Program and system for processing detection signal of radiation transmitting imaging object and generating radiation image

Abstract

[Problem] To provide a program and system capable of shortening calibration while maintaining the image quality of a reconstructed image. [Solution] In one embodiment, there is provided a system comprising a processor that processes a detection signal of radiation that has passed through an imaging subject to generate a radiation image. Applying a low tube voltage and applying a high tube voltage to the radiation tube are alternately switched in rotation of the radiation tube to form a tube voltage waveform, the tube voltage waveform has a rising portion from a low tube voltage to a high tube voltage, a falling portion from the high tube voltage to the low tube voltage, a high stability section between the rising portion and the falling portion, and a low stability section between the falling portion and the rising portion. The processor receives an input of a rotational speed and / or a number of views per revolution, determines a tube voltage waveform corresponding to the rotational speed and / or the number of views per revolution using a waveform determination model, calculating a low-energy average value of radiation corresponding to a low-stability section of the tube voltage waveform and a low-voltage section including a portion of a drop portion of the tube voltage waveform, and a high-energy average value of radiation corresponding to a high-stability section of the tube voltage waveform and a high-voltage section including another portion of the drop portion of the tube voltage waveform; and sets parameters for creating the radiation image on the basis of the low-energy average value and the high-energy average value.

Description

Technical Field

[0001] The present invention relates to a radiation imaging device, in particular to a device that realizes multi-energy

[0002] (multi-energy) photography technology. Background Art

[0003] As an imaging method using a radiation imaging apparatus represented by an X-ray CT apparatus (X-ray Computed Tomography System), a known imaging method is called dual-energy imaging, which switches the radiation tube voltage to perform data collection (for example, see Patent Documents 1 and 2).

[0004] In the case of X-ray CT systems, this imaging method utilizes the fact that the absorption spectrum of X-ray energy varies depending on the material to produce an image that emphasizes or suppresses (removes) specific substances in the subject. Specifically, for example, a subject composed of a first substance and a second substance is irradiated with first and second X-rays having different energy spectra, and first and second X-ray projection data corresponding to multiple views are collected.

[0005] Then, the first density image is reconstructed based on the first X-ray projection data, and the second density image is reconstructed based on the second X-ray projection data, so as to obtain a density distribution image representing the density distribution of a certain substance, or a monochrome image as an image of a certain energy spectrum, i.e., a dual-energy image, based on the first density image and the second density image.

[0006] 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 the 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, that is, each time the gantry of the X-ray CT apparatus rotates by one view.

[0007] Prior art literature

[0008] Patent Literature

[0009] Patent Document 1: Japanese Patent Application Laid-Open No. 2004-65975

[0010] Patent Document 2: Japanese Patent Application Laid-Open No. 2009-153829

[0011] Patent Document 3: Japanese Patent Application Laid-Open No. 2015-112478 Summary of the Invention

[0012] Problems to be solved by the invention

[0013] As described above, when collecting projection data while alternating the X-ray tube voltage between a low tube voltage and a high tube voltage, instantaneous switching from low tube voltage to high tube voltage and vice versa is desirable. However, due to transformer performance limitations, there are rise times from low tube voltage to high tube voltage and fall times from high tube voltage to low tube voltage. These rise and fall times of the tube voltage waveform 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 and the magnitude of the tube current.

[0014] Regarding tube current, the rise time remains largely unchanged regardless of whether the tube current is high or low. Meanwhile, the fall time is significantly affected compared to the rise time. For example, at a tube current of 400 mA, switching from a high tube voltage to a low tube voltage occurs relatively quickly, but at a tube current of 200 mA, the switch takes significantly longer. In other words, the fall time is relatively short at a tube current of 400 mA, while it is relatively short at a tube current of 200 mA. Furthermore, the rise and fall times often form curves rather than straight lines. When data is continuously sampled, that is, when the curved portion of the tube voltage waveform is included in the sampling, the energy spectrum of the irradiated X-ray beam changes, and the average energy of the sampled X-ray beam also changes.

[0015] Furthermore, when the number of collected views remains constant, changes in the rotation speed cause changes in the sampling rate, thus changing the period of switching from the high tube voltage to the low tube voltage. Consequently, the duration of the high stable range of the tube voltage waveform corresponding to the high tube voltage and the duration of the low stable range of the tube voltage waveform corresponding to the low tube voltage fluctuate. Specifically, as the rotation speed increases, the duration of both the high stable range and the low stable range decreases, while as the rotation speed decreases, the duration of both the high stable range and the low stable range increases.

[0016] On the other hand, to address variations in sensitivity between detector elements and deviations in the actual X-ray energy spectrum from the ideal spectrum, calibration is performed according to a pre-set image reconstruction protocol. As mentioned above, changes in the rotational speed or tube current cause variations in the energy spectrum of the emitted X-rays, necessitating calibration tailored to the type of rotational speed and tube current used. This necessity results in extended downtime of the X-ray CT system, increased imaging and calculation time for the complex calibration process, and increased service costs.

[0017] Therefore, there is a need for a new method that can shorten the imaging and calculation time for calibration while maintaining the image quality of the reconstructed image, reduce the downtime of the radiation imaging apparatus, and reduce the service cost.

