JP2007125129A - X-ray computed tomography apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To eanable photography based on optional tube voltage in an X-ray computed tomography (CT) apparatus. <P>SOLUTION: The x-ray CT apparatus has an X-ray tube, an X-ray detector arranged opposite to the X-ray tube and a rotation mechanism rotating a pair of the X-ray tube and the X-ray detector, and generates a tomogram of a subject based on the radiolucent area of the subject obtained at a large number of positions in the circumferential direction of rotation, wherein air data and a linearity correction coefficient with respect to optional consecutive tube voltage are calculated based on air data and a linearity correction coefficient measured with a small number of discrete tube voltage. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

  The present invention relates to an X-ray CT apparatus, and more particularly to an X-ray CT apparatus for generating a tomographic image of a human body and performing image diagnosis.

An X-ray CT apparatus is an apparatus that captures X-ray transmission images of a subject from various angles and reconstructs the obtained captured images to generate a tomographic image of the subject. In general, some image data is corrected prior to reconstruction calculation. Typical examples of such correction include air calibration and linearity correction.
The air calibration is an operation for normalizing the photographed image I of the subject with the air image Io photographed without placing the subject, and is expressed by the following equation.

Where l is the position on the X-ray beam and μ (l) is the absorption coefficient of the subject at position l. The reconstruction calculation is to calculate the absorption coefficient μ (l) of the subject from the data after air calibration based on the inverse Radon transform. When the absorption coefficient of the subject is uniform, Equation 1 is as follows.

Here, d is the path length in the subject of the X-ray beam. Air data is usually taken when the X-ray CT apparatus is activated. Usually, air data for several hundred to several thousand frames is photographed to improve S / N, and the average value is used for calibration. Since air data differs for each tube voltage of the X-ray tube, it is usually acquired for each tube voltage used for imaging.

  The linearity correction is a correction generally performed in an X-ray CT apparatus especially for the purpose of image diagnosis, and is performed on data after air calibration in order to improve the accuracy of the CT value. According to Equation 2, when a subject having a constant absorption coefficient such as water is photographed, the value of μd should change linearly with respect to the subject thickness d. However, in practice, the linearity is lowered due to the beam hardening effect and the distortion of the input / output characteristics of the X-ray detector, so that the CT value of the reconstructed image becomes non-uniform. Linearity correction is performed to avoid such a problem. There is Patent Document 1 as a conventional example of linearity correction. However, in Patent Document 1, linearity correction is called phantom calibration. In Patent Document 1, the nonlinear characteristic of μd is measured using a uniform water phantom whose subject thickness d is known, and this is corrected by polynomial transformation. Since the linearity correction of Patent Document 1 is optimized for a water phantom, effective correction accuracy is not always maintained depending on the composition of the subject. However, this is a very effective correction when a subject having a composition close to water, such as a human body, is targeted. Since the beam hardening effect changes depending on the tube voltage of the X-ray tube, the linearity correction polynomial is usually measured for each tube voltage used for imaging. The linearity correction polynomial is usually measured during maintenance of the CT apparatus.

JP-B 61-54412

  In order to acquire air calibration data and linearity correction data, it is usually necessary to perform various types of imaging, which requires a lot of time and labor for measurement. For this reason, at present, the number of tube voltages that can be photographed is normally limited to about 3 to 4 types. On the other hand, since the tube voltage is an important parameter that affects the image quality and exposure dose, there is a need to optimize the imaging conditions by increasing the types of tube voltage that can be set as much as possible.

  An object of the present invention is to provide a method for generating air data and linearity correction data of an arbitrary tube voltage based on air data and linearity correction data measured with a small number of types of tube voltages.

