Arithmetic processing device, radiation imaging device, x-ray CT device, and radiotherapy system

Abstract

The present invention comprises: an X-ray tube 5; an X-ray detector 4; and a scattered ray correction unit 13 that calculates a first scattered ray from X-ray projection data captured by an X-ray imaging device using a point-symmetric scattering kernel, and calculates a second scattered ray from the first scattered ray on the basis of the thickness of a subject to be imaged. The present invention calculates corrected X-ray projection data from the X-ray projection data and the second scattered ray, and reconstructs CT. The present invention thereby provides an arithmetic processing device, a radiation imaging device, an X-ray CT device, and a radiotherapy system with which it is possible to acquire X-ray projection data in which the degradation in image quality due to scattered rays is reduced without increasing calculation time compared to the prior arts.

Description

Processing unit, radiation imaging device, X-ray CT scanner, and radiation therapy system

[0001] The present invention relates to a computing device, a radiation imaging device, an X-ray CT device, and a radiation therapy system.

[0002] Patent Document 1 relates to a method, computer program product, and system for estimating scattered radiation in a radiographic projection of an object using a symmetric kernel, wherein the radiographic projection is generated by a two-dimensional imaging device irradiated by a radiation source spaced therefrom to provide space for the object, the imaging device measures incident radiation at a plurality of pixels at corresponding positions on a two-dimensional surface, and the radiographic projection consists of the two-dimensional surface and a plurality of radiation values ​​corresponding to a plurality of pixel positions on the imaging device, each radiation value including a primary radiation dose representing radiation reaching the corresponding pixel along a direct path from the radiation source and a scattered radiation dose representing other radiation reaching the corresponding pixel.

[0003] Non-patent document 1 describes using a convolutional filtering method to estimate the scattering distribution in images acquired by a digital subtraction angiography imaging system.

[0004] U.S. Patent Registration No. 8326011

[0005] Love LA, Kruger RA. Scatter estimation for a digital radiographic system using convolution filtering. Med Phys. 1987;14(2):178-185. doi:10.1118 / 1.596126

[0006] Radiation therapy is a treatment method that aims to eradicate tumors or alleviate pain by irradiating tumors in the patient's body with radiation such as X-rays, electron beams, proton beams, and heavy ion beams to deliver a sufficient dose.

[0007] In general, in radiation therapy, radiation (hereinafter also abbreviated as beam) accelerated by an accelerator system consisting of linear accelerators, synchrotrons, cyclotrons, etc. is transported to an irradiation nozzle and irradiated onto the tumor.

[0008] In radiation therapy, which uses radiation to treat cancer and tumors, it is necessary to understand the positional relationship between the target and dangerous organs within the patient's body on the day of treatment. For this purpose, a cone-beam CT (Cone Beam Computed Tomography) device is used, which takes images using an X-ray imaging device installed on the rotating support device of the radiation therapy apparatus.

[0009] In cone-beam CT imaging, an X-ray beam is first directed from an X-ray source towards the subject. A portion of the X-ray beam that has attenuated within the subject is detected by an X-ray detector such as an FPD (Flat Panel Detector) and stored as X-ray projection data.

[0010] Typically, the X-ray source and X-ray detector are designed to rotate around the subject, acquiring X-ray projection data of the subject from multiple angles. A three-dimensional tomographic image (cone-beam CT image) of the subject is reconstructed by applying back projection processing to the X-ray projection data obtained from multiple angles.

[0011] Cone-beam CT images used in radiation therapy are acquired to confirm the location of tumors and normal organs, therefore, the image quality of cone-beam CT images is crucial. However, it is known that the image quality of cone-beam CT images can deteriorate due to various factors.

[0012] Scattered radiation is one of the factors that significantly affects the image quality of cone-beam CT images.

[0013] Compared to the linear detectors used in diagnostic CT, the FPDs used in cone-beam CT imaging have a larger detection area, resulting in a greater amount of scattered X-rays (scattered radiation) from within the subject being mixed into the X-ray projection data.

[0014] When scattered radiation is mixed into X-ray projection data, errors occur during the reconstruction process of cone-beam CT images, resulting in artifacts in the cone-beam CT images. The presence of these artifacts reduces the contrast between body tissues and impairs the visibility of organ locations, thus requiring countermeasures.

[0015] To improve the image quality of cone-beam CT images, it is important to estimate the scattered radiation contained in the X-ray projection data and correct it by subtracting it from the X-ray projection data.