[0018] Technical Solution

[0019] In the present disclosure, one or more of the aforementioned problems or problems related thereto are at least partially solved by a system including a processor that processes detection signals of radiation transmitted through an imaging object to generate a radiographic image. Here, the radiation is radiation emitted by a radiation tube while rotating around the imaging object, and 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 highly stable interval between the rising portion and the falling portion, and a less stable interval between the falling portion and the rising portion. The processor receives an input of a rotational speed and / or a number of views per rotation, determines a tube voltage waveform corresponding to the rotational speed and / or the number of views per rotation using a waveform determination model, calculates a low energy average value of the radiation corresponding to a low voltage interval including a low stable interval of the tube voltage waveform and a portion of a falling portion of the tube voltage waveform, and a high energy average value of the radiation corresponding to a high voltage interval including a high stable interval of the tube voltage waveform and another portion of the falling portion of the tube voltage waveform, and sets parameters for producing a radiographic image based on the low energy average value and the high energy average value.

[0020] In another aspect of the present disclosure, the above-identified problems or one or more problems related thereto are at least partially solved by a program for processing detection signals of radiation transmitted through an imaging object to generate a radiographic image. Here, the above-mentioned radiation is radiation irradiated toward an imaging object by a radiation tube while rotating around the imaging object. 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 application of 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 stability interval between the rising portion and the falling portion, and a low stability interval between the falling portion and the rising portion. The program further causes the processor to perform the following steps: receiving an input of a rotation speed and / or a number of views per rotation, determining a tube voltage waveform corresponding to the rotation speed and / or the number of views per rotation using a waveform determination model, calculating a low energy average value of radiation corresponding to a low stable interval of the tube voltage waveform and a low voltage interval including a portion of a falling portion of the tube voltage waveform, and a high energy average value of radiation corresponding to a high stable interval of the tube voltage waveform and a high voltage interval including another portion of the falling portion of the tube voltage waveform, and setting parameters for producing a radiographic image based on the low energy average value and the high energy average value. BRIEF DESCRIPTION OF THE DRAWINGS

[0021] Figure 1 It is a diagram schematically showing the configuration of an X-ray CT system according to this embodiment.

[0022] Figure 2 It is a diagram showing the configuration of main parts of an X-ray tube and an X-ray detection unit.

[0023] Figure 3 This is a diagram showing the energy distribution of the X-ray beam that changes according to the tube voltage.

[0024] Figure 4 It is a diagram showing a graph of the time change of the tube voltage and a graph of the time change of the tube current.

[0025] Figure 5 This is a diagram showing the tube voltage waveform.

[0026] Figure 6 This is a diagram showing the tube voltage waveform.

[0027] Figure 7 This is a diagram explaining a method for obtaining a beam hardening correction coefficient to be applied. DETAILED DESCRIPTION

[0028] Hereinafter, embodiments of the invention will be described, but it should be noted that the invention is not limited thereto.

[0029] Figure 1 This is a block diagram showing the structure of an X-ray CT system 100 in this embodiment. While this disclosure uses a medical X-ray CT system as an example, the present invention is also applicable to non-destructive inspection systems such as dental CT systems and luggage inspection systems. The X-ray CT system 100 includes an operation console 1, an imaging table 10, and a gantry 20.

[0030] The operation console 1 is configured as a computer. Specifically, the operation console 1 includes an input device 2, such as a keyboard or mouse, which receives input from the operator; a central processing unit 3, which performs scan control processing, preprocessing, image generation processing, and the like; and a data collection buffer 5, which collects X-ray detector data collected by the gantry 20. Furthermore, the operation console 1 includes a monitor 6, which displays multi-energy images generated by the image generation processing; and a storage device 7, which stores programs, X-ray detector data, X-ray projection data (projection data), dual-energy images, and the like. Radiographic conditions are input from the input device 2 and stored in the storage device 7.

[0031] The imaging table 10 includes a cradle 12 on which the subject 71 is placed and fed into or taken out of an opening 20a (described later) of the gantry 20. The cradle 12 is moved vertically and horizontally by a motor built into the imaging table 10.

[0032] The gantry 20 has an opening 20a for conveying a subject 71 to be imaged. The gantry 20 also includes an X-ray tube 21; an X-ray control unit 22 that controls the X-ray tube voltage and X-ray emission timing in the X-ray tube 21; and a collimator 23 with an opening that shapes the X-rays emitted from the X-ray tube 21 into a fan-shaped X-ray beam 81. Furthermore, the gantry 20 includes a collimator control unit 27 that controls the opening 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 the analog data received from the detector elements of the X-ray detector 24 and converts the analog data into digital signals for subsequent processing.

[0033] The gantry 20 also includes a gantry rotation unit 15 that holds the X-ray tube 21, collimator 23, and X-ray detector 24 and rotates about the body axis of the subject 71; and a rotation control unit 26 that controls the gantry rotation unit 15. Furthermore, the gantry 20 includes a gantry control unit 29 that exchanges control signals between the operation console 1 and the X-ray control unit 22, the rotation control unit 26, and the radiographic table 10. It should be noted that in actual use, the 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 quality of the X-ray beam 81. The gantry rotation unit 15 holds these filters between the collimator 23 and the opening 20a, but illustration and detailed description are omitted here.

