Program and system for processing detection signal of radiation transmitted through imaging subject to generate radiological image

Abstract

To provide a method capable of shortening calibration while maintaining the image quality of a reconstructed image.SOLUTION: The tube voltage is alternately switched during rotation so as to form a waveform having a rising part from a low tube voltage to a high tube voltage, a falling part from the high tube voltage to the low tube voltage, a high steady-state interval between the rising part and the falling part, and a low steady-state interval between the falling part and the rising part. The method also comprises: receiving input for a rotation speed or number of views per rotation; identifying a tube voltage waveform corresponding to the rotation speed or the number of views per rotation; calculating a low-energy average of radiation corresponding to a low voltage interval including the low steady-state interval of the tube voltage waveform and a part of the falling part of the tube voltage waveform, and a high-energy average of radiation corresponding to a high voltage interval including the high steady-state interval of the tube voltage waveform and another part of the falling part of the tube voltage waveform; and setting a parameter used for creating the image in accordance with the low-energy average and the high-energy average.SELECTED DRAWING: Figure 6

Description

[Technical Field]

[0001] The present invention relates to a radiation imaging apparatus, and more particularly to a technique for realizing multi-energy imaging. [Background technology]

[0002] As an imaging method using a radiation imaging device represented by an X-ray CT device (X-ray Computed Tomography System), there is known an imaging method called dual-energy imaging, in which data is collected by switching the radiation tube voltage (see, for example, Patent Documents 1 and 2).

[0003] In the case of an X-ray CT scanner, this imaging method utilizes the difference in X-ray energy absorption spectrum between materials to obtain an image in which a specific material in the subject is emphasized or suppressed (removed). Specifically, for example, an object containing a first material and a second material is irradiated with first and second X-rays having different energy spectra to collect first and second X-ray projection data corresponding to multiple views. Then, a first density image is reconstructed based on the first X-ray projection data, and a second density image is reconstructed based on the second X-ray projection data. Based on the first and second density images, a density distribution image showing the density distribution of a certain material or a monochromatic image of a certain energy spectrum, a dual-energy image, is obtained.

[0004] As a method for collecting the first X-ray projection data and the second X-ray projection data as described above, for example, a method can be considered in which projection data is collected while the X-ray tube voltage is alternately switched between a low tube voltage and a high tube voltage for each view, i.e., each time the gantry of the X-ray CT device rotates by one view. [Prior art documents] [Patent documents]

[0005] [Patent Document 1] Japanese Patent Application Laid-Open No. 2004-65975 [Patent Document 2] Japanese Patent Application Laid-Open No. 2009-153829 [Patent Document 3] Japanese Patent Application Laid-Open No. 2015-112478 Summary of the Invention [Problem to be solved by the invention]

[0006] As described above, when acquiring projection data while alternately switching the X-ray tube voltage between low and high tube voltages, it is desirable to switch from low to high tube voltage and from high to low tube voltage instantaneously. However, due to limitations in transformer performance, rise times from low to high tube voltage and fall times from high to low tube voltage occur. These tube voltage waveform rise and fall times are affected not only by the potential difference between the high and low tube voltages, but also by the switching speed between the high and low tube voltages, the magnitude of the tube current, and other factors.

[0007] Regarding tube current, the rise time does not vary significantly whether the tube current is high or low, while the fall time is more significantly affected. For example, when the tube current is 400 mA, switching from high to low tube voltage occurs in a relatively short time, but when the tube current is 200 mA, switching from high to low tube voltage takes a relatively long time. That is, when the tube current is 400 mA, the fall time is relatively short, and when the tube current is 200 mA, the fall time is relatively short. Furthermore, the rise time and fall time are often curved rather than linear. When data is sampled continuously, i.e., when the curved portion of the tube voltage waveform is included in the sampling, changes occur in the energy spectrum of the irradiated X-ray beam, and the average energy of the sampled X-ray beam changes.

[0008] Furthermore, when the number of acquired views is constant, a change in the rotation speed changes the sampling rate, which in turn changes the cycle for switching from high to low tube voltage. This causes the duration of the high-voltage steady-state section of the tube voltage waveform corresponding to the high tube voltage and the duration of the low-voltage steady-state section of the tube voltage waveform corresponding to the low tube voltage to fluctuate. That is, as the rotation speed increases, the durations of both the high and low steady-state sections become shorter, and as the rotation speed decreases, the durations of both the high and low steady-state sections become longer.

[0009] On the other hand, to solve the problems of variations in sensitivity of individual detector elements and deviations of the actual X-ray energy spectrum from the ideal energy spectrum, calibration is performed according to an image reconstruction protocol (preset). As mentioned above, changes in the rotation speed and the magnitude of the tube current cause fluctuations in the energy spectrum of the irradiated X-rays, so calibration must be performed for each type of rotation speed and tube current used. This requirement results in longer downtime when the X-ray CT system cannot be used, longer imaging times for the complicated calibration and its calculations, and increased service costs.

[0010] Therefore, a new method is required that can shorten the time required for calibration imaging and calculation while maintaining the image quality of the reconstructed image, reduce the downtime of the radiation imaging apparatus, and reduce service costs. [Means for solving the problem]

[0011] The present disclosure at least partially addresses one or more of the above-identified problems or related problems by a system including a processor that processes detection signals of radiation transmitted through an imaging target to generate a radiological image. The radiation is radiation irradiated toward the imaging target while a radiation tube rotates around the imaging target, and the radiation includes low-energy radiation generated by applying a low tube voltage to the radiation tube and high-energy radiation generated by applying a high tube voltage to the radiation tube. The application of the low tube voltage and the high tube voltage to the radiation tube are alternately switched during the rotation of the radiation tube, to form a tube voltage waveform having a rising portion from the low tube voltage to the high tube voltage, a falling portion from the high tube voltage to the low tube voltage, a high steady-state section between the rising portion and the falling portion, and a low steady-state section between the falling portion and the rising portion. The processor receives input of the rotation speed and / or the number of views per rotation, identifies a tube voltage waveform corresponding to the rotation speed and / or the number of views per rotation using a waveform identification model, calculates a low-energy average value of radiation corresponding to a low-voltage section including a low steady-state section of the tube voltage waveform and a part of the falling portion of the tube voltage waveform, and a high-energy average value of radiation corresponding to a high-voltage section including a high steady-state section of the tube voltage waveform and another part of the falling portion of the tube voltage waveform, and sets parameters used to create a radiographic image according to the low-energy average value and the high-energy average value.

