CT scanning method for X-ray source tube voltage adaptive modulation

Abstract

The invention discloses an X-ray source tube voltage adaptive modulation CT scanning method, and relates to the technical field of tube voltage regulation, and the method comprises the following steps: S1, pre-scanning a patient; s2, drawing a scanning part contour according to a pre-scanning result; s3, selecting a tube voltage modulation optimization model; s4, determining a corresponding tube voltage modulation scheme according to the part contour and the optimization model; s5, scanning the patient by adopting a tube voltage modulation scheme, and acquiring projection data of a plurality of energy spectrums by utilizing a double-layer detector or a photon counting detector; and S6, performing dual-energy decomposition on the projection of each angle, and reconstructing to obtain a cross-sectional image of the measured part of the patient. According to the method, the tube voltage modulation scheme is obtained according to the parameter model, the image quality is effectively improved, the utilization rate of the radiation dose is improved, and a better CT image is obtained through less dose.

Description

Technical Field

[0001] The invention relates to the technical field of tube voltage regulation, and in particular to a CT scanning method for adaptively modulating the tube voltage of an X-ray source. Background Art

[0002] The main influencing factors of CT dose include two aspects: patient factors and CT parameter factors.

[0003] Patient-related factors mainly include the patient's organs being measured and the patient's own physical parameters. Different clinical detection tasks require different doses. For the examination of high-contrast structures, a lower dose can be used. In addition, the patient's body size will also affect the dose. Studies have shown that BMI (body mass index) is positively correlated with radiation dose. Patients with larger bodies need to increase the dose to ensure the image quality required for the examination. When performing CT examinations on children, a lower dose is required because they are more sensitive to radiation. In addition to the above two factors, whether the patient's scanning position is located in the center of the CT field of view during the scan will also affect the dose.

[0004] CT parameter factors include: tube voltage, tube current, exposure time and pitch. The size of the tube voltage determines the range of X-ray photon energy. The larger the tube voltage, the greater the maximum energy that the photon can have. The tube current determines the number of photons released per unit time. Electrons hit the anode target to produce photons. The larger the current, the more photons are produced. The length of exposure time directly affects the radiation dose received by the subject. The size of the pitch will first cause different degrees of overlap in the X-ray beam, and secondly, it will also have a certain impact on the exposure time. The following formula is used to describe the effect of the above parameters on CT dose:

[0005] CTDI vol =B0+B1*(body size)+B2*(kVp 2 )+B3*(mAs)+B4*(1 / pitch)

[0006] Where B0-B4 are the estimated regression coefficients, CTDI vol is the volume CT dose index (volume CT doseindex). It can be seen from the formula that different influencing factors have different degrees of influence on the dose when they change.

[0007] 1. Analysis of current research status and existing problems at home and abroad

[0008] Due to the uncontrollability of patient factors, the CT dose control methods currently studied at home and abroad mainly focus on adjusting and improving CT scanning technology, which are mainly divided into low-dose scanning methods, low-dose reconstruction methods, and dose control through other advanced technologies.

[0009] 1.1 Low-dose scanning method

[0010] According to the above-mentioned CT influencing parameters, the main dose control methods include: automatic tube current modulation, tube voltage selection, hybrid scanning, low-angle scanning, anatomical landmark detection scanning, etc.

[0011] 1) Automatic tube current modulation

[0012] The effect of tube current on image quality is mainly reflected in the noise. If the tube current is reduced under the same object, the number of photons that finally reach the detector will be reduced accordingly. After image reconstruction, the proportion of noise contained in the image will increase accordingly. Automatic tube current modulation (ATCM) adjusts the tube current at different angles according to the X-ray path and attenuation at different angles to ensure that the dose is reduced and the noise is reduced without losing image quality. This method is more effective for parts such as shoulders and hips where the difference between sagittal distance and coronal distance is large. ATCM can achieve a stronger dose reduction effect by defining more sophisticated models, organ-specific scheme customization, and more scientific dose index constraints.

