JP6142289B2 - Scattered ray correction apparatus, scattered ray correction method, and X-ray imaging apparatus - Google Patents

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. For example, Patent Document 1 proposes a scattered radiation component reduction method that can directly derive an image obtained by reducing the scattered radiation component from an image obtained by X-ray imaging.

Japanese Patent No. 4703221

  However, the scattered radiation component reduction method proposed in Patent Document 1 performs scattered radiation correction when the pixel value of a measurement image obtained by X-ray imaging is equal to or greater than a threshold value (the pixel value of a projected image is equal to or less than the threshold value). This is equivalent to performing the scattered radiation correction only when the pixel value of the measurement image obtained by X-ray imaging is equal to or less than the threshold value (the pixel value of the projection image is equal to or greater than the threshold value). The effect of line correction becomes discontinuous and unnatural.

  In view of the above situation, the present invention is capable of directly deriving an image obtained by reducing scattered radiation components from an image obtained by X-ray imaging, and a scattered radiation correction apparatus in which the effect of scattered radiation correction does not become unnatural. An object of the present invention is to provide 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 one aspect of the present invention, for a projection image obtained by X-ray imaging, x is used to calculate x of the projection image using a correction function y = f (x). A scattered radiation correction apparatus that performs scattered radiation correction using a pixel value y as a pixel value after scattered radiation correction, and the correction function is a logistic function y = α / [1 + exp [−β (x−γ )]] Is a configuration (first configuration) that is an inverse function of a function obtained by parallel translation.

  By making the correction function y = f (x) an inverse function of a function obtained by translating a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve, A projection image (projection image including a scattered radiation component) obtained by X-ray imaging can be converted by a correction function y = f (x) to obtain a projection image with a reduced scattered radiation component. That is, an image with a reduced scattered radiation component can be directly derived from an image obtained by X-ray imaging. Further, since the scattered radiation correction is continuously performed on the pixel values of the projection image obtained by X-ray imaging, the effect of the scattered radiation correction does not become unnatural. In dental CT imaging, since the subject is limited to the human head (particularly the jaw), the function form of the correction function is suitable for scattered radiation correction, and the scattered radiation component can be appropriately reduced.

  In order to achieve the above object, in the scattered radiation correction apparatus according to another aspect of the present invention, x is measured using a correction function y = f (x) for a measurement image obtained by X-ray imaging. A scattered radiation correction apparatus that performs scattered radiation correction using the measured image pixel value as y and the scattered radiation corrected pixel value, and the correction function is a logistic function y = α / [1 + exp [−β ( x-γ)]] is a function (second structure) which is a function obtained by parallel movement.

  X-ray imaging by making the correction function y = f (x) a function obtained by translating a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve The measurement image (measurement image including the scattered radiation component) obtained in (1) can be converted by the correction function y = f (x) to obtain a measurement image with a reduced scattered radiation component. That is, an image with a reduced scattered radiation component can be directly derived from an image obtained by X-ray imaging. In addition, since the scattered radiation correction is continuously performed on the pixel values of the measurement image obtained by X-ray imaging, the effect of the scattered radiation correction does not become unnatural. In dental CT imaging, since the subject is limited to the human head (particularly the jaw), the function form of the correction function is suitable for scattered radiation correction, and the scattered radiation component can be appropriately reduced.

  In the scattered radiation correction apparatus having the first or second configuration, the parameter value may be adjusted according to the projection angle of the X-ray imaging (third configuration).

  Since the total amount of scattered radiation detected by the X-ray detector during X-ray imaging varies depending on the projection angle of X-ray imaging, the function form of the correction function suitable for scattered radiation correction is the projection angle of X-ray imaging. It depends on. According to the third configuration, even if the projection angle of X-ray imaging changes, a function form of a correction function suitable for scattered radiation correction can be obtained each time, and scattered radiation components can be appropriately reduced.