[0012] In another aspect of the present disclosure, one or more of the above-identified problems or related problems are at least partially addressed by a program that processes detection signals of radiation transmitted through an imaging target to generate a radiation image. The radiation is radiation irradiated toward the imaging target while a radiation tube rotates around the imaging target. The radiation includes low-energy radiation generated by applying a low tube voltage to the radiation tube and high-energy radiation generated by applying a high tube voltage to the radiation tube. The application of the low tube voltage and the high tube voltage to the radiation tube are alternately switched during the rotation of the radiation tube, thereby forming a tube voltage waveform having a rising portion from the low tube voltage to the high tube voltage, a falling portion from the high tube voltage to the low tube voltage, a high steady-state section between the rising portion and the falling portion, and a low steady-state section between the falling portion and the rising portion. The program further causes the processor to receive input of the rotation speed and / or the number of views per rotation, identify a tube voltage waveform corresponding to the rotation speed and / or the number of views per rotation using a waveform identification model, calculate a low-energy average value of radiation corresponding to a low-voltage section including a low steady-state section of the tube voltage waveform and a part of the falling portion of the tube voltage waveform, and a high-energy average value of radiation corresponding to a high-voltage section including a high steady-state section of the tube voltage waveform and another part of the falling portion of the tube voltage waveform, and set parameters used to create a radiographic image according to the low-energy average value and the high-energy average value. [Brief explanation of the drawings]

[0013] [Figure 1] 1 is a diagram illustrating a schematic configuration of an X-ray CT system according to an embodiment of the present invention. [Figure 2] 2 is a diagram showing the configuration of the main parts of an X-ray tube and an X-ray detection unit. FIG. [Figure 3] FIG. 10 is a diagram showing the energy distribution of an X-ray beam that changes depending on the tube voltage. [Figure 4] 10A and 10B are graphs showing changes in tube voltage over time and changes in tube current over time. [Figure 5] FIG. 10 is a diagram showing a tube voltage waveform. [Figure 6] FIG. 10 is a diagram showing a tube voltage waveform. [Figure 7] FIG. 10 is a diagram illustrating a method for obtaining a beam hardening correction coefficient to be applied. DETAILED DESCRIPTION OF THE INVENTION

[0014] Hereinafter, embodiments of the present invention will be described, but the present invention is not limited thereto.

[0015] 1 is a block diagram showing the configuration of an X-ray CT device 100 according to this embodiment. While a medical X-ray CT device will be described as an example in this disclosure, the present invention can also be applied to non-destructive testing devices such as dental CT devices and CT devices for baggage inspection. This X-ray CT device 100 includes an operation console 1, an imaging table 10, and a scanning gantry 20.

[0016] The operation console 1 has a computer configuration. Specifically, the operation console 1 includes an input device 2 such as a keyboard or a mouse that accepts input from an operator, a central processing unit 3 that executes scan control processing, preprocessing, image generation processing, etc., and a data acquisition buffer 5 that collects X-ray detector data acquired by the scan gantry 20. The operation console 1 also includes a monitor 6 that displays multi-energy images generated by the image generation processing, etc., and a storage device 7 that stores programs, X-ray detector data, X-ray projection data, dual-energy images, etc. The imaging conditions are input from the input device 2 and stored in the storage device 7.

[0017] The imaging table 10 is provided with a cradle 12 on which the subject 71 is placed and which is moved in and out of an opening 20a (described later) of the scan gantry 20. The cradle 12 is moved up and down and in a horizontal linear motion by a motor built in the imaging table 10.

[0018] The scanning gantry 20 has an opening 20a through which a subject 71 to be imaged is carried. The scanner gantry 20 also has an X-ray tube 21, an X-ray control unit 22 that controls the X-ray tube voltage, X-ray emission timing, etc. of the X-ray tube 21, and a collimator 23 having an aperture that shapes the X-rays emitted from the X-ray tube 21 into a fan-shaped X-ray beam 81. The scanner gantry 20 also has a collimator control unit 27 that controls the aperture of the collimator 23, an X-ray detector 24 that detects the X-rays emitted from the X-ray tube 21, and a data acquisition system (DAS) 25 that collects X-ray detector data (also called raw data) from the output of the X-ray detector 24. The DAS 25 samples analog data received from detector elements of the X-ray detector 24 and converts the analog data into a digital signal for subsequent processing.

[0019] The scanner gantry 20 further includes a gantry rotation unit 15 that holds the X-ray tube 21, collimator 23, and X-ray detector 24 and rotates around the body axis of the subject 71, and a rotation control unit 26 that controls the gantry rotation unit 15. The scanner gantry 20 also includes a gantry control unit 29 that exchanges control signals between the operation console 1 and the X-ray control unit 22, rotation control unit 26, and imaging table 10. In practice, the scanner gantry 20 includes a beam-forming X-ray filter that spatially controls the dose of the X-ray beam 81 and an X-ray filter that controls the radiation quality of the X-ray beam 81, and the gantry rotation unit 15 holds these filters between the collimator 23 and the opening 20a, but illustration and detailed description thereof are omitted here.