[0013] 2) Automatic tube voltage selection

[0014] The automatic tube voltage selection technology automatically selects the appropriate tube voltage based on the cross-sectional size of different measurement positions of the patient and different imaging purposes. The impact of tube voltage on image quality is mainly reflected in the contrast. Since different materials have different absorption capacities for energy X-ray photons, lower energy photons can better reflect the differences between materials of similar density and have better effects on soft tissue imaging. For soft tissue and larger models, the optimal tube voltage will be higher than 100kV. For high-density tissue and smaller models, the optimal kV will be lower. However, a reduction in tube voltage means a decrease in the penetration ability of X-ray photons. In order to achieve the signal-to-noise ratio required for imaging, the corresponding tube current needs to be adjusted. Therefore, in actual applications, the tube voltage needs to be selected based on multiple factors such as dose and examination site.

[0015] 3) Other technologies

[0016] Hybrid scanning achieves two different scanning ranges by setting two different source-to-detector distances, and then reconstructs the data through a specific algorithm. This method effectively reduces the total scanning dose by reducing the scanning range of some angles while ensuring image quality.

[0017] The low-angle scanning method is mainly based on the theory of compressed sensing. It uses mathematical methods to find the minimum number of angles required to reconstruct the image. The core of this method lies in the reconstruction method. Due to the problem of missing data, reconstruction is required through an iterative method. A smaller number of scanning angles can significantly reduce the radiation dose, but this method has a low computing speed and requires a longer reconstruction time. In addition, it relies more on the accuracy of the measured data.

[0018] Anatomical landmark detection scanning technology performs a low-dose spiral scan on the patient, then uses artificial intelligence to identify the specific detection site or organ and plan the scanning parameters by itself. This method can effectively reduce the scanning range of CT and thus reduce the total radiation dose.

[0019] 1.2 Low-dose reconstruction method

[0020] In order to ensure good image quality when scanning at a lower dose, the original reconstruction algorithm needs to be improved. Currently, the CT reconstruction algorithms on the market are mainly divided into three categories: analytical reconstruction algorithms, iterative reconstruction algorithms, and artificial intelligence reconstruction algorithms.

[0021] The analytical reconstruction algorithm is mainly based on FBP (Filtered back projection), and its main feature is fast calculation speed, but it has high requirements on the quality of collected data. There are also methods to apply it to the reconstruction of low-dose CT. The original data is processed by a special noise reduction algorithm and then reconstructed by FBP.

[0022] The iterative algorithm sets up a loss function and repeatedly iterates to find the optimal image. The reconstruction speed is slow but it can effectively reduce image noise. By improving image quality, it can reduce the dose requirement under the same image quality requirements. However, this type of algorithm also has some limitations. When the dose is too low, the iterative algorithm will cause image quality degradation and distortion.

[0023] As an emerging algorithm, artificial intelligence algorithm has great potential. It can reconstruct high-quality images at low dose levels. At the same time, its speed and image quality are greatly improved compared with iterative reconstruction methods. Introducing artificial intelligence reconstruction algorithms in clinical examinations can effectively reduce the radiation exposure of patients. This type of algorithm is also constantly improving with the continuous advancement of artificial intelligence technology, and many methods based on different network structures have emerged. In future CT scans, artificial intelligence algorithms may become the mainstream reconstruction method.