[0034] Figure 2 This figure shows 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 transport direction of the imaging table 10 (usually coinciding with the thickness direction of the X-ray beam 81 or the body axis direction of the subject 71) is defined as the z-axis direction, and the direction perpendicular to the y-axis and z-axis directions (channel direction) is defined as the x-axis direction.

[0035] These components are supported on a predetermined base of the gantry rotating portion 15 so as to maintain the positional relationship shown in the figure. That is, the X-ray tube 21 and the X-ray detector 24 are arranged to face each other across the opening 20a. Then, the X-rays emitted from the X-ray tube 21 are transmitted by the X-ray detector 24. Figure 2 The slit formed by the collimator 23 (not shown) forms a fan-shaped X-ray beam 81 having a predetermined thickness (cone angle) and spread (fan angle).

[0036] The X-ray tube 21 has a structure in which a cathode sleeve 21 s having a focusing electrode and a cathode filament built therein and a rotating target electrode 21 t are housed in a housing 21 h , and generates X-rays diverging from an X-ray focal point F.

[0037] The X-ray detection unit 24 is a so-called multi-column X-ray detector that is composed of a plurality of, for example, 64 detection element columns arranged in the z-axis direction (the thickness direction of the X-ray beam 81). The detection element column is composed of a plurality of, for example, 1,000 X-ray detection elements 24a arranged in the channel direction CH (the expansion direction of the X-ray beam 81). Here, each detection element column is numbered 1, 2, 3, ..., 64 from the end. In this way, the so-called 64-column multi-slice X-ray CT is realized. However, here, the 64-column detection element column is an example, and the present invention is not limited to this. The X-ray detector 24 forms an X-ray detection surface 24s that detects the X-ray beam 81 that has passed through the subject 71 through these multiple X-ray detection elements 24a. The X-ray detection element 24a is composed of a so-called solid-state detector, for example, by a combination of a scintillator and a photodiode.

[0038] The central processing unit 3 includes a scan control unit 32, a preprocessing unit 34, and an image generator 35. The central processing unit 3 is, for example, a processor such as a CPU (Central Processing Unit). The central processing unit 3 performs the functions of the scan control unit 32, the preprocessing unit 34, and the image generator 35 by reading and executing a program stored in the storage device 7. The above program is an example of an implementation of the control program involved in the present invention.

[0039] 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 to perform multi-energy imaging of the subject 71. Specifically, the scan control unit 32 controls each of the above components to rotate the X-ray tube 21 and the X-ray detector 24 around the subject 71 and collect X-ray projection data.

[0040] In this embodiment, the target tube voltages are the first tube voltage V1 and the second tube voltage V2. For example, the first tube voltage V1 is 80 kV and the second tube voltage V2 is 140 kV. In other embodiments, the first tube voltage V1 is 85 kV and the second tube voltage V2 is 130 kV. In other embodiments, 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 kV to 120 kV, and the second tube voltage V2 is 110 kV to 200 kV 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 kV to 100 kV, preferably 40 kV to 70 kV. In other embodiments, in addition to applying the first tube voltage V1 and the second tube voltage V2, a third tube voltage V3 is applied. The third tube voltage V3 is a higher potential than the first tube voltage V1 and the second tube voltage V2. In this case, the effects of the fall time can be mitigated by repeatedly applying a low voltage, a high voltage, and a medium voltage. The number of tube voltage types can be any number greater than two, but to facilitate understanding, the following description focuses on an embodiment in which dual-energy imaging is performed using only two tube voltage types.

[0041] X-rays emitted when the tube voltage of the X-ray tube 21 is the first tube voltage V1 are referred to as first X-rays, and X-rays emitted when the tube voltage is the second tube voltage V2 are referred to as second X-rays. Since the tube voltages of the first X-rays and the second X-rays are different, the first energy spectrum of the first X-rays and the second energy spectrum of the second X-rays are different spectra.

[0042] Figure 3 The figure shows the energy spectrum of X-rays emitted when representative tube voltages of 100 kV, 150 kV, and 200 kV are applied to the X-ray tube 21. As shown in the figure, when the tube voltage is 100 kV, 150 kV, or 200 kV, the energy of the emitted X-rays is distributed as shown, being lower than the applied kV.

[0043] In a preferred embodiment, the scan control unit 32 switches the X-ray tube voltage V for each view via the X-ray control unit 22. Accordingly, the X-ray detector 24 collects X-ray projection data for the number of views required for image reconstruction (corresponding to X-ray projection data for multiple views, equivalent to 180° + fan angle α, or 360°). This collects 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. In other preferred embodiments, the X-ray control unit 22 switches the X-ray tube voltage V for every two views. Sampling to generate detection signals is performed a first time in each low-voltage interval and a second time in each high-voltage interval. The first and second times may be the same or different.