[0020] 2 is a diagram showing the configuration of the main parts of the X-ray tube 21 and the X-ray detection unit 24. Here, the vertical direction is defined as the y-axis direction, the direction of transport of the imaging table 10 (which usually coincides with the thickness direction of the X-ray beam 81 or the body axis direction of the subject 71) as the z-axis direction, and the direction perpendicular to the y-axis and z-axis directions (channel direction) as the x-axis direction.

[0021] These components are supported on a predetermined base of the gantry rotating unit 15 and maintain the positional relationship as shown in the figure. That is, the X-ray tube 21 and the X-ray detector 24 are arranged opposite each other with the opening 20a in between. Then, X-rays emitted from the X-ray tube 21 pass through a slit formed by the collimator 23 (not shown in Fig. 2), thereby forming a fan-shaped X-ray beam 81 having a predetermined thickness (cone angle) and spread (fan angle).

[0022] The X-ray tube 21 has a structure in which a cathode sleeve 21s incorporating a focusing electrode and a cathode filament, and a rotating target electrode 21t are housed in a housing 21h, and generates X-rays that diverge from an X-ray focal point F.

[0023] The X-ray detection unit 24 is a so-called multi-row X-ray detector, which is configured by arranging a plurality of, for example, 1,000 X-ray detection elements 24a in a channel direction CH (the direction in which the X-ray beam 81 spreads), and arranging a plurality of, for example, 64 detection element rows in the z-axis direction (the thickness direction of the X-ray beam 81). Here, the detection element rows are numbered 1, 2, 3, . . . , 64 from the end. This realizes a so-called 64-row multi-slice X-ray CT. However, the 64-row detection element row here is merely an example, and the present invention is not limited to this. The X-ray detector 24 forms an X-ray detection surface 24s, which detects the X-ray beam 81 transmitted through the subject 71, using the plurality of X-ray detection elements 24a. The X-ray detection elements 24a are configured as a so-called solid-state detector, for example, by combining a scintillator and a photodiode.

[0024] The central processing unit 3 has a scan control unit 32, a pre-processing unit 34, and an image generation unit 35. The central processing unit 3 is, for example, a processor such as a CPU (Central Processing Unit). The central processing unit 3 executes the functions of the scan control unit 32, the pre-processing unit 34, and the image generation unit 35 by reading and executing a program stored in the storage device 7. The program is an example of an embodiment of the control program according to the present invention.

[0025] The scan control unit 32 controls the X-ray control unit 22, the rotation control unit 26, the collimator control unit 27, and the imaging table 10 via the gantry control unit 29 so as to perform multi-energy imaging of the subject 71. Specifically, the scan control unit 32 controls the above-mentioned units to rotate the X-ray tube 21 and the X-ray detector 24 around the subject 71 and collect X-ray projection data.

[0026] In this embodiment, the target tube voltages are a first tube voltage V1 and a second tube voltage V2. For example, the first tube voltage V1 is 80 kV, and the second tube voltage V2 is 140 kV. In another embodiment, the first tube voltage V1 is 85 kV, and the second tube voltage V2 is 130 kV. In another embodiment, the first tube voltage V1 is 100 kV, and the second tube voltage V2 is 150 kV. Preferably, the first tube voltage V1 is 50 to 120 kV, and the second tube voltage V2 is 110 to 200 kV, which is higher than the first tube voltage V1. In this embodiment, the difference between the first tube voltage V1 and the second tube voltage V2 is 20 to 100 kV, and preferably 40 to 70 kV. In another embodiment, in addition to the first tube voltage V1 and the second tube voltage V2, a third tube voltage V3 is applied, with the third tube voltage V3 being a higher potential than the first tube voltage V1 and the second tube voltage V2. In this case, the influence of the fall time can be reduced by repeatedly applying a low voltage, a high voltage, and a medium voltage. Any number of tube voltages greater than two can be used, but to facilitate reader understanding, the following description will focus on an embodiment in which dual-energy imaging is performed using only two tube voltages.

[0027] The X-rays irradiated when the tube voltage of the X-ray tube 21 is the first tube voltage V1 are referred to as first X-rays, and the X-rays irradiated when the tube voltage is the second tube voltage V2 are referred to as second X-rays. Since the first X-rays and the second X-rays have different tube voltages, the first energy spectrum of the first X-rays and the second energy spectrum of the second X-rays are different spectra.

[0028] 3 shows the energy spectrum of X-rays irradiated when typical tube voltages of 100 kV, 150 kV, and 200 kV are applied to the X-ray tube 21. As shown in the figure, whether the tube voltage is 100 kV, 150 kV, or 200 kV, the energy of the irradiated X-ray radiation is distributed as shown in the figure as an energy lower than the applied kV.

[0029] In a preferred embodiment, the scan control unit 32 switches the X-ray tube voltage V every view via the X-ray control unit 22, and the X-ray detector 24 accordingly acquires X-ray projection data for the number of views required for image reconstruction (X-ray projection data for multiple views corresponding to 180° + fan angle α or 360°). As a result, first X-ray projection data p1 corresponding to the first tube voltage V1 and second X-ray projection data p2 corresponding to the second tube voltage V2 are acquired. In another preferred embodiment, the X-ray control unit 22 switches the X-ray tube voltage V every two views. Sampling for generating detection signals is performed a first number of times in each low voltage section and a second number of times in each high voltage section, and the first number of times and the second number of times may be the same or different.

[0030] The preprocessing unit 34 receives digital X-ray projection data from the DAS 25 and performs preprocessing on the X-ray projection data obtained by multi-energy imaging. Specifically, the preprocessing unit performs offset correction, logarithmic conversion, sensitivity correction to correct for sensitivity non-uniformity between channels in the raw data collected by the data acquisition unit 25, X-ray dose correction to correct for extreme signal intensity reduction or signal loss due to strong X-ray absorbers such as metal parts, X-ray scatter correction, X-ray beam hardening correction, and other preprocessing. As preprocessing, the first X-ray projection data p1 and the second X-ray projection data p2 are subjected to fan-parallel conversion, which converts fan-beam X-rays into parallel beam X-rays. For example, the fan-parallel conversion method disclosed in Japanese Patent Application Laid-Open No. 62-49831 can be used to prevent image degradation.