[0024] 1.3 Other advanced technologies

[0025] In addition to the improvement of scanning and reconstruction methods, technological breakthroughs in CT components will also enhance the dose efficiency of CT scanning. X-ray detectors are the basic components of CT systems, which determine the image quality and dose efficiency. At present, multi-row detectors with larger areas have been developed. More rows can scan a larger range, thereby reducing the exposure time and exposure of whole-organ scans. Double-layer detectors are a type of detector used in dual-energy CT. By superimposing two layers of detectors to measure information of two different spectra, spectral images can be generated, which can effectively reduce artifacts and reduce radiation dose. In addition, the emergence of photon technology detector CT (Photon-counting detector CT, PCD-CT) provides a new solution for dose control. This detector can measure the energy of each incident X-ray photon individually. The main advantages include: no electronic noise, higher radiation dose efficiency and better spatial resolution. In actual clinical applications, compared with traditional energy-integrating detector CT (Energy-integrating detector CT, EID-CT), PCD-CT has shown higher image quality and better dose control capabilities.

[0026] Tube filters are mainly used to filter the low-energy X-rays produced by the CT tube. These rays cannot penetrate the human body to achieve the purpose of imaging. If they are not filtered out, they will produce a large radiation dose. In current research, tin filters have better performance and are often used in low-dose CT.

[0027] As people pay more attention to dose, dose control methods have been continuously improved and developed. The emergence of technologies such as photon counting detectors and artificial intelligence algorithms will provide new possibilities for dose control. Most products on the market combine various technologies to achieve better results. However, there will be more CT examinations in the future, and there will inevitably be new challenges and requirements for dose control. According to research, the dose is linearly related to the tube current and linearly related to the square of the tube voltage, so tube voltage modulation has more potential. At the same time, due to technical limitations, only images under the same spectral conditions can be reconstructed in the same scan, and the choice of tube voltage is not flexible. There is still room for improvement in this area. Summary of the invention

[0028] The present invention aims to solve the problem of image reconstruction under different tube voltages and ensure the accuracy and availability of images. Therefore, the present invention discloses a CT scanning method with adaptive modulation of X-ray source tube voltage.

[0029] The specific scheme of the present invention is as follows:

[0030] A CT scanning method with adaptive modulation of X-ray source tube voltage, the specific steps are as follows:

[0031] S1. Pre-scan the patient.

[0032] S2, drawing the scan part contour according to the pre-scan result;

[0033] S3, select the tube voltage modulation optimization model, such as the noise homogenization principle and the contrast-to-noise ratio dose rate maximization principle;

[0034] S4, determining the tube voltage modulation scheme corresponding to the measured part according to the part contour and the optimization model;

[0035] S5, scanning the patient using a tube voltage modulation scheme, using a double-layer detector or a photon counting detector to obtain projection data of multiple energy spectra;

[0036] S6. Perform dual-energy decomposition on the projection at each angle and reconstruct a tomographic image of the measured part of the patient.

[0037] Furthermore, the specific steps of S1 include:

[0038] S1-1. Determine the part of the patient's body that needs to be scanned;

[0039] S1-2. Measure the contour and size of the scanned part of the patient. The measurement methods include but are not limited to: frontal and lateral scanning, optical visual detection, positioning scanning, and spiral positioning scanning;

[0040] S1-3. Using the above measurement method, obtain the contour and size information of the scanned part of the patient.

[0041] The specific steps of S2 include:

[0042] S2-1. Draw the outline of the scanned part according to the pre-scan result.

[0043] The specific steps of S3 include:

[0044] S3-1. Select the tube voltage modulation optimization model, such as the noise homogenization principle and the contrast-to-noise ratio-dose rate maximization principle.

[0045] Furthermore, the specific steps of S4 include:

[0046] S4-1, calculating the simulated ray path length of the scanned part at different scanning angles according to the scanned part and the contour and size information of the scanned part obtained in step S1-3;

[0047] S4-2, using each projection angle obtained in S4-1 to simulate the ray path length, and using the noise formula and the dose formula to calculate the noise and dose corresponding to each projection angle at different tube voltages;

[0048] S4-3. Using the noise and dose lists corresponding to different tube voltages at various angles calculated in S4-2, according to the noise homogenization principle and the contrast-to-noise-dose ratio maximization principle selected in S3, the tube voltage used at each angle is optimized.