[0016] As a method for estimating scattered radiation, a method called the SKS method (Scatter Kernel Superposition method), which involves convolving a symmetrical scattering kernel onto X-ray projection data, is known, as described in Non-Patent Literature 1. In this method, it is assumed that X-rays incident on each point on the X-ray projection data are scattered point-symmetrically in a Gaussian distribution, and the scattered radiation distribution is estimated by convolving a scattering kernel representing the spread of scattered radiation onto the X-ray projection data. In reality, when X-rays are irradiated onto a plate-shaped object at a point, the scattered radiation is distributed point-symmetrically around the point of incidence. Therefore, by using the method described in Non-Patent Literature 1, the scattered radiation distribution could be estimated with relatively good accuracy when the shape of the object was close to that of a plate.

[0017] On the other hand, when X-rays are applied to a triangular prism-shaped object, the object's shape becomes asymmetrical with respect to the X-ray incidence point, resulting in an asymmetrical distribution of scattered radiation. In such cases where the object's shape is asymmetrical with respect to the X-ray incidence point, the accuracy of scattered radiation estimation using this method decreases. In particular, since the human body, which is the target of radiation therapy, is not plate-shaped but closer to an ellipse, the method described in Non-Patent Document 1 suffers from reduced accuracy in estimating scattered radiation, requiring different countermeasures.

[0018] Furthermore, as a method for estimating scattered radiation while considering the asymmetry of the scattering radiation's spread, there is a technique described in Patent Document 1. Patent Document 1 describes a method for estimating scattered radiation by convolving an asymmetric scattering kernel into X-ray projection data. The asymmetry of the scattering kernel is formulated by the difference in subject thickness between two points.

[0019] The method described in Patent Document 1 uses an asymmetric scattering kernel to accurately estimate scattered radiation even for objects with an elliptical shape. However, it has the drawback of increasing computation time due to the increased number of convolution calculations required when calculating scattered radiation.

[0020] The present invention provides a processing unit, a radiography apparatus, an X-ray CT apparatus, and a radiotherapy system that can acquire X-ray projection data with reduced image quality degradation due to scattered radiation without increasing computation time compared to conventional methods.

[0021] The present invention includes multiple means for solving the above problems, but one example is to calculate a first scattered ray from X-ray projection data captured by an X-ray imaging device using a point-symmetric scattering kernel, and then calculate a second scattered ray from the first scattered ray based on the thickness of the subject being imaged.

[0022] According to the present invention, it is possible to obtain X-ray projection data with reduced image quality degradation due to scattered radiation without increasing computation time compared to conventional methods. Other problems, configurations, and effects will be clarified by the following description of the embodiments.

[0023] This figure shows an overview of the radiotherapy system of the embodiment. This is a schematic diagram showing the configuration of the cone-beam CT imaging device of the embodiment. This is a flowchart showing the flow of cone-beam CT reconstruction using the cone-beam CT imaging device of the embodiment. This is a flowchart showing the details of the scattered radiation processing flow in the processing unit of the embodiment. This figure shows an overview of the first scattered radiation estimation process by the processing unit of the comparative example. This figure shows the X-ray imaging system of the cylindrical phantom evaluated in the preliminary study. This figure shows the X-ray imaging system of the cylindrical phantom evaluated in the preliminary study. This figure shows the X-ray projection data, scattered radiation, and estimated scattered radiation of the cylindrical phantom evaluated in the preliminary study. This figure shows the X-ray projection data, scattered radiation, and estimated scattered radiation of the cylindrical phantom evaluated in the preliminary study. This figure shows an overview of the second scattered radiation distribution estimation process in the processing unit of the embodiment. This figure shows a comparison of cone-beam CT reconstruction images of the cylindrical phantom due to differences in scattered radiation correction in the processing unit of the embodiment. This is a flowchart showing the details of the scattered radiation processing flow in the processing unit of the modified example 1. This is a schematic diagram showing the configuration of the cone-beam CT imaging device of the modified example 2.

[0024] Embodiments of the arithmetic processing unit, radiation imaging device, X-ray CT device, and radiation therapy system of the present invention will be described with reference to Figures 1 to 13. In the drawings used herein, the same or corresponding components are denoted by the same or similar reference numerals, and repeated descriptions of these components may be omitted.

[0025] First, the overall configuration of the radiation therapy system will be explained using Figure 1. Figure 1 is a diagram showing the schematic configuration of the radiation therapy system according to this embodiment.

[0026] The radiotherapy system 150 shown in Figure 1 mainly comprises a bed 2 that supports the patient (subject 1), a rotating support device 3 surrounding the bed 2, an X-ray tube 5, an X-ray detector 4, a control computer 110, a therapeutic radiation control device 120, a therapeutic radiation generator 130, a therapeutic radiation irradiation device 140, and the like.