[0044] The preprocessing unit 34 receives digital X-ray projection data from the DAS 25 and preprocesses the X-ray projection data obtained through multi-energy imaging. Specifically, the preprocessing unit 34 performs preprocessing on the first X-ray projection data p1 and the second X-ray projection data p2, including offset correction, logarithmic transformation, sensitivity correction to correct for sensitivity unevenness between channels in the raw data collected by the data acquisition device 25, X-ray dose correction to correct for extreme signal intensity reduction or signal loss caused by strong X-ray absorbers such as metal parts, X-ray scatter correction, and X-ray beam hardening correction. Furthermore, as preprocessing, fan-to-parallel conversion is performed on the first X-ray projection data p1 and the second X-ray projection data p2, converting fan beam X-rays into parallel beam X-rays. For example, the fan-to-parallel conversion method disclosed in Japanese Patent Application Laid-Open No. 62-49831 can prevent image degradation.

[0045] The image generator 35 generates a dual energy image DE in which a specific substance is emphasized or suppressed, based on the first X-ray projection data p1 and the second X-ray projection data p2 preprocessed by the preprocessor 34 .

[0046] As a method of generating the dual energy image DE, a method of performing weighted subtraction processing in the image data space and a method of performing weighted subtraction processing in the projection data space are conceivable, but either method may be adopted.

[0047] 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, and a second image P2 is reconstructed based on second X-ray projection data p2, and a weighted subtraction processing is performed between the first image P1 and the second image P2 to generate the dual-energy image DE. Alternatively, in a method for generating a dual-energy image DE by performing weighted subtraction processing in projection data space, a weighted subtraction processing is performed on a view-by-view basis between the first X-ray projection data p1 and the second X-ray projection data p2, and the dual-energy image DE is reconstructed based on the processed X-ray projection data obtained as a result of the weighted subtraction processing.

[0048] Image reconstruction can be performed using, for example, a 3D image reconstruction method based on the conventionally known Feldkamp method, other 3D image reconstruction methods, or 2D image reconstruction methods, for example, through the following steps. First, a Fast Fourier Transform (FFT) is applied to the X-ray projection data to convert it into the frequency domain. This data is then multiplied by a reconstruction function, Kernel(j), and then inverse Fourier transformed. The X-ray projection data obtained by multiplying the data with the reconstruction function, Kernel(j), is then back-projected to obtain a tomographic image (xy plane) corresponding to the slice of the subject 71 when it is sliced ​​along the body axis (z-axis).

[0049] 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. Figure 4 As shown in graph G1, the switching control unit 321 switches the X-ray tube voltage of the X-ray tube 21 between the first tube voltage V1 and the second tube voltage V2. In graph G1, the horizontal axis represents time, and the vertical axis represents tube voltage. Consequently, the X-ray tube 21 rotates about the subject's body axis while alternately emitting the first and second X-rays. The voltages of the first and second X-rays can be set by the operator using the input device 2.

[0050] like Figure 4 As shown in graph G2, 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 of the X-ray tube 21 when the first tube voltage V1 is applied, and the second tube current A2 is the tube current of the X-ray tube 21 when the second tube voltage V2 is applied. For example, the first tube current A1 is 300 mA, and the second tube current A2 is 100 mA.

[0051] When the X-ray tube 21 is in the state of the first tube voltage V1, the switching control unit 321 switches from the first tube voltage V1 to the second tube voltage V2 after making the X-ray tube current A lower than the first tube current A1. Then, after switching to the second tube voltage, the switching control unit 321 sets the tube current A to the second tube current A2. In addition, when the X-ray tube 21 is in the state of the second tube voltage V2, the switching control unit 321 switches from the first tube voltage V1 to the second tube voltage V2 after making the X-ray tube current A lower than the second tube current A1. Then, after switching to the first tube voltage, the switching control unit 321 sets the tube current A to the first tube current A1. Figure 4 In FIG. 1 , 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 represented by symbol Ts.

[0052] In this example, the tube current A is zero at the switching timing Ts. However, the tube current A at the switching timing Ts may not be zero. For example, the tube current A at the switching timing Ts may be approximately 1% of the first tube current A1 and the second tube current A2.

[0053] Figure 5 Graph showing the actual tube voltage applied to the X-ray tube 21. Figure 5 、 Figure 6 In FIG. 1 , the horizontal axis represents time and the vertical axis represents voltage (kV). When the tube voltage is switched between the first tube voltage V1 and the second tube voltage V2, it is ideal to perform the following operation: Figure 4 Such instantaneous switching, but the performance of transformers, rectifiers, etc. has its limits, so in practice, Figure 5 Specifically, as shown in the figure, the tube voltage waveform includes 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 highly stable portion between the rising and falling portions, and a less stable portion between the falling and rising portions.

[0054] In many CT devices, projection data of approximately 1000 views are collected per rotation regardless of the rotation speed. Therefore, when the rotation speed changes, the sampling rate and the switching cycle of the tube voltage also change. Figure 5, as an example, tube voltage waveforms for rotation speeds of 1.0 s / rev and 0.6 s / rev are shown. 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 not approximately 1000, but is set to approximately 2000, approximately 500, or the like. 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.

[0055] like Figure 5 As shown, when the rotation speed is 0.6 sec / rev, the rising and falling parts do not change significantly compared to the case of 1.0 sec / rev, but the high and low stable intervals are shorter. The same is true when the number of views per rotation increases.