[0049] Furthermore, the specific steps of S5 include:

[0050] S5-1, transmitting the tube voltage modulation scheme obtained in S4-3 to the CT control unit;

[0051] S5-2. Scan the patient using a tube voltage modulation scheme and use a double-layer detector or a photon counting detector to obtain projection data of multiple energy spectra.

[0052] Furthermore, the specific steps of S6 include:

[0053] S6-1. Utilize energy spectrum decomposition technology to perform energy spectrum decomposition at each angle using two or more groups of projection information with different energies, and then reconstruct a virtual monoenergetic image using the decomposition results at each angle to obtain the required tomographic image of the patient's measured part.

[0054] Compared with the prior art, the present invention has the following advantages: (1) the present invention can accurately design a parameter model and simulate the model according to actual conditions, further obtaining more accurate data and providing high-precision guidance for later use; (2) the tube voltage modulation scheme obtained by the present invention based on the parameter model can effectively improve image quality and at the same time improve the utilization rate of radiation dose, thereby obtaining better images with less dose. BRIEF DESCRIPTION OF THE DRAWINGS

[0055] Figure 1 It is a schematic diagram of the overall steps of the present invention.

[0056] Figure 2 It is a schematic diagram of the curve of the shoulder theoretical calculation results in the dose model verification of the present invention.

[0057] Figure 3 It is a schematic diagram of the curve of the shoulder simulation measurement results in the dose model verification of the present invention.

[0058] Figure 4 It is a schematic diagram of the curve of the theoretical calculation results of the chest in the dose model verification of the present invention.

[0059] Figure 5 It is a schematic diagram of the curve of the chest simulation measurement results in the dose model verification of the present invention.

[0060] Figure 6 It is a schematic diagram of the curve of the theoretical calculation results of the abdomen in the dosage model verification of the present invention.

[0061] Figure 7 It is a schematic diagram of the curve of the abdomen simulation measurement results in the dose model verification of the present invention.

[0062] Figure 8 It is a curve diagram of the theoretical calculation results of the hip in the dose model verification of the present invention.

[0063] Fig. 9 It is a schematic diagram of the curve of the hip simulation measurement results in the dose model verification of the present invention.

[0064] Fig.10 This is a schematic diagram of the shoulder measurement result curve of the lowest noise group with equal dose in the present invention.

[0065] Fig.11 It is a schematic diagram of the chest measurement result curve of the lowest noise group with equal dose in the present invention.

[0066] Fig.12 This is a schematic diagram of the abdominal measurement result curve of the lowest noise group with equal dose in the present invention.

[0067] Fig.13 This is a schematic diagram of the hip measurement result curve of the lowest noise group with equal dose in the present invention.

[0068] Fig.14 It is a schematic diagram of the comparison curve of the lowest noise group of equal dose in the present invention.

[0069] Fig.15 It is a curve diagram of the abdominal dose model and the noise model in the equal dose lowest noise group of the present invention.

[0070] Fig.16 It is a curve schematic diagram of the shoulder dose model and the noise model in the equal dose lowest noise group of the present invention.

[0071] Fig.17 It is a curve diagram of the hip dose model and noise model in the equal dose lowest noise group of the present invention.

[0072] Fig.18 It is a curve diagram of the chest dose model and the noise model in the equal-dose lowest noise group of the present invention.

[0073] Fig.19 This is a schematic diagram of the shoulder measurement results curve of the lowest noise dose group of the present invention.

[0074] Fig. 20 This is a schematic diagram of the chest measurement results curve of the lowest noise dose group of the present invention.

[0075] Fig.21 This is a schematic diagram of the abdominal measurement results curve of the lowest noise dose group of the present invention.

[0076] Fig. 22This is a schematic diagram of the hip measurement results curve of the lowest noise dose group of the present invention.

[0077] Fig.23 It is a schematic diagram of the comparison curve of the lowest noise dose group of the present invention.

