in each row, the degree of expansion / contraction processing or reduction processing in the x-axis direction for the curve C1 in FIG. For this reason, the change in the ideal value ratio in the vertical direction is often not smooth but uneven. Therefore, it is necessary to apply a treatment for smoothing the change in the ideal value ratio in the vertical direction. However, if the same treatment as that in the horizontal direction is performed, the change in the ideal value ratio in the horizontal direction becomes uneven again. In this embodiment, the least square method is used.

  When the least square method is applied to y = A / (x−C) + B, the value becomes very complicated, and the values of A, B, and C cannot be obtained. Therefore, when the logarithm z = logy is taken with respect to the ideal value ratio, z is proportional to x locally, so the least square method is applied to z = Ax + B, and then z is returned to y. The ideal value ratio was calculated. In the present embodiment, the range to which the least square method is applied is divided into four, and the least square method is applied separately while making the boundary smooth for each. Note that the number of divisions in the range to which the least squares method is applied may be other than four.

  Further, the image processing apparatus 10 sets each region as a target region of interest in order to make the change in the ideal value ratio of each region calculated in the above-described procedure more smooth, and averages the ideal value ratios of the regions around the target region. Smoothing processing is performed with the value as the ideal value ratio of the region of interest.

  At the end of the ideal value ratio calculation process, the image processing apparatus 10 calculates the value of the ideal value ratio of each pixel based on the ideal value ratio of the center position of each region. For example, when the pixel that is the target of calculating the ideal value ratio is located on a line segment that connects the center positions of adjacent regions, the ideal value ratio changes linearly on the line segment, If the pixel whose ideal value ratio is to be calculated is not located on the line segment that connects the center positions of adjacent areas, the ideal value ratio of the four center positions surrounding the pixel is used. The ideal value ratio of the first point located on the line segment that connects the center positions of a pair of regions adjacent in the horizontal direction with the same horizontal coordinate is linear on the line segment. Next, the second point of the second point located on the line segment that connects the center positions of another set of regions adjacent to each other in the horizontal direction and having the same horizontal coordinate as that of the pixel is calculated. The ideal value ratio is calculated assuming that the ideal value ratio varies linearly on the line segment. If the ideal value ratio is linearly changed on the line segment connecting the first point and the second point, the ideal value ratio value of each pixel can be changed. Although it can be calculated, the ideal value ratio value of each pixel may be obtained by other methods. For example, the horizontal direction may be replaced with the vertical direction in the above method.

<Scattered ray component calculation processing>
For each pixel of the measurement image, the image processing apparatus 10 calculates the ideal value ratio calculated by the above-described ideal value ratio calculation processing, the ratio of the luminance value of the measurement image to the average luminance value in the white image, and the projection angle of X-ray imaging Then, the scattered radiation component of the measurement image is calculated using the curve data shown in FIG. The curve data shown in FIG. 15 is stored in the HDD 107 in the form of a data table in which the values of constants A, B, and C of y = A / (x−C) + B are associated with the ideal value ratio and the projection angle. Has been.

  Here, when the pixel is located toward the end of the X-ray detector 9A and the ideal value ratio is low, the image processing apparatus 10 performs a correction process to reduce the scattered radiation component calculated as described above. You may do it.

<Correction process>
The image processing apparatus 10 performs a correction process for removing the scattered radiation component calculated by the above-described scattered radiation component calculation process from the luminance value of the pixel of the measurement image for each pixel of the measurement image. Thereby, the image which reduced the scattered ray component can be derived | led-out directly from the measurement image obtained by X-ray imaging.

  The brightness value of the measurement image also reflects scattered rays from the subject and the flat panel detector. Therefore, even if the subject is a metal that does not pass X-rays at all, such as an AuAgPd alloy with a high Au content, the measurement image must be displayed. The luminance value appears in the. In the process of removing the metal artifact, if the scattered radiation component is still included, the metal portion cannot be extracted well, so that the metal artifact cannot be sufficiently removed. However, it is expected that metal artifacts can be removed if the transmitted X-rays of these metal parts can be accurately calculated by removing the scattered radiation component.

  Further, in general, CT values for dental use cannot calculate an accurate CT value. This is considered to be mainly caused by scattered rays from the subject and the flat panel detector. However, if the scattered radiation component can be removed from the measurement image, only the luminance value by the transmitted X-rays obtained by the X-ray attenuation of the subject can be obtained as data. Therefore, it can be expected that a more accurate CT value can be calculated.

  Although one embodiment of the present invention has been described above, the scope of the present invention is not limited to this, and various modifications can be made without departing from the spirit of the invention.