[0056] and then, Figure 5 The tube voltage waveform shown 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 decreases more gradually, while when the current value is large, the voltage decreases more steeply. In contrast, the voltage increase does not change significantly regardless of whether the current value is small or large. Such changes in the tube voltage waveform caused by the tube current can also be determined by applying an existing measuring device such as an oscilloscope to the X-ray tube 21. In a preferred embodiment, a tube current of a predetermined value within the range of 20mA to 1000mA is applied to the X-ray tube 21. More preferably, a tube current of a predetermined value within the range of 50mA to 500mA is applied to the X-ray tube 21. The tube current can be switched in increments of 50mA. In other embodiments, the tube current can be switched in increments of 10mA. The values ​​of the tube current for the first X-ray and the second X-ray can be set by the operator using the input device 2.

[0057] Figure 6 A portion of the tube voltage waveform is shown. Figure 6 In the example, the voltage rising interval begins at t0 and ends at t1. The high stability interval begins at t1 and ends at t2. The voltage falling interval begins at t2 and ends at t3. The low stability interval begins at t3 and ends at t4. Then, these four intervals are repeated in the same order.

[0058] When sampling the detection signal of the X-ray beam based on such a voltage waveform, as shown in FIG. Figure 6As shown, sampling can be performed by setting the high X-ray energy interval (high tube voltage interval Vh) from the start time between t0 and t1 to the end time between t2 and t3, and setting the low X-ray energy interval (low tube voltage interval Vl) from the start time between t2 and t3, which starts simultaneously with the end of the high tube voltage interval Vh, to the end time between t4 and t5. In a preferred embodiment, the high tube voltage interval Vh and the low tube voltage interval Vl have the same duration. In other embodiments, the duration of the high tube voltage interval Vh is 0 to 20% longer than the duration of the low tube voltage interval Vl. In other embodiments, the duration of the low tube voltage interval Vl is 0 to 20% longer than the duration of the high tube voltage interval Vh.

[0059] Able to make various changes Figure 6 The timing at which the high-voltage section Vh starts (corresponding to the end of the low-voltage section Vl) and the timing at which the high-voltage section Vh ends (corresponding to the start of the low-voltage section Vl) are shown, and are implemented. However, when the high-voltage section Vh starts at time t0 and ends at time t2, the high-voltage section Vh includes a portion whose value is close to the low voltage V1, and on the other hand, the low-voltage section Vl includes a portion whose value is close to the high voltage V2, so it is not preferable. It is expected that the high-voltage portion is included in the high-voltage section Vh and the low-voltage portion is included in the low-voltage section Vl, regardless of whether it is in the voltage rising section or the voltage falling section. In Figure 6 In the example, sampling is performed once during the high tube voltage interval Vh and once during the low tube voltage interval Vl. However, similar problems arise when sampling is performed twice or more during the high tube voltage interval Vh and twice or more during the low tube voltage interval Vl. Furthermore, starting the low tube voltage interval Vl simultaneously with the end of the high tube voltage interval Vh offers advantages such as minimizing the radiation dose to the imaging subject. However, providing a small time interval between the end of the high tube voltage interval Vh and the start of the low tube voltage interval Vl, and / or between the end of the low tube voltage interval Vl and the start of the high tube voltage interval Vh, can also lead to voltage rising and voltage falling periods.

[0060] In order to include high voltage portions within the high-voltage transistor voltage interval Vh and low voltage portions within the low-voltage transistor voltage interval Vl, both during the voltage-rising and voltage-falling periods, the start time of the high-voltage transistor voltage interval Vh can be set after a predetermined delay time has elapsed from time t0 when the second transistor voltage V2 is applied. In other embodiments, the start time of the high-voltage transistor voltage interval Vh can be set at the time when the transistor voltage exceeds a predetermined threshold. The end time of the high-voltage transistor voltage interval Vh can be set at the time when half of the repetition period (or a predetermined time interval such as 55% of the repetition period) has elapsed from the start time of the high-voltage transistor voltage interval Vh. In other embodiments, the end time of the high-voltage transistor voltage interval Vh can be set at the time when the transistor voltage falls below a predetermined threshold. The sampling timing performed by the DAS 25 can be set based on the results of such automatic analysis of the voltage waveform.

[0061] As described above, as the rotation speed increases and / or the number of views per rotation increases, the stable range of the tube voltage waveform shortens. When the current applied to the X-ray tube 21 is low, the voltage decreases more gradually, while when the current is high, the voltage decreases more steeply. Therefore, regardless of how the start and end times of the high tube voltage interval Vh are set, the energy of the first X-ray based on the tube voltage in the low tube voltage interval Vl and the energy of the second X-ray based on the tube voltage in the high tube voltage interval Vh both fluctuate over time. Consequently, the average energy of the first X-ray, the average energy value, and the average energy value of the second X-ray also fluctuate depending on the rotation speed, the number of views per rotation, and the tube current. It should be noted that the average value referred to in this specification is not simply an average of energy values; it may also refer to a value calculated by taking into account corrections for X-ray absorption characteristics or a value based on a representative value.