[0078] Fig.24 It is a curve diagram of the abdominal dose model and the noise model in the lowest noise dose group of the present invention.

[0079] Fig.25 It is a curve schematic diagram of the shoulder dose model and the noise model in the lowest noise dose group of the present invention.

[0080] Fig.26 It is a curve diagram of the hip dose model and the noise model in the lowest noise dose group of the present invention.

[0081] Fig. 27 It is a curve diagram of the chest dose model and the noise model in the lowest noise dose group of the present invention.

[0082] Fig.28 This is a diagram showing the simulation results of tube voltage modulation of the present invention. DETAILED DESCRIPTION

[0083] It should be noted that, in the absence of conflict, the embodiments of the present invention and the features in the embodiments may be combined with each other.

[0084] In the description of the present invention, it should be understood that the terms "center", "longitudinal", "lateral", "up", "down", "front", "back", "left", "right", "vertical", "horizontal", "top", "bottom", "inside", "outside" and the like indicate positions or positional relationships based on the positions or positional relationships shown in the accompanying drawings, and are only for the convenience of describing the present invention and simplifying the description, rather than indicating or implying that the device or element referred to must have a specific orientation, be constructed and operated in a specific orientation, and therefore cannot be understood as limiting the present invention. In addition, the terms "first", "second", and the like are only used for descriptive purposes, and cannot be understood as indicating or implying relative importance or implicitly indicating the number of technical features indicated. Thus, features defined as "first", "second", and the like may explicitly or implicitly include one or more of the features. In the description of the present invention, unless otherwise specified, "multiple" means two or more.

[0085] In the description of the present invention, it should be noted that, unless otherwise clearly specified and limited, the terms "installed", "connected", and "connected" should be understood in a broad sense, for example, it can be a fixed connection, a detachable connection, or an integral connection; it can be a mechanical connection or an electrical connection; it can be a direct connection, or it can be indirectly connected through an intermediate medium, or it can be the internal communication of two components. For ordinary technicians in this field, the specific meanings of the above terms in the present invention can be understood by specific circumstances.

[0086] Example: Figure 1 As shown, a CT scanning method with adaptive modulation of X-ray source tube voltage, the specific steps are as follows:

[0087] S1. Pre-scan the patient.

[0088] The specific steps of S1 include:

[0089] S1-1. Determine the part of the patient that needs to be scanned;

[0090] S1-2. Measure the contour and size of the scanned part of the patient. The measurement methods include but are not limited to: frontal and lateral scanning, optical visual detection, positioning scanning, and spiral positioning scanning;

[0091] S1-3. Using the above measurement method, obtain the contour and size information of the scanned part of the patient.

[0092] S2. Draw the contour of the scanned part according to the pre-scanning result.

[0093] S3. Select the tube voltage modulation optimization model, such as the noise homogenization principle and the contrast-to-noise ratio-dose rate maximization principle.

[0094] First, five phantoms were set up, corresponding to four different parts of the human body and a cylindrical simulation phantom. The phantoms were made of water.

[0095] The pixel coordinates of the central pixel are taken as the optimization target of modulation. In order to simplify the model, only the central ray passing through the central pixel is considered, and the noise of the central pixel is regarded as the direct summation of the projection variances at various angles. is the noise at the center of the reconstructed image, N P represents the number of angles, Ni represents the number of photons detected by the detector at angle i, and the variance of a single projection can be regarded as N i The inverse of the noise of the central pixel can be given by formula (1):

[0096]

[0097] N i The size of the incident photon number N 0iand attenuation, and the attenuation coefficient of the material is affected by the type of material and the energy of the photon, so N i It is expressed as formula (2):

[0098]

[0099] where e -μELi represents the attenuation of X-rays when passing through an object at angle i, L i is the object path length that the X-ray passing through the central pixel passes through, Sp i (E) represents the energy spectrum distribution at a certain tube voltage level. According to formula (1) and formula (2), the relationship between the noise of the central pixel and the path length and the tube voltage can be obtained as formula (3):

[0100]

[0101] Among them, E max is the maximum energy size under the tube voltage, N 0i is the number of initial photons released by the tube at angle i after filtering, Sp i (E) is the energy spectrum used at angle i, E is the photon energy, μ E is the attenuation coefficient of a certain substance for photons of energy E, L i It is the attenuation path length of the central ray beam at angle i passing through the object being measured.