[0027] The therapeutic radiation generator 130 is composed of equipment that generates various types of radiation, such as X-rays, proton beams, and heavy ion beams such as deuterium and carbon. Its configuration varies depending on the type of radiation to be generated, and various known configurations can be adopted.

[0028] In configurations that irradiate with proton beams or heavy ion beams, the therapeutic radiation generator 130 includes, for example, an ion source, a linac, and a synchrotron. The synchrotron includes a deflection magnet, a quadrupole magnet, a radio frequency accelerator, a radio frequency emitter, and an emitter deflector.

[0029] The ion source is connected to a linac, which is connected to a synchrotron. Particles generated from the ion source are pre-accelerated by the linac and then injected into the synchrotron. The particle beam, further accelerated in the synchrotron, is then emitted into the transport system.

[0030] The transport system includes multiple deflection electromagnets and a quadrupole electromagnet (not shown) and connects the synchrotron to the therapeutic radiation irradiation device 140. Furthermore, a portion of the therapeutic radiation generator 130 (the transport system) and the therapeutic radiation irradiation device 140 are mounted on a cylindrical rotating support device 3 and can rotate together with the rotating support device 3. The particle beam emitted from the synchrotron passes through the therapeutic radiation generator 130, is focused by the quadrupole electromagnet, and then its direction is changed by the deflection electromagnet before it enters the therapeutic radiation irradiation device 140.

[0031] The therapeutic radiation irradiation device 140 is equipped with two scanning electromagnets, a dose monitor, and a position monitor. The scanning electromagnets are positioned perpendicular to each other and generate a magnetic field by excitation current, which can deflect the particle beam so that it reaches a desired position in a plane perpendicular to the beam axis at the target location. The dose monitor measures the amount of irradiated particle beam. The position monitor can detect the position through which the particle beam has passed. The particle beam that has passed through the therapeutic radiation irradiation device 140 reaches the target within the subject 1. When treating patients with cancer or the like, subject 1 represents the patient, and the target represents a tumor or the like.

[0032] The bed on which the object 1 is placed is called the bed 2. The bed 2 can move in three mutually perpendicular axial directions based on instructions from the therapeutic radiation control device 120, and can further rotate about each axis. By these movements and rotations, the position of the object 1 including the target can be moved to a desired position.

[0033] [[ID=……]] The therapeutic radiation control device 120 is a device that controls the irradiation and stop of particle beams. In order to control devices such as the therapeutic radiation generation device 130 and the therapeutic radiation irradiation device 140, it is electrically connected to the therapeutic radiation generation device 130, the therapeutic radiation irradiation device 140, the bed 2, the console, etc.

[0034] Next, an X-ray CT apparatus is provided which includes an X-ray tube 5, an X-ray detector 4, and a scatter correction unit 13, calculates X-ray projection data and corrected X-projection data from second scatter radiation, and reconstructs a CT. Also, a calculation processing device (scatter correction unit 13) that calculates first scatter radiation from X-ray projection data taken by an X-ray imaging device using a point-symmetric scatter kernel and calculates second scatter radiation from the first scatter radiation based on the thickness of the object to be imaged will be described in detail with reference to FIG. 2. FIG. 2 shows an example of the cone beam CT imaging apparatus 100 and the attached control computer 110 of the present embodiment.

[0035] As shown in FIG. 2, the cone beam CT imaging apparatus 100 of the present embodiment includes a bed 2 that supports the object 1, a rotary support device 3 that surrounds the bed 2, one or more pairs of X-ray imaging devices attached to the rotary support device 3 so as to rotate integrally with the rotary support device 3, a rotation device (not shown) that rotates the rotary support device 3, and a control computer 110.

[0036] Among these, the X-ray imaging device has a pair of X-ray tubes 5 that perform X-ray imaging of the object 1 on the bed 2, an X-ray detector 4, and an X-ray aperture 6 (X-ray collimator). Also, the X-ray CT apparatus is substantially constituted by the parts of the cone beam CT imaging apparatus 100 and the control computer 110 excluding the scatter correction unit 13 and the control unit 10, and the radiation imaging apparatus is substantially constituted by the scatter correction unit 13, the control unit 10, the X-ray tube 5, and the X-ray detector 4.

[0037] The X-ray imaging apparatus is composed of one or more X-ray tubes 5, X-ray detectors 4, and X-ray shutters 6.

[0038] The X-ray tube 5 generates X-rays by receiving the application of a high voltage and the supply of a filament current from a high voltage generation unit (not shown). The irradiation time interval of the X-rays to the subject 1 is, for example, 10 times per second. The high voltage generation unit applies a high voltage to the X-ray tube 5 and supplies a tube current according to the imaging parameters determined by the imaging parameter setting unit 15 in the control computer 110.