  In the scattered radiation correction apparatus having the third configuration, a storage unit that stores each value of the parameter corresponding to each of the plurality of projection angles in association with the projection angle, and the parameter stored in the storage unit It is good also as a structure (4th structure) provided with the interpolation part which interpolates the value of the said parameter corresponding to the said projection angle which does not respond | correspond to any of these values.

  According to such a configuration, since the number of parameter values stored in the storage unit can be reduced, it is possible to save the trouble of storing each parameter value in the storage unit in association with the projection angle.

  In the scattered radiation correction apparatus having the fourth configuration, the interpolation unit may use Catmull-Rom spline interpolation (fifth configuration).

  According to such a configuration, the calculation process of interpolation can be simplified, and the burden on the interpolation unit can be reduced.

  In order to achieve the above object, in the scattered radiation correction method according to one aspect of the present invention, x is projected onto a projection image obtained by X-ray imaging using a correction function y = f (x). A scattered radiation correction method is performed in which scattered radiation correction is performed using the pixel value of the image as y and the pixel value after the scattered radiation correction, and the correction function is a logistic function y = α / [1 + exp [−β (x -γ)]] is an inverse function of the function obtained by translation.

  In addition, in the scattered radiation correction method according to another aspect of the present invention to achieve the above object, x is calculated using a correction function y = f (x) for a measurement image obtained by X-ray imaging. It is a scattered radiation correction method in which scattered radiation correction is performed using the pixel value of the measurement image as y and the pixel value after the scattered radiation correction, and the correction function is a logistic function y = α / [1 + exp [−β ( x-γ)]] is a function obtained by translating.

  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 logarithmically converts the measurement image to generate a projection image, and the projection image or the measurement image. It is set as the structure provided with the scattered radiation correction apparatus of the said any one of the said 1st-5th structure which performs scattered radiation correction | amendment with respect to it.

  According to one aspect of the present invention, a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve is obtained by translating a projection image obtained by X-ray imaging. Since the scattered radiation correction is performed using an inverse function of the obtained function, an image with a scattered radiation component reduced can be directly derived from an image obtained by X-ray imaging.

  According to another aspect of the present invention, a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve is translated with respect to a measurement image obtained by X-ray imaging. Since the scattered radiation correction is performed using the obtained function, an image in which the scattered radiation component is reduced can be directly derived from an image obtained by X-ray imaging.

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 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 each example of a measurement image and a projection image. It is a figure which shows the example of a division | segmentation of a projection image. It is a graph which shows the example of a relationship between the pixel value of the projection image containing a scattered radiation component, and the pixel value of the projection image which does not contain a scattered radiation component. 13 is a graph showing an example in which a correction function is fitted to the set of points shown in FIG. 12. It is a graph which shows an example of the relationship between a projection angle and the value of the parameter of a correction function. It is a figure which shows the geometry of calculation of a Monte Carlo simulation. It is a side view 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.

  An image output from the X-ray detection unit 9 and received by the image processing apparatus 10 is called a measurement image. Each pixel value of the measurement image can be regarded as representing a value approximately proportional to the number of X-ray quanta that has reached each detection element of the X-ray detection unit 9.

  The reconstruction algorithm used in the image reconstruction processing program reconstructs the cross section using the line integral of the X-ray line attenuation coefficient instead of the number of X-ray quanta. For this reason, in the image reconstruction process, it is necessary to convert the measurement image indicating the distribution of the number of X-ray quanta into an image (projection image) indicating the distribution of the line integral of the line attenuation coefficient. Hereinafter, the conversion (logarithmic conversion) from the measurement image to the projection image will be described in detail.

The pixel label is i, the pixel value of the measurement image when there is no subject is I ₀ (i), the pixel value of the measurement image when the subject to be reconstructed this time is placed is I (i), and the X-ray When the path is s and the linear attenuation coefficient at the position s is μ (s), the following equation (1) is established. This equation (1) is an equation representing how X-rays are attenuated.