[0102] The above noise model involves the number of incident photons and the spectrum, both of which will affect the radiation dose generated during scanning, and constraints need to be added. The dose constraint in tube voltage modulation is not equivalent to the total number of incident photons. At the same time, under the same tube current, different tube voltages will also affect the number of photons. Therefore, the following dose expression is given as formula (4):

[0103]

[0104] Dose is the dose deposited during CT scanning, width j is the beam width covered by the object at angle i, μ enE is the absorption coefficient of a certain substance for photons of energy E, width i It will significantly affect the accuracy of the dose model. For example, for the shoulder phantom, if only the central ray is considered, the dose deposition at 0° is significantly less than that at 90°. However, the actual situation is the opposite, because more beams are covered at 0°. In addition, M is the mass of the scanned part of the object. However, since the shape and size of the phantom used are fixed, the subsequent dose exploration only considers the molecular part.

[0105] For the above two models, both include N 0i, so we need to clarify N 0i The relationship with the tube voltage is as follows:

[0106] N 0i = kV i 2

[0107] Among them, kV i is the tube voltage used at angle i, and the tube voltage modulation is set to, under a fixed dose, by adjusting the tube voltage at different angles to minimize the image noise of the central pixel. Due to the complexity of formula (3) and formula (4), it is difficult to combine the two equations to obtain an analytical solution. Therefore, by calculating the noise and dose data at different angles and tube voltages, the final modulation scheme is obtained by using the optimization method.

[0108] S4, determining the tube voltage modulation scheme corresponding to the measured part according to the scanned part contour and the optimization model;

[0109] The specific steps of S4 include:

[0110] S4-1, calculating the simulated ray path length of the scanned part at different scanning angles according to the scanned part and the contour and size information of the scanned part obtained in step S1-3;

[0111] S4-2, using each projection angle obtained in S4-1 to simulate the ray path length, and using the noise formula and the dose formula to calculate the noise and dose corresponding to each projection angle at different tube voltages;

[0112] The noise formula is

[0113]

[0114] The dosage formula is:

[0115]

[0116] S4-3. Using the noise and dose lists corresponding to different tube voltages at various angles calculated in S4-2, according to the noise homogenization principle and the contrast-to-noise-dose ratio maximization principle selected in S3, the tube voltage used at each angle is optimized.

[0117] First, the dose and noise are listed as follows using the above formulas:

[0118]

[0119] Assume the modulation scheme matrix is ​​an unknown number, x view，kV The values ​​of are only 0 and 1. When the value is 1, it means that kV is used as the tube voltage under the angle view. For a certain angle, only one kV corresponding to x is 1.

[0120]

[0121] Expand the above three matrices into one-dimensional form

[0122]

[0123] After that, we can get the equations of total dose and total noise. By controlling one of the equations as a constraint, we can use the existing shaping linear optimization method to quickly get the modulation scheme. The following is the optimization equation for the modulation scheme with the minimum total noise when the total dose is fixed.

[0124]

[0125] The total noise is fixed, and the equation for minimizing the total dose only requires f(x) = Noise a ×modulation′ a Noise a and Dose a ×modulation′ a Dose in ≤const a exchange.

[0126] The dose model simulation measurement was performed. By comparing the theoretical calculated energy deposition and the simulated measurement results, it can be seen that the dose variation trends of the two phantoms with tube voltage and projection angle are basically consistent for different phantoms. For the oval-shaped phantoms (chest, abdomen, hip), the continuity of the theoretical calculation results at around 80° is poor. By comparing the measured results with the theoretical calculation results (such as Figure 2-Figure 9 This dosage model is feasible.