[0062] For example, when the first tube voltage V1 is 100 kV and the second tube voltage V2 is 150 kV, Figure 6 At the moment t0 of irradiation Figure 3 The energy distribution of X-rays shown at a tube voltage of 100 kV changes rapidly toward time t1. Figure 3 The energy distribution shown in the figure is shown in the figure with a tube voltage of 150 kV. From time t1 to time t2, the Figure 3 The energy distribution of X-ray irradiation is shown as follows: from time t2 to time t3, the energy distribution of X-ray irradiation is shown as follows: Figure 3 The energy distribution shown at the tube voltage of 150kV is Figure 3 The energy distribution shown in FIG. 1 is a tube voltage of 100 kV. The distribution changes rapidly at first and then gradually. Then, from time t3 to time t4, the distribution maintains Figure 3 The energy distribution of X-ray irradiation at a tube voltage of 100 kV is shown.

[0063] Then, when the rotation speed increases, or the number of views per rotation increases, the time intervals between times t1 and t2 and between times t3 and t4 decrease. When the rotation speed decreases, or the number of views per rotation decreases, the time intervals between times t1 and t2 and between times t3 and t4 increase. Furthermore, regardless of the magnitude of the tube current, the time intervals between times t0 and t1 and between times t1 and t2 do not significantly change. That is, regardless of the magnitude of the tube current, the tube voltage waveform follows a relatively steep rising curve between times t0 and t1. When the tube current is large, the time interval between times t2 and t3 decreases, while the time interval between times t3 and t4 increases. On the other hand, when the tube current is small, the time interval between times t2 and t3 increases, while the time interval between times t3 and t4 decreases. In other words, when the tube current is small, the tube voltage waveform follows a gently falling curve between times t2 and t3.

[0064] The inventors of the present invention have discovered that the above-mentioned changes in the tube voltage waveform have regularity and can be modeled. Specifically, the tube voltage waveform is determined as follows.

[0065] Time t1: a time point determined by a polynomial that takes the potential difference between the second tube voltage V2 and the first tube voltage V1 as a function.

[0066] Time t2, t4: time points determined by the rotation speed and / or the number of views per rotation (for example, when the rotation speed is 1.0 sec / rotation and 1000 views are collected per rotation, t2 is 1 millisecond after t0 and t4 is 2 milliseconds after t0; when the rotation speed is 1.0 sec / rotation and 500 views are collected per rotation, t2 is 2 milliseconds after t0 and t4 is 4 milliseconds after t0).

[0067] Time t3: A time point determined by a polynomial that takes the potential difference between the second tube voltage V2 and the first tube voltage V1 and the tube current as functions.

[0068] Between times t0 and t1: A curve defined by a polynomial that uses the potential difference between the second tube voltage V2 and the first tube voltage V1 as a function. When the potential difference between the second tube voltage V2 and the first tube voltage V1 is large, the time interval between t0 and t1 becomes longer. When the potential difference between the second tube voltage V2 and the first tube voltage V1 is small, the time interval between t0 and t1 becomes shorter.

[0069] Between time t1 and time t2: a straight line determined by the second tube voltage V2.

[0070] Between times t2 and t3: A curve defined by a polynomial that uses the potential difference between the second tube voltage V2 and the first tube voltage V1 and the tube current as functions. If the potential difference between the second tube voltage V2 and the first tube voltage V1 is large, the time interval between t2 and t3 becomes longer. If the potential difference between the second tube voltage V2 and the first tube voltage V1 is small, the time interval between t2 and t3 becomes shorter. If the tube current is large, the time interval between t2 and t3 becomes shorter. If the tube current is small, the time interval between t2 and t3 becomes longer.

[0071] Between time t3 and time t4: a straight line determined according to the first tube voltage V1.

[0072] In this embodiment, the curve portion of the tube voltage waveform is modeled using a polynomial. However, in other embodiments, it is also possible to model using a mathematical model approximated by a combination of a linear function, a quadratic function, and a cubic function.

[0073] The tube voltage waveform can also be modeled as an artificial intelligence (AI) model constructed through deep learning. During deep learning, combinations of the first tube voltage V1, the second tube voltage V2, the tube current, the rotational speed, and the number of views / rev, along with the actual tube voltage waveforms measured for these combinations, are used as training data. This allows the construction of a fully learned AI model.

[0074] On the other hand, as described in Patent Document 3, beam hardening correction has been performed on image data collected through 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 from a penetrating object shifts toward higher energy values ​​in the case of polychromatic X-rays, due to the spectral correlation of the actual object's light attenuation properties. In the reconstructed image of the object, linear, spectrally correlated light attenuation can be observed via a shift in grayscale values ​​relative to theoretical values. In particular, shifts in grayscale values ​​in the reconstructed image caused by high-charge and high-density materials (such as bone) or beam-hardened virtual images can hinder accurate image interpretation.

[0075] To mitigate or eliminate this beam hardening effect, beam hardening correction is applied to image data collected through multi-energy imaging. A beam hardening correction coefficient is used. The beam hardening correction coefficient is obtained by performing a calibration scan using multi-energy imaging, such as a water phantom. During calibration, the coefficient is set so that the CT value of water is 0 and the CT value of air is -1000. However, as mentioned above, the average energy of X-rays varies depending on the rotational 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. Therefore, to accurately set the beam hardening correction coefficient, calibration has traditionally been performed using a combination of the rotational 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.