  While the left side of equation (1) is the pixel value of the measurement image, the content of the exponent function on the right side is the pixel value of the projection image obtained by inverting the sign. Therefore, by solving the equation (1) regarding the contents of the exponential function, a conversion equation for conversion from the measurement image to the projection image (logarithmic conversion) can be obtained. Thus, the logarithmic conversion formula is as shown in the following formula (2).

  Here, an example of the measurement image is shown in FIG. 10A, and a projection image obtained by logarithmically converting the measurement image of FIG. 10A is shown in FIG.

The image processing apparatus 10 performs scattered ray correction on the projection image using the correction function y = f (x). Here, the correction function y = f (x) is an inverse function of a function obtained by translating a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve. . Furthermore, in this embodiment, the logistic functions α, β, and γ are set to α = 2a, β = 2 / a, and γ = 0. Therefore, in this embodiment, the correction function y = f (x) is expressed by the following equation (3), and a is a parameter for finely adjusting the function form of the correction function y = f (x).

  By using such a correction function y = f (x), a projection image (projection image including a scattered radiation component) obtained by X-ray imaging is converted by the correction function y = f (x), and the scattered radiation component is converted. The projected image can be reduced. That is, it is possible to directly derive a projection image in which scattered radiation components are reduced from a projection image obtained by X-ray imaging. Further, since the scattered radiation correction is continuously performed on the pixel values of the projection image obtained by X-ray imaging, the effect of the scattered radiation correction does not become unnatural.

  Hereinafter, a procedure for setting the value of the parameter a will be described. The procedure for setting the value of the parameter a may be performed by the image processing apparatus 10, but may be performed by another image processing apparatus.

  First, two projection images that differ only in the presence or absence of scattered radiation components obtained by X-ray imaging of the same subject at the same projection angle are prepared by some means. Specific examples of the means will be described later. Note that the projection image that does not include the scattered radiation component is ideally a projection image that does not include the scattered radiation component at all, but the scattered radiation component may remain to some extent.

  Next, each of the projection image including the scattered radiation component and the projection image not including the scattered radiation component is divided into a plurality of regions by the same division method. An example of this division is shown in FIG. In the division example shown in FIG. 11, the projection image is equally divided. For example, the dentition portion may be finely divided, and the portions other than the dentition may be divided roughly. Further, the size of the area when dividing may be arbitrary, but it is desirable to adjust in consideration of the influence of noise included in the projection image.

  Next, an average pixel value in each region is calculated for each of the projection image including the scattered radiation component and the projection image not including the scattered radiation component.

  Next, the correspondence relationship between the average pixel values in the same region of the projection image including the scattered radiation component and the projection image not including the scattered radiation component is plotted as one point in the graph. By performing this plot for all regions, a set of points indicating the correspondence relationship between the average pixel values in the same region of the projection image including the scattered radiation component and the projection image not including the scattered radiation component is obtained. An example of this point aggregate is shown in FIG.

  Then, the correction function represented by the above equation (3) is added to a set of points indicating the correspondence relationship between the average pixel values in the same region of the projection image including the scattered radiation component and the projection image not including the scattered radiation component. Adjust the value of parameter a to fit. The correction function represented by the above equation (3) is fitted to a set of points indicating the correspondence relationship between the average pixel values in the same region of the projection image including the scattered radiation component and the projection image not including the scattered radiation component. An example is shown in FIG. White squares in FIG. 13 indicate points on the correction function expressed by the above equation (3).

  Here, the total amount of scattered radiation detected by the X-ray detector 9 during X-ray imaging varies depending on the projection angle of X-ray imaging. For this reason, the function form of the correction function that fits a set of points indicating the correspondence relationship between the average pixel values in the same region of the projection image including the scattered radiation component and the projection image not including the scattered radiation component is an X-ray imaging function. It depends on the projection angle. Therefore, the value of parameter a is adjusted according to the projection angle of X-ray imaging.