[0039] An X-ray shutter 6 is attached to the X-ray irradiation port side of the X-ray tube 5. The X-ray shutter 6 limits the irradiation field of the X-rays generated from the X-ray tube 5. Specifically, the X-ray shutter 6 movably supports a plurality of shutter blades made of a material (such as lead) that shields X-rays. By adjusting the positions of the plurality of shutter blades, the size and shape of the X-ray irradiation field change. The X-ray shutter 6 moves the position of the shutter blades in response to the supply of a drive signal from the control unit 10.

[0040] The X-ray detector 4 detects the X-rays generated from the X-ray tube 5 and transmitted through the subject 1, and generates a current signal corresponding to the intensity of the detected X-rays. A data collection circuit (not shown) is connected to the X-ray detector 4 to collect the current signal output from the X-ray detector 4. The data collection circuit amplifies the collected current signal and digitally converts the amplified current signal to generate X-ray projection data, which is a digital signal. The X-ray projection data is transferred to the control unit 10 of the control computer 110 and stored in the storage unit 12 together with the time series information.

[0041] By irradiating X-rays from the X-ray tube 5 while the rotary support device 3 rotates and generating X-ray projection data with the X-ray detector 4, the X-ray projection data necessary for reconstructing the cone beam CT image is acquired.

[0042] The control computer 110 includes a control unit 10, a display unit 11, a storage unit 12, a scatter correction unit 13, a reconstruction unit 14, and an imaging parameter setting unit 15.

[0043] The control unit 10 controls the operation of the X-ray imaging apparatus and the rotation support device 3. The display unit 11 displays the X-ray projection data generated by the X-ray imaging apparatus and the cone-beam CT image created by the reconstruction unit 14 on the screen. The storage unit 12 stores various information. The scattered radiation correction unit 13 estimates the scattered radiation distribution contained in the X-ray projection data acquired by the X-ray imaging apparatus and corrects it by subtracting it from the X-ray projection data. The reconstruction unit 14 reconstructs the cone-beam CT image from the X-ray projection data. The imaging parameter setting unit 15 sets the imaging parameters of the X-ray tube 5 and the X-ray detector 4.

[0044] The control computer 110 and the therapeutic radiation control device 120 described above may each have a central processing unit (CPU) and memory connected to this CPU, or they may be configured as a single computer, and are not particularly limited in any way.

[0045] Furthermore, the control processes for the actions to be performed may be combined into a single program, divided into multiple programs, or a combination of these.

[0046] Some or all of the programs contained within each device may be implemented using dedicated hardware, or they may be modularized. Furthermore, various programs may be installed on each device via a program distribution server or external storage media, or existing devices may be updated.

[0047] Furthermore, each device may be an independent device connected by a wired or wireless network, or two or more devices may be integrated into a single unit.

[0048] Figure 3 is a flowchart showing an example of the cone-beam CT image reconstruction process in this embodiment.

[0049] As shown in Figure 3, when processing begins, the imaging parameter setting unit 15 first sets the imaging parameters (step S201). Specifically, the imaging parameters refer to the tube voltage and tube current applied to the X-ray tube 5, the exposure time, the number of exposures per unit time, and the imaging angle.

[0050] Next, the scattered radiation correction unit 13 reads out the scattered radiation correction parameters that have been previously stored in the storage unit 12 (step S202).

[0051] Next, the control unit 10 operates the rotation support device 3, X-ray detector 4, X-ray tube 5, X-ray diaphragm 6, etc., based on the imaging parameters set in step S201 to acquire X-ray projection data (step S203).

[0052] Subsequently, the scattered radiation correction unit 13 applies scattered radiation correction processing to the acquired X-ray projection data and stores the corrected X-ray projection data in the storage unit 12 (step S204).

[0053] Next, the reconstruction unit 14 reads the corrected X-ray projection data stored in the storage unit 12, reconstructs the cone-beam CT image by back-projecting the corrected X-ray projection data, and displays it on the display unit 11 (step S205).

[0054] Next, we will explain in detail the scattered radiation correction process in step S204 of Figure 3, which is a characteristic process in this embodiment, using Figure 4. Figure 4 is a flowchart showing an example of the scattered radiation correction process in this embodiment.

[0055] As shown in Figure 4, when the scattered radiation correction process (step S204) is started, the scattered radiation correction unit 13 first reads the X-ray projection data from the storage unit 12 (step S301).

[0056] Subsequently, the scattered radiation correction unit 13 calculates the subject thickness distribution τ(x,y) based on the following equation (1) (step S302).