[0076] However, by modeling the tube voltage waveform as described above, if 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 is determined, the corresponding tube voltage waveform can be determined. If the tube voltage waveform can be determined, the average energy of the X-ray beam based on it can be determined, and the beam hardening correction coefficient to be applied can be obtained.

[0077] Figure 7 This diagram illustrates 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, conventional methods such as calibration imaging using a phantom can be used to obtain the BH vector for the X-ray beam at energies corresponding to major tube voltages, such as 80 kVp, 100 kVp, 120 kVp, and 140 kVp.

[0078] Then, when the average energy of the X-ray beam obtained by modeling the tube voltage waveform is, for example, 90 kVp, as shown in FIG. Figure 7 As shown in FIG, by using the average of the BH vector of 80 kVp and the BH vector of 100 kVp, a BH vector of 90 kVp can be obtained. In another example, when the BH vector obtained is a value (X kVp) between 120 kVp and 140 kVp,

[0079] Using the weighted average of the BH vectors of 120kVp and 140kVp, we can find XkVp

[0080] In this case, the value of the BH vector of XkVp in each channel is: the value of the BH vector of XkVp = ((140-X) × the value of the BH vector of 120kVp + (X-120) × the BH vector of 140kVp)

[0081] vector value) / 20(120 <X<140)。

[0082] Furthermore, it is known to those skilled in the art that the scattered X-rays generated by Compton scattering and the reflected

[0083] The energy of the electrons changes according to the energy of the incident X-rays, and the distribution of the scattered X-rays also changes according to the energy of the incident X-rays.

[0084] In order to remove or reduce the influence of such scattered X-rays, conventional

[0085] Perform scatter correction. This scatter correction is particularly important in dental CT systems and baggage inspection systems, where the scanned object is likely to contain metal. The coefficients used for this scatter correction can also be calculated based on the average value of the X-ray beam calculated using the tube voltage waveform model.

[0086] In addition, the average value of the X-ray beam can also be used to calculate the material decomposition

[0087] decomposition: The absorption value μ of the matrix material pair used in the calculation of matrix material decomposition.

[0088] Furthermore, the rate of X-ray absorption varies depending on the type and concentration of elements constituting the imaging object. When substances with higher element numbers are present at higher densities, the X-ray dose received by the detector is reduced. Furthermore, those skilled in the art know that the degree of absorption varies depending on the energy of the incident X-rays. To eliminate or mitigate the effects of this X-ray absorption, absorption correction has traditionally been performed.

[0089] This absorption correction is used in dental CT equipment or other equipment where there is a high possibility that the scanned object contains metal.

[0090] This is particularly important in CT scanners for medical examinations. 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.

[0091] Any of the beam hardening correction coefficient, the scattering correction coefficient, and the absorption correction coefficient can be generated for each of the plurality of detector elements 24a, or can be generated for each of the plurality of detector elements 24a.

[0092] The image generator 35 can also reconstruct an image using the difference between the beam hardening correction coefficient calculated by the preprocessor 34 based on the tube voltage waveform model and the beam hardening correction coefficient calculated by imaging the phantom.

[0093] By setting the parameters for radiographic image creation based on the low-energy and high-energy averages calculated from the tube voltage waveform model, calibration imaging and calculation time can be shortened, which can reduce CT system downtime and service costs.

[0094] In a preferred embodiment of the present invention, a function for updating the tube voltage waveform model is provided. Due to repeated use of the CT system or aging, the waveform of the voltage output by the power supply unit, or degradation of the target of the X-ray tube 21, the shape of the tube voltage waveform may change. Alternatively, even if the tube voltage waveform remains unchanged, the energy distribution of the resulting X-ray beam may change. In such cases, the tube voltage waveform model generated at factory shipment or at the start of use may not be suitable for the current CT system. In such cases, the tube voltage waveform model is corrected based on projection data measured during calibration.

[0095] It should be noted that the present invention is not limited to the present embodiment, and various modifications can be made without departing from the gist of the invention.

[0096] Furthermore, a program for causing a computer to function as each means for performing control or processing in the above-mentioned X-ray CT apparatus is also an example of an embodiment of the invention.

[0097] Description of Reference Numerals

[0098] 1: Operation console 2: Input device 3: Central processing unit 5: Data acquisition buffer 6: Monitor 7: Storage device 10: Radiographic table 12: Bracket 15: Gantry rotation unit 20: Scanner 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: Preprocessing unit 35: Image generator 71: Subject / imaging object 81: X-ray beam 100: X-ray CT apparatus

Claims

1. A system comprising a processor that processes a detection signal of radiation transmitted through an imaging object to generate a radiographic image, wherein: The radiation is emitted toward the imaging subject by a radiation tube while rotating around the imaging subject. 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 low tube voltage and the high tube voltage are applied to the radiation tube by alternately switching 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 highly stable section between the rising portion and the falling portion, and a less stable section between the falling portion and the rising portion. The processor Accepts input of rotation speed and / or number of views per rotation, determining a tube voltage waveform corresponding to the rotation speed and / or the number of views per rotation using a waveform determination model, calculating a low energy average value of the radiation corresponding to a low stable interval of the tube voltage waveform and a low voltage interval including a portion of a falling portion of the tube voltage waveform, and a high energy average value of the radiation corresponding to a high stable interval of the tube voltage waveform and a high voltage interval including another portion of the falling portion of the tube voltage waveform, Parameters for producing the radiographic image are set based on the low energy average value and the high energy average value.