  In the present embodiment, when performing X-ray imaging 510 times by equally dividing 360 degrees that circulate around the subject, the above formula is applied only to a projection image having a specific projection angle among 510 projection angles. Fitting of the correction function represented by (3) is executed, and the value of the parameter a corresponding to the remaining projection angle is indirectly derived by interpolation. Therefore, the HDD 107 stores each value of the parameter a corresponding to each specific projection angle in association with the projection angle, and stores an interpolation program for interpolating the value of the parameter a. For example, the relationship between the projection angle after the interpolation program is executed and the value of the parameter a is as shown in FIG. The interval between the specific projection angles may be arbitrary, but it is desirable to adjust in consideration of a trade-off between the accuracy of the value of parameter a obtained by interpolation and the degree to which the value of parameter a after interpolation vibrates.

  As the interpolation to be applied, an interpolation in which the transition of the value of the parameter a after the interpolation is smooth and the value of the parameter a after the interpolation does not vibrate as much as possible is desirable. be able to. Catmull-Rom spline interpolation is particularly desirable because it is easy to compute as follows. For example, Newton interpolation, Lagrangian interpolation, or the like may be applied when a slight vibration is allowed with respect to the value of the parameter a after interpolation or when the number of given points is small.

Value of parameter a obtained by Catmull-Rom spline interpolation
Is the label of the projection image corresponding to the projection angle of interest (the order when 510 projection images are arranged in the projection angle order) z, and the value of the parameter a by fitting the correction function represented by the above equation (3) Z _i (i = 1,2,3,4) of the labels of the projected image obtained directly from the label z of the projected image of interest, one before, one before, one after, Is expressed by the following equation (4).

  Finally, an example of means for preparing two projection images that differ only in the presence or absence of scattered radiation components obtained by X-ray imaging of the same subject at the same projection angle will be described.

<< Theory as a premise of means for preparing two projection images that differ only in the presence or absence of scattered radiation components >>
Before explaining the contents of means for preparing two projection images that differ only in the presence or absence of scattered radiation components, the theory that is the premise thereof will be explained. The theory is a theory for calculating a luminance value by scattered rays based on a luminance value of a pixel of a measurement image. The theory is constructed using simulation results of Monte Carlo simulation.

  Here, the geometry of the Monte Carlo simulation calculation is shown in FIG. FIG. 15A is a top view, and FIG. 15B 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. 16A and the top view shown in FIG. 16B, and a cylindrical cortical bone having a thickness of 2 mm is placed 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 alternate long and short dash 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. 16 and a subject 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. 17 is the ratio of the luminance values 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. 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. 17, 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 indicated by the solid line in FIG. 17 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. 17 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. 18 and 19.

  The horizontal axis of the graph shown in FIG. 18 is the pixel number indicating the pixel position in the x-axis direction shown in FIG. 16, and the vertical axis of the graph shown in FIG. 18 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. 18 is a calculation result obtained in the case of (1) below. Similarly, a curve C2 in FIG. 18 is a calculation result obtained in the following case (2), and a curve C3 in FIG. 18 is a calculation result obtained in the following case (3).

  Further, the horizontal axis of the graph shown in FIG. 19 is a pixel number indicating the pixel position in the x-axis direction shown in FIG. 16, and the vertical axis of the graph shown in FIG. 19 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. 19 is a calculation result obtained in the case of (1) below. Similarly, a curve C12 in FIG. 19 is a calculation result obtained in the following case (2), and a curve C13 in FIG. 19 is a calculation result obtained in the following case (3).

(1) When the water phantom combination shown in FIG. 16 is a subject in which no metal is placed in the portion surrounded by the alternate long and short dash line shown in FIG. 16 (2) The water phantom combination shown in FIG. 16, 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. When the object is a subject and the X-ray irradiation direction is aligned with the x-axis (3) A combination of the water phantoms shown in FIG. 16, 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. 18 and FIG. 19, even if the luminance values (corresponding to the luminance values of the pixels in the measurement image) of all X-rays detected by the pixels of the X-ray detector 9A 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) above, and is 5 in the case of (3) above (see FIG. 18). 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. 19, they are 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.