[0127] The modulation scheme is obtained by table lookup method, and two different constraint schemes are obtained: (1) keeping the total dose unchanged and reducing the image noise; (2) keeping the total noise unchanged and reducing the dose used for scanning.

[0128] (1) Keep the total dose constant and reduce image noise

[0129] For the equal-dose minimum noise group, more scanning angles will use the tube voltage of 140 kV to reduce the projection variance on the longer path. For the shorter path, using a smaller voltage can distribute more of the overall dose to the projection angle of the long path. At the same time, a small voltage will lead to an increase in the projection variance, which can reduce the variance difference between different scanning angles to a certain extent and enhance the isotropy of the noise. The data of the equal-dose minimum noise group are plotted into a graph (such as Fig.10 As shown in the figure), and in the case of equal dose, noise model and dose model are modeled for different parts respectively (as shown in the figure). Figure 11-Figure 14 shown).

[0130] (2) Keep the total noise constant and reduce the scanning dose

[0131] For the equal noise minimum dose group, the data of the equal noise minimum dose group are plotted into a graph (e.g. Fig.15 As shown in the figure, it can be seen that the overall tube voltage is lower than that of the equal-dose group. For longer paths, more 140kV tube voltages are still used because the noise of longer paths is more sensitive to tube voltage in terms of noise reduction. The benefits of using high voltage at these scanning angles are greater. Under the condition of equal noise, noise models and dose models are modeled for different parts respectively (such as Figure 16-Figure 19 shown).

[0132] By analyzing the results of theoretical calculations, the dose changes in the equal-dose group are all ≤1%, which can be considered as unchanged dose. For phantoms with large length and width changes (chest, hip), the noise reduction effect is obvious and can reach more than 30%. The total noise in the equal-noise group can remain unchanged, and the dose of various phantoms can be reduced to a certain extent. Similarly, for phantoms with large length and width changes, the dose reduction can reach 30%-50%. The noise reduction and dose reduction effects of the two groups of modulation schemes are relatively obvious, as shown in the following table.

[0133]

[0134] By comparing the results after tube voltage modulation (such as Fig. 20 As shown in the table below, the doses of the four phantoms in the equal-dose group were all ≤1% compared with those in the control group, which means that the doses remained unchanged. The doses in the equal-noise lowest-dose group were significantly improved. For phantoms with a smaller length-width difference, the dose reduction could reach 10%, while for phantoms with a larger length-width difference, the effect could reach 30%-50%, as shown in the table below.

[0135]

[0136] The tube voltage used at different scanning angles in tube voltage modulation should be positively correlated with the attenuation of the central ray. For a homogeneous phantom, the tube voltage is positively correlated with the path length of the central ray passing through the phantom. The two modulation schemes of the equal-dose minimum noise group and the equal-noise minimum dose group are the optimal modulation schemes.

[0137] S5, scanning the patient using a tube voltage modulation scheme, using a double-layer detector or a photon counting detector to obtain projection data of multiple energy spectra;

[0138] The specific steps of S5 include:

[0139] S5-1, transmitting the tube voltage modulation scheme obtained in S4-3 to the CT control unit;

[0140] S5-2. Scan the patient using a tube voltage modulation scheme and obtain projection data of multiple energy spectra using a double-layer detector or a photon counting detector.

[0141] S6. Perform dual-energy decomposition on the projection at each angle and reconstruct a tomographic image of the measured part of the patient.

[0142] The specific steps of S6 include:

[0143] S6-1. Utilize energy spectrum decomposition technology to perform energy spectrum decomposition at each angle using two or more groups of projection information with different energies, and then reconstruct a virtual monoenergetic image using the decomposition results at each angle to obtain the required tomographic image of the patient's measured part.