[0031] The image generating unit 35 generates a dual-energy image DE in which a specific material is emphasized or suppressed, based on the first X-ray projection data p1 and the second X-ray projection data p2 preprocessed by the preprocessing unit 34.

[0032] As a method for generating the dual-energy image DE, a method for performing weighted subtraction processing in image data space and a method for performing weighted subtraction processing in projection data space are conceivable, and either method may be adopted.

[0033] In a method for generating a dual-energy image DE by performing weighted subtraction processing in image data space, a first image P1 is reconstructed based on first X-ray projection data p1, a second image P2 is reconstructed based on second X-ray projection data p2, and weighted subtraction processing is performed between the first image P1 and the second image P2 to generate the dual-energy image DE.On the other hand, in a method for generating a dual-energy image DE by performing weighted subtraction processing in projection data space, weighted subtraction processing is performed between the first X-ray projection data p1 and the second X-ray projection data p2 on a view-by-view basis, and a dual-energy image DE is reconstructed based on the processed X-ray projection data obtained as a result.

[0034] Image reconstruction can be performed using, for example, a conventionally known three-dimensional image reconstruction method based on the Feldkamp method, other three-dimensional image reconstruction methods, or two-dimensional image reconstruction methods, and is performed, for example, by the following procedure. First, these X-ray projection data are subjected to a fast Fourier transform (FFT) to convert them into the frequency domain, and then multiplied by a reconstruction function Kernel(j) to perform an inverse Fourier transform. Then, a back projection process is performed on the X-ray projection data multiplied by the reconstruction function Kernel(j), and a tomographic image (xy plane) corresponding to the same slice when the subject 71 is sliced ​​in the body axis direction (z-axis direction) is obtained.

[0035] The operation of the X-ray CT apparatus 100 according to this embodiment will be described. The X-ray tube 21 and the X-ray detector 24 rotate around the body axis of the subject to collect projection data. At this time, the switching control unit 321 switches the X-ray tube voltage of the X-ray tube 21 between a first tube voltage V1 and a second tube voltage V2, as shown in graph G1 of FIG. 4. In graph G1, the horizontal axis represents time and the vertical axis represents tube voltage. As a result, the X-ray tube 21 alternately irradiates the subject with first and second X-rays while rotating around the body axis of the subject. The voltage values ​​of the first and second X-rays can be set by an operator using the input device 2.

[0036] 4, the switching control unit 321 switches the X-ray tube current A of the X-ray tube 21 between a first tube current A1 and a second tube current. The first tube current A1 is the tube current when the X-ray tube 21 has a first tube voltage V1, and the second tube current A2 is the tube current when the X-ray tube 21 has a second tube voltage V2. For example, the first tube current A1 is 300 mA, and the second tube current A2 is 100 mA.

[0037] When the X-ray tube 21 is at the first tube voltage V1, the switching control unit 321 reduces the X-ray tube current A to less than the first tube current A1 and then switches the tube voltage from the first tube voltage V1 to the second tube voltage V2. After switching to the second tube voltage, the switching control unit 321 sets the tube current A to the second tube current A2. Furthermore, when the X-ray tube 21 is at the second tube voltage V2, the switching control unit 321 reduces the X-ray tube current A to less than the second tube current A1 and then switches the tube current A from the first tube voltage V1 to the second tube voltage V2. After switching to the first tube voltage, the switching control unit 321 sets the tube current A to the first tube current A1. In FIG. 4, the timing of switching from the first tube voltage V1 to the second tube voltage V2 and the timing of switching from the second tube voltage V2 to the first tube voltage V1 (switching timing) are indicated by the symbol Ts.

[0038] In this example, the tube current A is zero at the switching timing Ts. However, the tube current A at the switching timing Ts does not have to be zero. For example, the tube current A at the switching timing Ts may be about 1 / 100 of the first tube current A1 and the second tube current A2.

[0039] Fig. 5 is a diagram showing the actual tube voltage applied to the X-ray tube 21. In Figs. 5 and 6, the horizontal axis represents time, and the vertical axis represents voltage (kV). When switching the tube voltage between the first tube voltage V1 and the second tube voltage V2, ideally the switching would be instantaneous as shown in Fig. 4. However, due to performance limitations of the transformer, rectifier, and the like, in reality, voltage rising and falling sections occur as shown in Fig. 5. Specifically, as shown in the diagrams, the tube voltage waveform includes a rising section in which the tube voltage rises from a low tube voltage to a high tube voltage, a falling section in which the tube voltage rises from a high tube voltage to a low tube voltage, a high steady section between the rising and falling sections, and a low steady section between the falling and rising sections.

[0040] In many CT devices, projection data of approximately 1000 views is collected per rotation, regardless of the rotation speed. Therefore, when the rotation speed changes, the sampling rate and the tube voltage switching period also change. FIG. 5 shows, as an example, tube voltage waveforms when the rotation speed is 1.0 second / rotation and when the rotation speed is 0.6 second / rotation. These tube voltage waveforms can be determined by applying an existing measurement device such as an oscilloscope to the X-ray tube 21. In other embodiments, the number of views per rotation is set to approximately 2000, approximately 500, or the like, rather than approximately 1000. Preferably, a fixed or variable number of views between 200 and 4000 is collected per rotation. The rotation speed and / or the number of views per rotation can be set by the operator using the input device 2.

[0041] As shown in Figure 5, when the rotation speed is 0.6 seconds / rotation, there is no significant change in the rising and falling parts compared to when it is 1.0 seconds / rotation, but the high and low steady-state periods become shorter. The same is true when the number of views per rotation increases.