[0057]

[0058] Here, x and y are positions on the X-ray projection data, I ｐ (x, y) is X-ray projection data, I ０ (x, y) is the X-ray projection data acquired when subject 1 is not present, called an air image, and μ is the linear attenuation coefficient of the X-rays, which ranges from 0.15 to 0.02 (mm) depending on the energy of the X-ray tube 5. －１ It takes the value of ).

[0059] In this way, the subject thickness can be calculated from X-ray projection data. Note that the subject thickness distribution τ(x,y) can also be determined from cone-beam CT images, treatment planning CT images for each imaging angle, or the standard (average) thickness of subject 1, in addition to being calculated from X-ray projection data.

[0060] Next, the scattered radiation correction unit 13 uses the following equation (2) based on the SKS method to determine the first scattered radiation distribution I ｓ １ｓｔ Calculate (x, y). In this way, the first scattered ray can be calculated based on a point-symmetric scattering kernel and X-ray projection data.

[0061]

[0062] Here, g(x-x', y-y') is a symmetric scattering kernel. A two-dimensional Gaussian distribution can be used for the scattering kernel. Figure 5 shows an image of calculating estimated scattered rays (synonymous with first scattered rays) using the SKS method, which is a comparative example.

[0063] As mentioned above, the accuracy of scattered radiation decreases with the SKS method for subjects with an elliptical shape. Our preliminary studies revealed that the SKS method tends to overestimate the scattered radiation distribution within the subject in many cases. Figures 6 (top view) and 7 (side view) show an example of an X-ray imaging system evaluated by our preliminary studies. In the systems shown in Figures 6 and 7, X-rays emitted from the X-ray tube 5 pass through the cylindrical phantom 60, which is the subject, and are incident on the X-ray detector 4.

[0064] Figures 8 and 9 show examples of X-ray projection data and ground truth scattered radiation evaluated using this system, estimated scattered radiation calculated by the SKS method, and the X-direction profile of the scattered radiation.

[0065] As shown in Figures 8 and 9, the ground truth scattered radiation shows a distribution that is convex downwards toward the center of the cylindrical phantom 60, which is the subject of the study. However, the estimated scattered radiation calculated by the SKS method, which is a comparative example, shows a distribution that is convex upwards toward the center of the cylindrical phantom 60. This is thought to be because the cylindrical phantom 60 is thicker in the center than at the edges, so the attenuation of scattered radiation is greater near the center, and the ground truth scattered radiation is convex downwards toward the center.

[0066] On the other hand, the SKS method uses a symmetric scattering kernel, which means it cannot take into account the large attenuation of scattered radiation in the center of the cylindrical phantom 60, and is thought to overestimate scattered radiation, especially where the subject 1 is thick.

[0067] While conducting such evaluations on various systems, the inventors discovered that the ground truth scattered radiation can be reproduced by reducing, or more preferably by determining, the estimated scattered radiation calculated by the SKS method based on the object thickness.

[0068] Therefore, in this embodiment, the first scattered ray distribution calculated by the SKS method is attenuated based on the object thickness to obtain the second scattered ray distribution I ｓ ２ｎｄ Calculate (x, y) (step S304). Specifically, calculate using the following formula (3).

[0069]

[0070] Thus, the second scattered ray can be calculated based on a function including one or more attenuation adjustment parameters, the object thickness, and the first scattered ray. Here, f(τ) represents the attenuation effect of the first scattered ray distribution based on the object thickness.

[0071] Figure 10 shows an image illustrating the calculation of the second scattered ray distribution. Since f(τ) represents the effect of attenuation of the first scattered ray within the subject, it is desirable that it has a value range of 0 to 1, and the function f(τ) including the attenuation adjustment parameter can take values ​​in the range of 0 to 1. However, it is also possible to have a value of 1 or greater to weaken the effect of attenuation of the first scattered ray.

[0072] The functional form of f(τ) can be determined empirically, for example, equations (4) through (6) below are used.

[0073]

[0074]

[0075]

[0076] Here, a, b, and c in equations (4) through (6) are damping adjustment parameters that adjust the damping effect and are determined empirically. Also, τｍａｘ This is the maximum thickness of the subject, and empirically, 100 to 1000 mm is used.

[0077] In equation (6), the first scattered rays are attenuated linearly as the thickness of the subject increases, whereas in equations (4) and (5), the effect of attenuating the first scattered rays can be gradually weakened as the thickness of the subject increases.

[0078] It is desirable to appropriately select which of equations (4) to (6) to use as f(τ) depending on the energy of the X-ray tube 5 and the distance between the X-ray tube 5 and the X-ray detector 4. Note that the functional form of f(τ) is not limited to these.