  16 is a combination of the water phantoms shown in FIG. 16, 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 alternate long and short dash 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. 18). Hereinafter, the luminance value obtained by the pixel of the X-ray detector 9A when the bone or metal on the X-ray path that passes through the region of interest 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. 20 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. 20 is obtained from the white circle and white triangle in FIGS. 18 and 19 with the ideal value ratio fixed at 0.1. 20 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. 20 is obtained by changing the ideal value ratio. A curve C21 in FIG. 20 is a graph with an ideal value ratio of 1, a curve C22 in FIG. 20 is a graph with an ideal value ratio of 0.3, and a curve C23 in FIG. 20 has an ideal value ratio of 0.1. A curve C24 in FIG. 20 is a graph with an ideal value ratio of 0.03.

  Here, each curve in FIG. 20 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 close 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 X-ray detection is performed at 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 means for preparing two projection images that differ only in the presence or absence of scattered radiation components >>
The means for preparing two projection images that differ only in the presence or absence of scattered radiation components is a computer that executes ideal value ratio calculation processing, scattered radiation component calculation processing, correction processing, and conversion processing. The computer executes processing in the order of ideal value ratio calculation processing, scattered radiation component calculation processing, correction processing, and conversion processing.

<Ideal value ratio calculation processing>
In the ideal value ratio calculation process, for example, the computer collects a measurement image of 512 × 512 pixels every 16 × 16 pixels, divides it into a total of 1024 regions of 32 in the horizontal direction and 32 in the vertical direction. After calculating the ideal value ratio, 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 computer first calculates 32 curves obtained by moving the region in the horizontal direction (x in the horizontal direction x, white in the ideal value ratio of the center position of 32 regions in the horizontal direction × 32 regions in the vertical direction). 16 is a combination of the measurement image and the water phantom shown in FIG. 16 and represented by a one-dot chain line shown in FIG. 16 (a). 32 curves (longitudinal area) obtained by calculating each from the data obtained by performing Monte Carlo simulation calculation on the object where no metal is placed in the enclosed part, and then moving the area in the vertical direction Is a combination of the measurement image and the water phantom shown in FIG. 16, where x is a curve obtained by setting y as the ratio of the luminance value of all X-rays at the center position to the average luminance value in the white image. Performing ideal value ratio calculation processing in step of calculating from each of the data obtained by performing the calculation of Monte Carlo simulations to those not including the metal part surrounded by a chain line shown in (a) in the subject. 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 computer changes the curve C1 in FIG. 18 in the x-axis direction so that the curve (the range of x is limited to only the part 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.

  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.

  Furthermore, in order to make the change in the ideal value ratio of each region calculated by the above-described procedure smoother, the computer targets each region as the target region and pays attention to the average value of the ideal value ratios of the regions around the target region. Smoothing processing is performed to obtain an ideal value ratio of the region.

  At the end of the ideal value ratio calculation process, the computer calculates the ideal value ratio value 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>
The computer, for each pixel of the measurement image, from the ideal value ratio calculated by the above-described ideal value ratio calculation process, 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, The scattered radiation component of the measurement image is calculated using the curve data shown in FIG. The curve data shown in FIG. 20 is stored in the computer 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. Stored in the department.

  Here, when the pixel is located toward the end of the X-ray detector 9A and the ideal value ratio is low, the computer performs correction processing to reduce the scattered radiation component calculated as described above. Also good.

<Correction process>
The computer performs, for each pixel of the measurement image, correction processing for removing the scattered radiation component calculated by the above-described scattered radiation component calculation processing from the luminance value of the pixel of the measurement image.

<Conversion processing>
The computer performs logarithmic conversion using the above equation (2) on each of the measurement image obtained by X-ray imaging and the measurement image after the correction process, and converts the result into a projection image. Thereby, two projection images that differ only in the presence or absence of scattered radiation components can be prepared.