2. The system according to claim 1, comprising: a workbench on which the imaging object is placed; a gantry rotatably supporting a detector for detecting the radiation transmitted through the imaging object and the radiation tube; a storage medium storing the waveform determination model; a user interface for inputting the rotation speed and / or the number of views per rotation; and An image reconstruction device includes the processor.

3. The system according to claim 1, wherein: The parameters include a beam hardening correction factor.

4. The system according to claim 1, wherein: The parameters include X-ray absorption coefficient and / or X-ray scattering coefficient.

5. The system according to claim 1, wherein: The low voltage interval includes a portion of the rising portion of the tube voltage waveform. The high voltage interval includes another part of the rising portion of the tube voltage waveform, The waveform determination model makes the voltage drop portion relatively gentle when the current value of the power applied to the radiation tube is low and makes the voltage drop portion relatively steep when the current value is high, while the voltage rise portion does not change significantly whether the current value is low or high.

6. The system according to claim 1, wherein: The tube voltage waveform is collected by measuring the power applied to the radiation tube. In the waveform determination model, the curve of the voltage drop portion of the measured waveform is expressed by a polynomial approximation.

7. The system according to claim 6, wherein: Preprocessor Separating the tube voltage waveform into the low voltage interval and the high voltage interval according to a predetermined threshold value, The sampling for generating the detection signal is performed for a first time in each of the low voltage intervals and for a second time in each of the high voltage intervals. The first number of times is the same as the second number of times.

8. The system according to claim 7, wherein: The detector includes a plurality of detector elements extending along the circumferential direction of the gantry, The radiation tube irradiates a fan beam or a cone beam spreading in the circumferential direction toward the detector as the radiation, The detector detects the radiation that has passed through the phantom to generate a phantom detection signal. generating a first phantom-based beam hardening correction coefficient corresponding to the low-energy radiation based on the phantom detection signal; generating the first phantom-based beam hardening correction coefficient for each of the plurality of detector elements, The parameters include a first model-based beam hardening correction coefficient corresponding to the low voltage interval, The radiographic image is image reconstructed by using a first difference between the first model-based beam hardening correction coefficient and the first phantom-based beam hardening correction coefficient.

9. The system according to claim 8, wherein: generating a second phantom-based beam hardening correction coefficient corresponding to the high-energy radiation based on the phantom detection signal; generating a second phantom-based beam hardening correction coefficient for each of the plurality of detector elements, The parameters include second model-based beam hardening correction coefficients corresponding to the high voltage and low voltage intervals, The radiographic image is image reconstructed by 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 transmitted through an imaging object to generate a radiographic image, wherein: The radiation is emitted toward the imaging subject by a radiation tube while rotating around the imaging subject. 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 low tube voltage and the high tube voltage are applied to the radiation tube by alternately switching 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 highly stable section between the rising portion and the falling portion, and a less stable section between the falling portion and the rising portion. The program causes the processor to execute: Accepts input of rotation speed and / or number of views per rotation, determining a tube voltage waveform corresponding to the rotation speed and / or the number of views per rotation using a waveform determination model, calculating a low energy average value of the radiation corresponding to a low stable interval of the tube voltage waveform and a low voltage interval including a portion of a falling portion of the tube voltage waveform, and a high energy average value of the radiation corresponding to a high stable interval of the tube voltage waveform and a high voltage interval including another portion of the falling portion of the tube voltage waveform, Parameters for producing the radiographic image are set based on the low energy average value and the high energy average value.

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

12. The program according to claim 10, wherein The parameters include a beam hardening correction factor.

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

14. The program according to claim 10, wherein The low voltage interval includes a portion of the rising portion of the tube voltage waveform. The high voltage interval includes another part of the rising portion of the tube voltage waveform, The waveform determination model makes the voltage drop portion relatively gentle when the current value of the power applied to the radiation tube is low and makes the voltage drop portion relatively steep when the current value is high, while the voltage rise portion does not change significantly whether the current value is low or high.

15. The program according to claim 11, wherein The tube voltage waveform is collected by measuring the power applied to the radiation tube. In the waveform determination model, the curve of the voltage drop portion of the measured waveform is expressed by a polynomial approximation.

16. The program according to claim 15, wherein The program causes the processor to execute: separating the tube voltage waveform into the low voltage interval and the high voltage interval using a predetermined threshold value; The sampling for generating the detection signal is performed for a first time in each of the low voltage intervals and for a second time in each of the high voltage intervals. The first number of times is the same as the second number of times.

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

18. The program according to claim 17, wherein generating a second phantom-based beam hardening correction coefficient corresponding to the high-energy radiation based on the phantom detection signal; generating a second phantom-based beam hardening correction coefficient for each of the plurality of detector elements, The parameters include second model-based beam hardening correction coefficients corresponding to the high voltage and low voltage intervals, The radiographic image is image reconstructed by 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 claim 10 .

Images