  For example, the ratio of the luminance value of all X-rays to the average luminance value in the white image and the X-ray detector 9A even if the projection angle of X-ray imaging changes due to the limitation of the imaging range of X-ray imaging. When the curve indicating the relationship with the ratio of the scattered dose included in the total X-ray dose to the total X-ray dose detected at a certain pixel of the pixel hardly changes, the image processing apparatus 10 performs measurement in the scattered ray component calculation process. For each pixel of the image, the curve data shown in FIG. 15 is obtained from only two of the ideal value ratio calculated in the ideal value ratio calculation process described above and the ratio of the luminance value of the measurement image to the average luminance value in the white image. It may be used to calculate the scattered radiation component of the measurement image.

  In the X-ray imaging apparatus for dentistry or otolaryngology described in the embodiment described above, the subject is limited to the human head. However, the present invention is not limited to the case where the subject is a human head. If the subject is not a human head, the subject model used in the Monte Carlo simulation may be a model suitable for the actual subject, instead of the model shown in FIG.

DESCRIPTION OF SYMBOLS 1 Body part of arm type X-ray imaging apparatus according to one embodiment of the present invention 2 Base 3 Lower pole 4 Upper pole 5 Fixed arm 6 Turning arm 7 Head holding part 8 X-ray irradiation part 8A X-ray focus 8B X-ray aperture DESCRIPTION OF SYMBOLS 9 X-ray detection part 9A X-ray detector 9B Carbon 10 Image processing apparatus 11 Subject 12A-12C Water phantom 12D Cervical vertebra part 101 CPU
102 ROM
103 RAM
104 Communication interface 105 VRAM
106 Display 107 HDD
DESCRIPTION OF SYMBOLS 108 Input part 201 Virtual dental arch 202 Envelope virtual dental arch 203 Patient dental arch 204, 205 Virtual line 206 Rotating-arm center of rotation axis 207 Image reconstruction range P1 Imaging start position P2 Imaging end position W On X-ray detection unit X-ray beam width at

Claims (9)

A scattered radiation correction apparatus that performs scattered radiation correction on a measurement image obtained by X-ray imaging,
A normalized ideal value calculator that calculates a normalized ideal value, which is a luminance value of all normalized X-rays, for each pixel of the measurement image;
For each pixel of the measurement image, a scattered radiation component calculation unit that calculates a scattered radiation component of the measurement image from the normalized ideal value and the normalized luminance value of the measurement image;
A scattered radiation correction apparatus comprising: a correction process for removing a scattered radiation component calculated by the scattered radiation component calculation processing unit from a luminance value of each pixel of the measurement image for each pixel of the measurement image.   The scattered radiation component calculation unit, for each pixel of the measurement image, scatters the measurement image from the normalized ideal value, the normalized luminance value of the measurement image, and the projection angle of the X-ray imaging. The scattered radiation correction apparatus according to claim 1, which calculates a line component. A storage unit for storing the relationship between the normalized luminance value of all X-rays and the ratio of the scattered dose included in the total X-ray dose to the total X-ray dose;
The scattered radiation correction apparatus according to claim 1, wherein the scattered radiation component calculation unit calculates a scattered radiation component of the measurement image using the relationship stored in the storage unit.   The scattered radiation correction apparatus according to claim 3, wherein the storage unit stores a constant value used in a function indicating the relationship in the form of a data table associated with the normalized ideal value.   The normalized ideal value calculation unit divides the measurement image into a plurality of regions each of which is an aggregate of a plurality of pixels, obtains a luminance value for all regions representing the region for each region, The scattered radiation correction apparatus according to claim 1, wherein the normalized ideal value is calculated with reference to luminance values of representative all X-rays.   The normalized ideal value calculation unit corrects the normalized ideal value obtained by simulation with reference to luminance values of all X-rays representing the region, and uses the corrected idealized ideal value as a calculation result. The scattered radiation correction apparatus according to claim 5.   The scattered radiation correction apparatus according to claim 1, wherein the normalized ideal value is obtained when bone or metal on an X-ray path that passes through a region of interest of a subject is replaced with skin. A scattered radiation correction method for performing scattered radiation correction on a measurement image obtained by X-ray photography,
A normalized ideal value calculating step of calculating a normalized ideal value, which is a luminance value of all normalized X-rays, for each pixel of the measurement image;
For each pixel of the measurement image, a scattered radiation component calculation step for calculating a scattered radiation component of the measurement image from the normalized ideal value and the normalized luminance value of the measurement image;
A scattered radiation correction method, comprising: a correction processing step for removing the scattered radiation component calculated in the scattered radiation component calculation processing step from the luminance value of the pixel of the measurement image for each pixel of the measurement image. An X-ray irradiation unit that irradiates the subject with X-rays;
An X-ray detector that detects X-rays transmitted through the subject;
A measurement image generation unit that generates a measurement image using the detection result of the X-ray detection unit;
A projection image generation unit that logarithmically converts the measurement image to generate a projection image;
An X-ray imaging apparatus comprising: the scattered radiation correction apparatus according to claim 1, wherein the scattered radiation correction is performed on the measurement image.