[0042] Furthermore, the tube voltage waveform shown in FIG. 5 is affected by the tube current. Specifically, when the current value of the power applied to the X-ray tube 21 is small, the voltage falling section is relatively gentle, and when the current value is large, the voltage falling section is relatively steep. In contrast, the voltage rising section does not change significantly whether the current value is small or large. Such changes in the tube voltage waveform due to the tube current can also be identified by applying an existing measurement device such as an oscilloscope to the X-ray tube 21. In a preferred embodiment, a tube current of a predetermined value in the range of 20 mA to 1000 mA is applied to the X-ray tube 21. More preferably, a tube current of a predetermined value in the range of 50 mA to 500 mA is applied to the X-ray tube 21. The tube current can be increased or decreased in increments of 50 mA. In another embodiment, the tube current can be increased or decreased in increments of 10 mA. The values ​​of the tube current for the first X-ray and the second X-ray can be set by an operator using the input device 2.

[0043] Figure 6 shows a portion of a tube voltage waveform. In Figure 6, the voltage rising section starts at t0 and ends at t1. The high steady section starts at t1 and ends at t2. The voltage falling section starts at t2 and ends at t3. The low steady section starts at t3 and ends at t4. These four sections are repeated in the same order.

[0044] When an X-ray beam detection signal based on such a voltage waveform is sampled, as shown in FIG. 6 , sampling can be performed by defining the period from the start time between t0 and t1 to the end time between t2 and t3 as a high X-ray energy period (high tube voltage period Vh), and the period from the start time between t2 and t3, which starts simultaneously with the end of the high tube voltage period Vh, to the end time between t4 and t5 as a low X-ray energy period (low tube voltage period Vl). In a preferred embodiment, the high tube voltage period Vh and the low tube voltage period Vl have the same duration. In another embodiment, the duration of the high tube voltage period Vh is 0 to 20% longer than the duration of the low tube voltage period Vl. In another embodiment, the duration of the low tube voltage period Vl is 0 to 20% longer than the duration of the high tube voltage period Vh.

[0045] The timing of the start and end of the high-Tube voltage section Vh (corresponding to the end of the low-Tube voltage section Vl) shown in Figure 6 can be varied in various ways. However, starting the high-Tube voltage section Vh at time t0 and ending it at time t2 is not desirable because the high-Tube voltage section Vh includes a voltage portion close to the low V1, while the low-Tube voltage section Vl includes a voltage portion close to the high V2. It is desirable to include the high-voltage portion in both the voltage-rising section and the voltage-falling section in the high-Tube voltage section Vh, and the low-voltage portion in the low-Tube voltage section Vl. In the example of Figure 6, one sampling is performed in the high-Tube voltage section Vh and one sampling is performed in the low-Tube voltage section Vl. However, a similar problem occurs if two or more samplings are performed in the high-Tube voltage section Vh and two or more samplings are performed in the low-Tube voltage section Vl. Furthermore, starting the low tube voltage section Vl simultaneously with the end of the high tube voltage section Vh has the advantage of minimizing the radiation dose to the imaging subject, but problems with the voltage rise section and voltage fall section also arise when a short time interval is provided between the end of the high tube voltage section Vh and the start of the low tube voltage section Vl and / or between the end of the low tube voltage section Vl and the start of the high tube voltage section Vh.

[0046] In both the voltage rising section and the voltage falling section, the high-voltage portion is included in the high-Tube voltage section Vh, and the low-voltage portion is included in the low-Tube voltage section Vl. Therefore, the start time of the high-Tube voltage section Vh can be set to a predetermined delay time after the time t0 when the application of the second tube voltage V2 begins. In other embodiments, the start time of the high-Tube voltage section Vh can be set to the time when half the repetition period (or a predetermined time interval, such as 55% of the repetition period) has elapsed since the start time of the high-Tube voltage section Vh. In other embodiments, the end time of the high-Tube voltage section Vh can be set to the time when the tube voltage falls below a predetermined threshold. The timing of sampling performed by the DAS25 can be set based on the results of such automatic analysis of the voltage waveform.

[0047] As described above, as the rotation speed increases and / or the number of views per rotation increases, the steady-state section of the tube voltage waveform becomes shorter, and the voltage falling section becomes relatively gradual when the current value of the power applied to the X-ray tube 21 is small, and relatively steep when the current value is large. Therefore, regardless of how the start and end times of the high tube voltage section Vh are set, the energy of the first X-ray based on the tube voltage in the low tube voltage section Vl and the energy of the second X-ray based on the tube voltage in the high tube voltage section Vh vary over time. Therefore, the average energy, average energy, and average energy of the first X-ray and the average energy of the second X-ray also vary depending on the rotation speed, the number of views per rotation, and the tube current. Note that the "average" referred to in this specification does not necessarily mean a pure average of energy values, but may also be calculated with correction taking into account X-ray absorption characteristics or based on a representative value.

[0048] For example, if the first tube voltage V1 is 100 kV and the second tube voltage V2 is 150 kV, X-rays with the energy distribution shown for the 100 kV tube voltage in Figure 3 are irradiated at time t0 in Figure 6, and then the energy distribution rapidly changes to that shown for the 150 kV tube voltage in Figure 3 toward time t1. From time t1 to time t2, X-ray irradiation with the energy distribution shown for the 150 kV tube voltage in Figure 3 is maintained, and from time t2 to time t3, the energy distribution changes rapidly at first to that shown for the 100 kV tube voltage in Figure 3. Then, from time t3 to time t4, X-ray irradiation with the energy distribution shown for the 100 kV tube voltage in Figure 3 is maintained.