[0144] CT imaging measures and reconstructs the internal material information of an object based on the different attenuation abilities of different materials to X-rays. However, the same material exhibits different attenuations under different tube voltages. The projection data obtained using different tube voltages in the same scan cannot be directly used for reconstruction. Therefore, energy spectrum decomposition is performed in the projection domain to convert the attenuation information into material distribution information, and then the final image is obtained by reconstructing a virtual monoenergetic image.

[0145] Energy spectrum decomposition in the projection domain requires obtaining at least two sets of projection data generated by different energy spectra at the same scanning angle. In order to solve this problem, a double-layer detector or a photon counting detector is used to obtain different energy data at the same angle. After that, the attenuation information is converted into material distribution information through energy spectrum decomposition, and a virtual monoenergetic map is constructed to finally reconstruct the image.

[0146] In addition, it should be noted that the shapes and names of the parts and components of the specific embodiments described in this specification may be different. Any equivalent or simple changes made based on the concept, structure, features and principles of the patent of the present invention are included in the protection scope of the patent of the present invention. The technicians in the technical field of the present invention can make various modifications or supplements to the specific embodiments described or replace them in a similar manner, as long as they do not deviate from the structure of the present invention or exceed the scope defined by the claims, they should all fall within the protection scope of the present invention.

Claims

1. A CT scanning method with adaptive modulation of X-ray source tube voltage, characterized in that: The specific steps are as follows: S1. Pre-scan the patient. S2, drawing the contour of the scanned part according to the pre-scanning result; S3, select the tube voltage modulation optimization model, such as the noise homogenization principle and the contrast-to-noise ratio dose rate maximization principle; S4, determining the tube voltage modulation scheme corresponding to the measured part according to the part contour and the optimization model; S5, scanning the patient using a tube voltage modulation scheme, using a double-layer detector or a photon counting detector to obtain projection data of multiple energy spectra; S6. Perform dual-energy decomposition on the projection at each angle and reconstruct a tomographic image of the measured part of the patient.

2. A CT scanning method with adaptive modulation of X-ray source tube voltage as claimed in claim 1, characterized in that: The specific steps of S1 include: S1-1. Determine the part of the patient's body that needs to be scanned; S1-2. Measure the contour and size of the scanned part of the patient. The measurement methods include but are not limited to: frontal and lateral scanning, optical visual detection, positioning scanning, and spiral positioning scanning; S1-3. Using the above measurement method, obtain the contour and size information of the scanned part of the patient.

3. A CT scanning method with adaptive modulation of X-ray source tube voltage as claimed in claim 1, characterized in that: The specific steps of S4 include: S4-1, calculating the simulated ray path length of the scanned part at different scanning angles according to the scanned part and the contour and size information of the scanned part obtained in step S1-3; S4-2, using each projection angle obtained in S4-1 to simulate the ray path length, and using the noise formula and the dose formula to calculate the noise and dose corresponding to each projection angle at different tube voltages; S4-3. Using the noise and dose lists corresponding to different tube voltages at various angles calculated in S4-2, according to the noise homogenization principle and the contrast-to-noise-dose ratio maximization principle selected in S3, the tube voltage used at each angle is optimized.

4. A CT scanning method with adaptive modulation of X-ray source tube voltage as claimed in claim 1, characterized in that: The specific steps of S5 include: S5-1, transmitting the tube voltage modulation scheme obtained in S4-3 to the CT control unit; S5-2. Scan the patient using a tube voltage modulation scheme and use a double-layer detector or a photon counting detector to obtain projection data of multiple energy spectra.

5. The CT scanning method with adaptive modulation of X-ray source tube voltage according to claim 1, characterized in that: The specific steps of S6 include: S6-1. Utilize energy spectrum decomposition technology to perform energy spectrum decomposition at each angle using two or more groups of projection information with different energies, and then reconstruct a virtual monoenergetic image using the decomposition results at each angle to obtain the required tomographic image of the patient's scanned part.

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