  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, a phantom that imitates a human head is generated on a computer, and the scattering obtained by X-ray imaging of the same subject at the same projection angle by simulation using EGS (Electron-Gamma Shower) etc. Two projected images that differ only in the presence or absence of a line component may be prepared.

  Also, for example, in the above-described embodiment, the inverse function of the function obtained by translating the logistic function y = α / [1 + exp [−β (x−γ)]] representing the logistic curve with respect to the measurement image. However, since the projected image is an image obtained by logarithmically converting the measurement image, the logistic curve representing the logistic curve for the measurement image is used instead of the scattered ray correction performed in the above-described embodiment. Scattered ray correction may be performed using a function obtained by translating the function y = α / [1 + exp [−β (x−γ)]].

  Further, for example, if the function form of the correction function hardly changes even if the projection angle of X-ray imaging changes due to the limitation of the imaging range of X-ray imaging, the value of parameter a is set to one. It may be fixed.

  In the above-described embodiment, the HDD 107 stores only each value of the parameter a corresponding to a specific projection angle. However, each value of the parameter a corresponding to all projection angles at which X-ray imaging is performed is stored. May be. In addition, if the relationship between the projection angle and the value of the parameter a can be approximated by a function expression, the value of the parameter a corresponding to the projection angle may be obtained by calculation using the function expression.

  In the X-ray imaging apparatus for dental or otolaryngology described in the above-described embodiment, the subject is limited to the human head (particularly the jaw), and the subject has a similar shape every time. For this reason, it is considered that the method of distribution of scattered radiation components included in an image obtained by X-ray imaging may be considered to have small individual differences. Therefore, when the subject is a human head (particularly the jaw), it is expected that the scattered radiation component can be appropriately reduced by performing the scattered radiation correction with the same correction function and the same parameter a value transition. it can. However, the present invention is not limited to the case where the subject is a human head (particularly the jaw). For example, the transition of the correction function fixed coefficient or the parameter value included in the correction function according to the projection angle may be changed according to shooting conditions other than the projection angle including the type of subject.

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 9A X-ray detector 9B Carbon 9 X-ray detection unit 10 Image processing device 11 Subject 12A to 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 (8)

Scattered ray correction is performed on a projection image obtained by X-ray imaging using a correction function y = f (x), with x being a pixel value of the projected image and y being a pixel value after the scattered ray correction. A device,
Scattering ray correction, wherein the correction function is an inverse function of a function obtained by translating a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve apparatus. Scattered ray correction is performed on a measurement image obtained by X-ray imaging using a correction function y = f (x), with x being a pixel value of the measurement image and y being a pixel value after the scattered ray correction. A device,
The scattered radiation correction apparatus, wherein the correction function is a function obtained by translating a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve.   The scattered radiation correction apparatus according to claim 1, wherein the parameter value is adjusted according to a projection angle of the X-ray imaging. A storage unit that stores each value of the parameter corresponding to each of the plurality of projection angles in association with the projection angle;
The scattered radiation correction apparatus according to claim 3, further comprising: an interpolation unit that interpolates a value of the parameter corresponding to the projection angle that does not correspond to any of the parameter values stored in the storage unit.   The scattered radiation correction apparatus according to claim 4, wherein the interpolation unit uses Catmull-Rom spline interpolation. Scattered ray correction is performed on a projection image obtained by X-ray imaging using a correction function y = f (x), with x being a pixel value of the projected image and y being a pixel value after the scattered ray correction. A method,
Scattering ray correction, wherein the correction function is an inverse function of a function obtained by translating a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve Method. Scattered ray correction is performed on a measurement image obtained by X-ray imaging using a correction function y = f (x), with x being a pixel value of the measurement image and y being a pixel value after the scattered ray correction. A method,
The scattered radiation correction method, wherein the correction function is a function obtained by translating a logistic function y = α / [1 + exp [−β (x−γ)]] representing a logistic curve. 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 projection image or the measurement image.