[0079] Furthermore, since scattered radiation contains only low-frequency components, the subject thickness distribution τ(x,y) may be pre-blurred using a Gaussian distribution before calculating f(τ). By pre-blurring τ(x,y) in this way, abrupt changes in estimated scattered radiation at the subject boundary are prevented, and scattered radiation outside the subject can also be estimated with high accuracy. As shown in Figure 8, the scattered radiation distribution calculated by this invention in the region where the subject exists agrees well with the ground truth scattered radiation.

[0080] Subsequently, the estimated direct X-ray distribution is calculated by subtracting the second scattered radiation distribution calculated in step S304 from the X-ray projection data (step S305). Once the processes from steps S301 to S305 have been performed for all the X-ray projection data (step S306), the scattered radiation estimation process (step S204) is terminated.

[0081] Figure 11 shows a comparison between the SKS method used for comparison and the method of this embodiment. Figure 11 is an example in which X-ray projection data was acquired from 360 degrees on a cylindrical phantom and reconstructed as a cone-beam CT image.

[0082] As shown in Figure 11, when reconstructed using X-ray projection data that does not contain scattered radiation, the center of the cylindrical phantom is uniform, whereas in the case where scattered radiation is not corrected, the CT value decreases towards the center of the cylindrical phantom.

[0083] Furthermore, when scattered radiation correction is performed using the SKS method, which is used for comparison, the scattered radiation is excessively corrected within the subject, resulting in an increase in the CT value towards the center of the cylindrical phantom.

[0084] On the other hand, it can be seen that by correcting scattered radiation using the method of this embodiment, excessive decreases or increases in CT values ​​can be suppressed, and a cone-beam CT image close to that of a state without scattered radiation can be created.

[0085] Furthermore, when the average value of the CT value error within a slice was evaluated using a cone-beam CT image without scattered radiation as the correct image, the average value of the uncorrected CT value error was "-74.8 HU", the average value of the CT value error with the conventional technology was "57.7 HU", and the average value of the CT value error with the present invention was "14.5 HU". The present invention has the smallest average value of the CT value error, and image quality close to that of a cone-beam CT image without scattered radiation has been obtained.

[0086] Furthermore, as can be seen from equation (3), the present invention requires only one convolution calculation, which is fewer than the method using the asymmetric scattering kernel described in Non-Patent Document 1, thus reducing the calculation time. In particular, since the cone-beam CT reconstruction process is performed while the patient is in the treatment room, it is desirable that the correction process be completed in the shortest possible time, and therefore a very favorable result can be obtained.

[0087] Next, a modified example 1 of the present invention will be described using Figure 12. For brevity of explanation, only the differences from the embodiment will be described.

[0088] Modification 1 evaluates the scattered radiation correction effect of the reconstructed image and optimizes the attenuation adjustment parameters. Specifically, by sequentially evaluating the image quality of the reconstructed cone-beam CT image while changing the attenuation adjustment parameters, the attenuation adjustment parameters are optimized, and a higher quality cone-beam CT image can be obtained.

[0089] In the first modification, when the scattered ray correction process (step S204) starts, as shown in FIG. 12, the scattered ray correction unit 13 sets an initial value of the attenuation adjustment parameter (step S901). As the initial value of the attenuation adjustment parameter, a value previously stored in the storage unit 12 can be used. Alternatively, the user can set an arbitrary value through the display unit 11.

[0090] Subsequently, the scattered ray correction unit 13 reads out the X-ray projection data stored in the storage unit 12 (step S902). Then, based on the formula (1), the object thickness distribution is calculated (step S903). Then, based on the formula (2), the first scattered ray distribution is calculated (step S905), and further, based on the formula (3), the second scattered ray distribution is calculated (step S906), and the estimated direct X-ray distribution is calculated (step S906).

[0091] Subsequently, after repeating the processes from step S902 to step S906 for all the X-ray projection data (step S907), a cone beam CT image is reconstructed using the estimated direct X-ray obtained in step S906 (step S908).

[0092] At the time of reconstruction in this step S908, depending on the initial value of the attenuation adjustment parameter, the scattered rays may be corrected excessively or insufficiently. Therefore, a correction effect index representing the effect of the scattered ray correction is calculated (step S909). Various correction effect indexes can be considered, and for example, the mean square error obtained by the following formula (7) can be used.

[0093]

[0094] Here, δ(f(x, y, z, a)) in the formula (7) is represented by the following formula (8).