[0049] When the rotation speed increases or the number of views per rotation increases, the time interval between time t1 and time t2 and the time interval between time t3 and time t4 become shorter. When the rotation speed decreases or the number of views per rotation decreases, the time interval between time t1 and time t2 and the time interval between time t3 and time t4 become longer. Furthermore, regardless of the magnitude of the tube current, the time interval between time t0 and time t1 and the time interval between time t1 and time t2 do not change significantly. That is, regardless of the magnitude of the tube current, the tube voltage waveform exhibits a relatively steep rising curve between time t0 and time t1. When the magnitude of the tube current is large, the time interval between time t2 and time t3 becomes shorter, and the time interval between time t3 and time t4 becomes longer. On the other hand, when the magnitude of the tube current is small, the time interval between time t2 and time t3 becomes longer, and the time interval between time t3 and time t4 becomes shorter. That is, when the magnitude of the tube current is small, the tube voltage waveform forms a gentle descending curve between time t2 and time t3.

[0050] The inventors of the present invention have discovered that the above-mentioned changes in the tube voltage waveform have a pattern that can be modeled. Specifically, the tube voltage waveform is specified as follows: Time t1: A time point determined by a polynomial that is a function of the potential difference between the second tube voltage V2 and the first tube voltage V1. Times t2 and t4: points in time determined by the rotation speed and / or the number of views per rotation (for example, if the rotation speed is 1.0 second / rotation and 1000 views are collected per rotation, t2 will be 1 millisecond after t0 and t4 will be 2 milliseconds after t0; if the rotation speed is 1.0 second / rotation and 500 views are collected per rotation, t2 will be 2 milliseconds after t0 and t4 will be 4 milliseconds after t0). Time t3: A point in time determined by a polynomial that is a function of the potential difference between the second tube voltage V2 and the first tube voltage V1 and the magnitude of the tube current. Between time t0 and time t1: A curve specified by a polynomial that is a function of the potential difference between the second tube voltage V2 and the first tube voltage V1. If the potential difference between the second tube voltage V2 and the first tube voltage V1 is large, the time interval between time t0 and time t1 will be long. If the potential difference between the second tube voltage V2 and the first tube voltage V1 is small, the time interval between time t0 and time t1 will be short. Between time t1 and time t2: A straight line identified by the second tube voltage V2. Between time t2 and time t3: A curve specified by a polynomial that is a function of the potential difference between the second tube voltage V2 and the first tube voltage V1 and the magnitude of the tube current. If the potential difference between the second tube voltage V2 and the first tube voltage V1 is large, the time interval between time t2 and time t3 is long. If the potential difference between the second tube voltage V2 and the first tube voltage V1 is small, the time interval between time t2 and time t3 is short. If the tube current is large, the time interval between time t2 and time t3 is short. If the tube current is small, the time interval between time t2 and time t3 is long. Between time t3 and time t4: a straight line identified by the first tube voltage V1.

[0051] In this embodiment, the curved portion of the tube voltage waveform is modeled using a polynomial, but in other embodiments, it can also be modeled using a mathematical model that approximates it using a combination of a linear function, a quadratic function, and a cubic function.

[0052] The tube voltage waveform can also be modeled as an artificial intelligence (AI) model constructed using deep learning. During deep learning, a trained AI model can be constructed by repeatedly learning combinations of the first tube voltage V1, second tube voltage V2, tube current, rotation speed, and number of views / rotations, as well as the actual tube voltage waveforms measured under those combinations, as training data.

[0053] On the other hand, as described in Patent Document 3, beam hardening correction has been conventionally performed on image data collected by multi-energy imaging. Beam hardening correction is based on a phenomenon known as the "beam hardening effect," in which the average energy of X-rays emitted by a penetrating object shifts to higher energy values ​​in the case of polychromatic X-rays due to the spectral correlation of the ray attenuation performance of the actual object. In the reconstructed image of the object, linear, spectrally related ray attenuation can be observed via a shift in grayscale values ​​relative to the theoretical case. In particular, the shift in grayscale values ​​in the reconstructed image caused by high nuclear charge and high-density materials (such as bone) or beam hardening artifacts can cause the reconstructed image to be difficult to interpret correctly.

[0054] To reduce or eliminate this beam hardening effect, beam hardening correction is applied to image data acquired using multi-energy scanning. The beam hardening correction coefficient is used for this purpose. The beam hardening correction coefficient is determined by performing calibration, in which a phantom, such as a water phantom, is actually scanned using multi-energy scanning. During calibration, the CT value of water is set to 0, and the CT value of air is set to -1000. However, as mentioned above, the average energy of X-rays varies depending on the rotation speed, number of views per rotation, tube current, and potential difference between the second tube voltage V2 and the first tube voltage V1. Therefore, in the past, to correctly set the beam hardening correction coefficient, calibration was performed using a combination of the rotation speed, number of views per rotation, tube current, and potential difference between the second tube voltage V2 and the first tube voltage V1.

[0055] However, by modeling the tube voltage waveform as described above, it is possible to identify the corresponding tube voltage waveform by specifying the combination of the rotation speed, the number of views per rotation, the tube current, and the potential difference between the second tube voltage V2 and the first tube voltage V1. Once the tube voltage waveform is identified, it is possible to determine the average energy of the X-ray beam based on this, and obtain the beam hardening correction coefficient to be applied.

[0056] 7 is a diagram illustrating a method for obtaining the beam hardening correction coefficient to be applied. When the beam hardening correction coefficient is embodied as a beam hardening vector (BH vector) corresponding to a detector channel, the BH vector for an X-ray beam with an energy corresponding to a major tube voltage, such as 80 kVp, 100 kVp, 120 kVp, or 140 kVp, can be obtained using an existing method such as calibration imaging using a phantom.

[0057] Then, if the average energy of the X-ray beam obtained by modeling the tube voltage waveform is, for example, 90 kVp, the BH vector of 90 kVp can be obtained by using the average of the BH vector of 80 kVp and the BH vector of 100 kVp as shown in Figure 7. In another example, if the desired BH vector is a value (XkVp) between 120 kVp and 140 kVp, the BH vector of XkVp can be obtained by using the weighted average of the BH vector of 120 kVp and the BH vector of 140 kVp. In such a case, the value of the BH vector of XkVp is: BH vector value at XkVp = ((140-X) × 120kVp BH vector value + (X-120) × 140kVp BH vector value) / 20 (120 <X<140) It can be said that:

[0058] Furthermore, it is known to those skilled in the art that the energies of the scattered X-rays and recoil electrons generated by Compton scattering vary depending on the energy of the incident X-rays, and that the distribution of scattered X-rays also varies depending on the energy of the incident X-rays. Scatter correction has been performed to eliminate or reduce the effects of such scattered X-rays. This scatter correction is particularly important in dental CT systems and CT systems used for baggage inspection, where the scanned object is likely to contain metals. The coefficients used in this scatter correction can also be calculated based on the average value of the X-ray beam calculated from the tube voltage waveform model. The average value of the X-ray beam can also be used to calculate the absorption value μ of the basis material pair used in material decomposition calculations.

[0059] Furthermore, the rate of X-ray absorption varies depending on the type and concentration of elements that make up the imaging target, and the presence of high-density materials with high atomic numbers reduces the amount of X-rays received by the detector. Those skilled in the art also know that the degree of absorption varies depending on the energy of the incident X-rays. Absorption correction has traditionally been performed to eliminate or reduce the effects of such absorbed X-rays. This absorption correction is particularly important in dental CT systems and CT systems used for baggage inspection, where the scanned object is likely to contain metals. The coefficients used in this absorption correction can also be calculated based on the average value of the X-ray beam calculated from the tube voltage waveform model.

[0060] The beam hardening correction coefficient, the scatter correction coefficient, and the attenuation correction coefficient can be generated for each of the plurality of detector elements 24a or for a group of the plurality of detector elements 24a. The image generation unit 35 can also reconstruct an image using the difference between the beam hardening correction coefficient calculated by the preprocessing unit 34 from the tube voltage waveform model and the beam hardening correction coefficient calculated by imaging a phantom.

[0061] In this way, by setting the parameters used to create radiographic images according to the low-energy average value and high-energy average value calculated from the tube voltage waveform model, it is possible to shorten the time required for calibration imaging and calculation, which in turn reduces the downtime of the CT system and reduces service costs.

[0062] In a preferred embodiment of the present invention, a function for updating a tube voltage waveform model is provided. Repeated use of the CT device or aging can cause the shape of the tube voltage waveform to change due to factors such as the waveform of the voltage output by the power supply or deterioration of the target of the X-ray tube 21. Even if the tube voltage waveform remains unchanged, the energy distribution of the X-ray beam irradiated based on this may change. In such cases, the tube voltage waveform model generated at the time of shipment or initial use may no longer be compatible with the current CT device. In such cases, the tube voltage waveform model is corrected based on projection data measured during calibration.

[0063] The invention is not limited to the present embodiment, and various modifications are possible within the scope of the invention.

[0064] Furthermore, a program for causing a computer to function as each of the means for controlling and processing the X-ray CT apparatus is also an example of an embodiment of the invention. [Explanation of symbols]

[0065] 1: Operation console 2: Input device 3: Central processing unit 5: Data acquisition buffer 6: Monitor 7: Storage device 10: Examination table 12: Cradle 15: Gantry rotation unit 20: Scanning gantry 20a: Opening 21: X-ray tube 21h: Housing 21s: Cathode sleeve 21t: Target electrode 22: X-ray control unit 23: Collimator 24: X-ray detector 24a: X-ray detection element 24s: X-ray detection surface 25: DAS 26: Rotation control unit 27: Collimator control unit 29: Gantry control unit 32: Scan control unit 34: Pre-processing unit 35: Image generation unit 71: Subject / imaging target 81: X-ray beam 100: X-ray CT device

Claims

1. A system including a processor that processes a detection signal of radiation that has passed through an imaging object to generate a radiation image, the radiation is radiation that is irradiated toward the imaging target while a radiation tube rotates around the imaging target, the radiation includes low-energy radiation generated by applying a low tube voltage to the radiation tube and high-energy radiation generated by applying a high tube voltage to the radiation tube; application of the low tube voltage and the high tube voltage to the radiation tube are alternately switched during rotation of the radiation tube, to form a tube voltage waveform having a rising portion from the low tube voltage to the high tube voltage, a falling portion from the high tube voltage to the low tube voltage, a high steady section between the rising portion and the falling portion, and a low steady section between the falling portion and the rising portion; The processor: receiving input for rotation speed and / or number of views per rotation; identifying a tube voltage waveform corresponding to the rotation speed and / or the number of views per rotation using a waveform identification model; calculating a low-energy average value of the radiation corresponding to a low-voltage section including a low steady-state section of the tube voltage waveform and a part of the falling portion of the tube voltage waveform, and a high-energy average value of the radiation corresponding to a high-voltage section including a high steady-state section of the tube voltage waveform and another part of the falling portion of the tube voltage waveform; setting parameters used to create the radiation image in accordance with the low-energy average value and the high-energy average value; system.

2. a table on which the imaging target is placed; a gantry that rotatably supports the radiation tube and a detector that detects the radiation that has passed through the imaging target; a storage medium that stores the waveform identification model; a user interface for inputting the rotation speed and / or the number of views per rotation; an image reconstruction device comprising the processor; The system of claim 1 , comprising:

3. The system of claim 1 , wherein the parameters include a beam hardening correction factor.

4. The system of claim 1 , wherein the parameters include an X-ray absorption coefficient and / or an X-ray scattering coefficient.