[0095]

[0096] Here, a in the formulas (7) and (8) is the attenuation adjustment parameter, V is the evaluation region, N is the number of voxels included in the evaluation region, f(x, y, z, a) is the cone beam CT image obtained in step S908, f ｔａｒｇｅｔ is the target CT value, f ｔａｒｇｅｔ ｍｉｎ and f ｔａｒｇｅｔｍａｘ These are the minimum and maximum values ​​of the CT value being evaluated.

[0097] If scattered radiation is properly corrected, dips and excessive increases in CT values ​​are suppressed, and relatively uniform CT values ​​are obtained. Therefore, the correction effect can be evaluated by whether the soft tissue region, which makes up the majority of the human body, becomes uniform. If the soft tissue region is the target of evaluation, for example, f ｔａｒｇｅｔ = 0HU, f ｔａｒｇｅｔ ｍｉｎ = -300HU, f ｔａｒｇｅｔ ｍａｘ =300HU and similar models can be used.

[0098] It should be noted that this correction effect index is just one example and is not limited to it. For example, if scattered radiation is properly corrected, the uniformity of CT values ​​within the subject will improve, so the difference in CT values ​​between the edges and the center of subject 1 can also be considered as a correction effect index. Furthermore, the uniformity of the air region outside subject 1 may also be included as a correction effect index.

[0099] Returning to Figure 12, it is then determined whether the correction effect index has converged (step S910). If it has not converged, the process returns to step S901, updates the damping adjustment parameters, and repeats steps S901 to S909 again.

[0100] When using the correction effect index of equation (7), a smaller correction effect index suggests that the scattered radiation is being corrected appropriately. Therefore, to determine the convergence of the correction effect index, it is possible that the correction effect index falls below a predetermined value.

[0101] Alternatively, the process from step S901 to step S910 could be repeated a predetermined number of times.

[0102] Furthermore, updating the damping adjustment parameters can be done by calculating the derivative of the correction effect index using the gradient method to derive new damping adjustment parameters, or by using optimization methods such as the Nelder-Mead method.

[0103] Furthermore, since there is only one attenuation adjustment parameter, it can be evaluated comprehensively, similar to the grid method. Finally, the scattered radiation correction process is performed using the attenuation adjustment parameter that resulted in the smallest correction effect index, and the scattered radiation estimation process (step S204) is terminated.

[0104] In Modification 1, by finding attenuation adjustment parameters that reduce the correction effect index, the attenuation adjustment parameters can be optimized for each subject, further improving the image quality of cone-beam CT images.

[0105] Next, we will describe a modified example 2. Note that, for brevity, only the differences from the embodiment will be explained.

[0106] Modification 2 involves changing the attenuation adjustment parameter for each shooting angle. Specifically, by changing the attenuation adjustment parameter for each projected image, it becomes possible to accurately estimate scattered radiation by taking into account the difference in scattered radiation dose due to the difference in reflection of bed 2.

[0107] As shown in Figure 13, in cone-beam CT, the imaging range can be expanded by shifting the X-ray detector 4 from the rotational axis 1000 during imaging. When imaging in this way, the degree to which the bed 2 is reflected in the X-ray projection data changes depending on the imaging angle, and differences in scattered radiation dose also occur.

[0108] Our investigations have shown that at imaging angles where X-rays are directly irradiated onto the bed 2 (mainly from 40 to 140 degrees), the X-rays directly irradiated onto the bed 2 are scattered and a large amount is captured by the X-ray detector 4.

[0109] In contrast, to accurately estimate the scattered dose which changes with the shooting angle, the modified example 2 changes the attenuation adjustment parameter for each projection angle.

[0110] For example, at imaging angles where X-rays are directly irradiated onto bed 2, the estimated scattered dose can be increased by setting a smaller attenuation adjustment parameter.

[0111] The amount of change to the attenuation adjustment parameter can be determined in advance through imaging tests using a phantom or similar device that simulates a human body. Alternatively, the user may set the amount of change to the attenuation adjustment parameter to any value they choose.

[0112] Furthermore, the attenuation adjustment parameter may be changed discretely or continuously at any angle. For example, if equation (4) is used for f(τ), which represents the attenuation effect of the first scattered radiation distribution based on the subject thickness, the attenuation adjustment parameter typically has values ​​between 0.01 and 0.02. However, it may be set to 0.01 at the imaging angle where X-rays are directly irradiated onto bed 2 (40 to 140 degrees) and to 0.02 at other angles. Alternatively, it may be continuously changed using trigonometric functions or the like so that it is 0.01 at 90 degrees and 0.02 at 270 degrees.

[0113] Next, the effects of the above-described examples and modified versions will be explained.

[0114] The scattered radiation correction unit 13 in the above-described embodiment calculates a first scattered radiation from X-ray projection data captured by an X-ray imaging device using a point-symmetric scattering kernel, and calculates a second scattered radiation from the first scattered radiation based on the thickness of the subject being imaged.

[0115] This configuration eliminates the need for an asymmetric scattering kernel, thus reducing image degradation due to scattered radiation without increasing computation time compared to conventional methods. Consequently, it is possible to obtain cone-beam CT images with reduced image degradation due to scattered radiation.

[0116] Furthermore, since the attenuation of scattered X-rays within an object is mainly determined by the object's thickness distribution, the scattered X-rays can be estimated with higher accuracy by calculating the second scattered ray by reducing the first scattered ray based on the object's thickness.

[0117] Furthermore, by calculating the first scattered ray based on a point-symmetric scattering kernel and X-ray projection data, the computation process can be completed in a shorter time.

[0118] Furthermore, by calculating the thickness of the subject from the X-ray projection data, the thickness of each individual subject can be accurately determined without taking much time.

[0119] Furthermore, by calculating the second scattered radiation based on a function including one or more attenuation adjustment parameters, the object thickness, and the first scattered radiation, the second scattered radiation corresponding to the state of object 1 can be determined with greater accuracy.

[0120] Furthermore, by including the attenuation adjustment parameter in the function, which takes a value range of 0 or greater and 1 or less, the effect of attenuation of the first scattered rays within object 1 can be accurately represented, thereby further improving the accuracy of the calculation.

[0121] Furthermore, by changing the attenuation adjustment parameter for each shooting angle, the attenuation effect by the subject 1 can be adjusted according to the reflection range of the bed 2, and the increase in scattered X-rays due to the bed 2 can be taken into account, thereby further improving the image quality of the resulting cone-beam CT image.

[0122] Furthermore, by evaluating the scattered radiation correction effect of the reconstructed image and optimizing the attenuation adjustment parameters, higher-quality cone-beam CT images can be obtained.

[0123] <Other> The present invention is not limited to the embodiments described above, and various modifications and applications are possible. The embodiments described above are explained in detail for the purpose of clearly illustrating the present invention, and are not necessarily limited to those having all the configurations described.

[0124] 1...Subject (patient) 2...Bed 3...Rotation support device 4...X-ray detector (X-ray detector, X-ray imaging device, radiation imaging device) 5...X-ray tube (X-ray source, X-ray imaging device, radiation imaging device) 6...X-ray diaphragm 10...Control unit (radiation imaging device, X-ray imaging device, X-ray CT device) 11...Display unit 12...Memory unit 13...Scattered radiation correction unit (calculation processing unit, radiation imaging device, X-ray CT device) 14...Reconstruction unit (X-ray CT device) 15...Imaging parameter setting unit 60...Cylindrical phantom 100...Cone-beam CT imaging device (X-ray CT device) 110...Control computer 120...Therapeutic radiation control device 130...Therapeutic radiation generator 140...Therapeutic radiation irradiation device 150...Radiation therapy system 1000...Rotational axis

Claims

1. A processing unit that calculates a first scattered ray from X-ray projection data acquired by an X-ray imaging device using a point-symmetric scattering kernel, and calculates a second scattered ray from the first scattered ray based on the thickness of the subject being imaged.

2. A calculation processing device according to claim 1, wherein a second scattered radiation is calculated by reducing the first scattered radiation based on the thickness of the subject.

3. A processing unit according to claim 1 or 2, which calculates the first scattered ray based on the point-symmetric scattering kernel and the X-ray projection data.

4. A processing unit according to any one of claims 1 to 3, wherein the processing unit calculates the thickness of the subject from the X-ray projection data.

5. A calculation processing device according to any one of claims 1 to 4, wherein the second scattered ray is calculated based on a function including one or more attenuation adjustment parameters, the thickness of the subject, and the first scattered ray.

6. The arithmetic processing device according to claim 5, wherein the function including the damping adjustment parameter takes a value range of 0 or more and 1 or less.

7. A calculation processing device according to claim 4, wherein the attenuation adjustment parameter is changed for each shooting angle.

8. A radiation imaging apparatus comprising an X-ray source, an X-ray detector, and a processing unit according to any one of claims 1 to 7.

9. An X-ray CT apparatus comprising the radiation imaging apparatus described in claim 8, which calculates corrected X projection data from the X-ray projection data and the second scattered radiation, and reconstructs a CT scan.

10. An X-ray CT apparatus according to claim 9, wherein the scattered radiation correction effect of the reconstructed image is evaluated and the attenuation adjustment parameter is optimized.

11. A radiotherapy system comprising an X-ray CT apparatus according to claim 9 or 10.

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