5. the low-voltage section includes a part of the rising portion of the tube voltage waveform, the high-voltage section includes another part of the rising portion of the tube voltage waveform, 2. The system of claim 1, wherein the waveform specification model makes the voltage falling portion relatively gradual when a current value of the power applied to the radiation tube is low and makes the voltage falling portion relatively steep when the current value is high, while not significantly changing the voltage rising portion whether the current value is low or high.

6. the tube voltage waveform is collected by measuring power applied to the radiation tube; The system of claim 1 , wherein in the waveform identification model, a curve of a voltage falling portion of a measured waveform is expressed by an approximation of a polynomial.

7. The pre-processor is Separating the tube voltage waveform into the low voltage section and the high voltage section using a predetermined threshold; sampling for generating the detection signal is performed a first number of times in each of the low voltage sections and a second number of times in each of the high voltage sections; The system of claim 6 , wherein the first number of times is the same as the second number of times.

8. the detector includes a plurality of detector elements extending in a circumferential direction of the gantry; the radiation tube irradiates the detector with a fan beam or a cone beam diverging in the circumferential direction as the radiation; the detector detects the radiation transmitted through the phantom to generate a phantom detection signal; generating a first phantom-based beam-hardening correction factor corresponding to the low-energy radiation based on the phantom detection signal; the first phantom-based beam-hardening correction factor is generated for each of the plurality of detector elements; the parameters include a first model-based beam-hardening correction factor corresponding to the low voltage interval; The system of claim 7 , wherein the radiological image is reconstructed using a first difference between the first model-based beam-hardening correction coefficient and the first phantom-based beam-hardening correction coefficient.

9. generating a second phantom-based beam-hardening correction factor corresponding to the high-energy radiation based on the phantom detection signal; the second phantom-based beam-hardening correction factor is generated for each of the plurality of detector elements; the parameters include a second model-based beam hardening correction factor corresponding to a previous high and low voltage interval; The system of claim 8 , wherein the radiological image is reconstructed using a second difference between the second model-based beam-hardening correction coefficient and the second phantom-based beam-hardening correction coefficient.

10. A program for processing a detection signal of radiation that has passed through an imaging target to generate a radiation image, the radiation is radiation that is irradiated toward the imaging target while a radiation tube rotates around the imaging target, the radiation includes low-energy radiation generated by applying a low tube voltage to the radiation tube and high-energy radiation generated by applying a high tube voltage to the radiation tube; application of the low tube voltage and the high tube voltage to the radiation tube are alternately switched during rotation of the radiation tube, to form a tube voltage waveform having a rising portion from the low tube voltage to the high tube voltage, a falling portion from the high tube voltage to the low tube voltage, a high steady section between the rising portion and the falling portion, and a low steady section between the falling portion and the rising portion; The program receiving input for rotation speed and / or number of views per rotation; identifying a tube voltage waveform corresponding to the rotation speed and / or the number of views per rotation using a waveform identification model; calculating a low-energy average value of the radiation corresponding to a low-voltage section including a low steady-state section of the tube voltage waveform and a part of the falling portion of the tube voltage waveform, and a high-energy average value of the radiation corresponding to a high-voltage section including a high steady-state section of the tube voltage waveform and another part of the falling portion of the tube voltage waveform; a program causing a processor to set parameters used to create the radiation image in response to the low-energy average value and the high-energy average value;

11. The imaging target is placed on a table, the radiation transmitted through the imaging target is detected by a detector; the detector and the radiation tube are rotatably supported by a gantry; the waveform-specific model is pre-stored in a storage medium; The program according to claim 10 , wherein the rotation speed and / or the number of views per rotation are input via a user interface.

12. The program of claim 10 , wherein the parameters include a beam hardening correction factor.

13. The program according to claim 10 , wherein the parameters include an X-ray absorption coefficient and / or an X-ray scattering coefficient.

14. the low-voltage section includes a part of the rising portion of the tube voltage waveform, the high-voltage section includes another part of the rising portion of the tube voltage waveform, 11. The program according to claim 10, wherein the waveform specification model makes the voltage falling portion relatively gradual when a current value of the power applied to the radiation tube is low and makes the voltage falling portion relatively steep when the current value is high, while not significantly changing the voltage rising portion whether the current value is low or high.

15. the tube voltage waveform is collected by measuring power applied to the radiation tube; The program according to claim 11 , wherein in the waveform identification model, a curve of a voltage falling portion of the measured waveform is expressed by an approximate polynomial.

16. the program causes the processor to separate the tube voltage waveform into the low-voltage section and the high-voltage section using a predetermined threshold; sampling for generating the detection signal is performed a first number of times in each of the low voltage sections and a second number of times in each of the high voltage sections; The program of claim 15 , wherein the first number of times is the same as the second number of times.

17. the detector includes a plurality of detector elements extending in a circumferential direction of the gantry; the radiation tube irradiates the detector with a fan beam or a cone beam diverging in the circumferential direction as the radiation; the detector detects the radiation transmitted through the phantom to generate a phantom detection signal; generating a first phantom-based beam-hardening correction factor corresponding to the low-energy radiation based on the phantom detection signal; the first phantom-based beam-hardening correction factor is generated for each of the plurality of detector elements; the parameters include a first model-based beam-hardening correction factor corresponding to the low voltage interval; 17. The program of claim 16, wherein the radiological image is reconstructed using a first difference between the first model-based beam hardening correction coefficient and the first phantom-based beam hardening correction coefficient.

18. generating a second phantom-based beam-hardening correction factor corresponding to the high-energy radiation based on the phantom detection signal; the second phantom-based beam-hardening correction factor is generated for each of the plurality of detector elements; the parameters include a second model-based beam hardening correction factor corresponding to a previous high and low voltage interval; 18. The program of claim 17, wherein the radiation image is reconstructed using a second difference between the second model-based beam hardening correction coefficient and the second phantom-based beam hardening correction coefficient.

19. A non-transitory storage medium storing the program according to any one of claims 10 to 18.

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