The following is a brief description of an outline of typical inventions among the inventions disclosed in the present application.
(Means 1)
An X-ray tube; an X-ray detector disposed opposite to the X-ray tube; and a rotation mechanism for rotating a pair of the X-ray tube and the X-ray detector, and a plurality of positions in the circumferential direction of the rotation. In the X-ray CT apparatus for generating a tomographic image of the subject based on the X-ray transmission image of the subject acquired in step N, the N types of the X-ray tube (where N is a natural number of 2 or more) without the subject being arranged Recording means for recording measured values of air data photographed at a tube voltage of), means for calculating theoretical values of air data for different M types of tube voltages (where M is a natural number satisfying N <M), and , Means for calculating a difference value between the measured value of the air data and the theoretical value of the air data for the N types of tube voltages, and means for approximating the difference value of the air data with an approximate expression using the tube voltage as a variable. And using the approximate expression for the M types of tube voltages An X-ray comprising means for calculating a corrected air data by calculating a difference value of the data and adding it to a theoretical value of the air data, and means for calculating an air calibration using the corrected air data CT device. As a result, air data corresponding to a larger number (M types) of tube voltages can be generated from air data measured with a small number (N types) of tube voltages. An X-ray CT apparatus that can cope with tube voltage can be provided. Further, since only the difference between the theoretical value and the actual measurement value of the air data is approximated by an approximate expression using the tube voltage as a variable, the above approximation can be performed with high accuracy without being influenced by the theoretical value component. As a result, it is possible to reduce the decrease in air calibration accuracy caused by the approximation error.
(Means 2)
The X-ray CT apparatus according to means 1, further comprising second recording means for recording the corrected air data. As a result, air data can be calculated and recorded in advance for a large number of tube voltages, so that air calibration can be performed at high speed.
(Means 3)
Third recording means for actually measuring and recording photographing data at the N types of tube voltages with respect to a simulated subject composed of a substantially uniform material, and photographing data of the simulated subject with respect to the M types of tube voltages Means for calculating a theoretical value of the imaging data, means for calculating a difference value between the measured value of the photographing data and the theoretical value of the photographing data with respect to the N types of tube voltages, and the difference value of the photographing data as a variable of the tube voltage. And means for approximating with the approximate expression, and calculating the difference value of the photographic data for the M types of tube voltages using the approximate expression, and then adding to the theoretical value of the photographic data to calculate the corrected photographic data. Means for calculating a conversion formula for converting the corrected shooting data so that a CT value of the simulated subject generated with respect to the corrected shooting data is substantially constant, and shooting of the subject based on the conversion formula Data Is an X-ray CT apparatus according to the means 1 and 2 characterized by having a positive to means. As a result, linearity correction data corresponding to a larger number (M types) of tube voltages can be generated from shooting data measured with a small number (N types) of tube voltages, thereby reducing the time and effort required to shoot a simulated subject. In addition, it is possible to provide an X-ray CT apparatus that can handle a large number of tube voltages. In addition, since only the difference between the theoretical value and the actual measurement value of the photographed data is approximated by an approximate function using the tube voltage as a variable, the above approximation can be performed with high accuracy without being influenced by the theoretical value component. As a result, the decrease in linearity correction accuracy caused by the approximation error can be reduced.
(Means 4)
The X-ray CT apparatus according to means 3, further comprising fourth recording means for recording a parameter value defining the conversion formula. As a result, linearity correction data can be calculated and recorded in advance for a large number of tube voltages, so that the linearity correction calculation can be performed at high speed.

The following is a brief description of the effects obtained by the typical inventions among the inventions disclosed in the present application.
Air data and linearity correction data that enable X-ray CT measurement at an arbitrary tube voltage can be created with little effort and time. As a result, the tube voltage can be set with high accuracy, and the quality of the measurement image can be improved.

Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings.
FIG. 1 is a schematic front view of an X-ray CT apparatus according to an embodiment of the present invention. The X-ray CT apparatus according to the present embodiment includes an X-ray tube 1, an X-ray detector 2, a rotating plate 3, a couch top 4, a gantry 5, an imaging control means 100, a console 101, a frame memory 102, an air calibration. 103, linearity correction means 104, image reconstruction means 105, image display means 106, calculation means A 107, discrete air data table 108, calculation means B 109, continuous air data table 110, calculation means C 111, discrete linearity correction table 112, The calculation means D113, continuous linearity correction table 114, discrete μd table 115, continuous μd table 116, and the like are included. An imaging system consisting of a pair of an X-ray tube 1 and an X-ray detector 2 is fixed to a rotating plate 3, and the entire imaging system and the rotating plate 3 are stored inside a gantry 5. An opening 6 is provided at the center of the gantry 5, and a subject 10 is disposed near the center of the opening 6. In the present embodiment, a human body is assumed as the subject 10, and the normal subject 10 is measured while lying on the couch top 4. The rotating plate 3 is rotated by a drive motor (not shown) to capture an X-ray transmission image of the subject 10 from the entire circumference.

  In FIG. 1, a representative example of the distance between the X-ray generation point of the X-ray tube 1 and the X-ray input surface of the X-ray detector 2 is 1040 [mm]. A typical example of the diameter of the opening 8 is 500 [mm]. A typical example of the time required for the rotation of the rotating plate 3 is 0.4 [s]. As the X-ray detector 2, a known X-ray detector including a scintillator and a photodiode is used. The X-ray detector 2 is a multi-slice CT multi-layer sensor. A typical example of the number of channels in the horizontal direction on the paper is 900 channels, and a typical example of the number of slices in the vertical direction on the paper is 64 slices. A typical example of the size in the channel direction and slice direction of each detection element is 1 [mm]. A typical example of the number of times of photographing in one rotation of the photographing system is 900 times, and one photographing is performed every time the rotating plate 3 rotates 0.4 degrees.

  Next, the operation of the X-ray CT apparatus according to this embodiment will be described. The X-ray CT apparatus is provided with two types of imaging modes, a main imaging mode and a calibration imaging mode. The examiner gives an instruction through the console 101 to select the main photographing mode and the calibration photographing mode. In FIG. 1, the data flow in the main photographing mode is indicated by a solid line, and the data flow in the calibration photographing mode is indicated by a broken line. The operation in both shooting modes will be described below.

  First, the operation of the X-ray imaging apparatus in the main imaging mode will be described. The user instructs the start of shooting after designating shooting conditions such as the tube voltage and tube current at the time of shooting and the shooting range of the subject 10 through the console 101. At the same time that the start of photographing is instructed, the photographing control means 100 starts to rotate the rotating plate 3. When the rotation of the rotating plate 3 enters a constant speed state, the imaging control means 100 instructs the X-ray tube 1 to start X-ray irradiation and the X-ray detector 2 to start imaging, and starts imaging. Imaging data output from the X-ray detector 2 is sequentially stored in the frame memory 102. The air calibration unit 103 performs air calibration shown in Equation 1 on all channel and slice position data stored in the frame memory 102. At this time, air data corresponding to the imaging tube voltage is referred to from the continuous air data table 110. Subsequently, the linearity correction unit 104 reads the linearity correction data corresponding to the photographing tube voltage with reference to the continuous linearity correction table 114, and performs the linearity correction 104 by a method described later. Subsequently, the image reconstruction unit 105 reconstructs a tomographic image of the subject 10 using a known reconstruction algorithm, and the reconstruction result is displayed by a known image display unit 106 such as a monitor.

Next, the operation of the X-ray imaging apparatus in the calibration imaging mode will be described. In the calibration shooting mode, air data and linearity correction data are measured.
At the time of air data measurement, photographing is performed without the subject 10 and the couch top 4 being arranged. When the user instructs the start of air data shooting through the console 101, the rotating plate 3 starts rotating, and when the rotation enters a constant speed state, air data shooting is started. In photographing, air data for about one or two rotations of the rotating plate 3 is acquired and recorded in the frame memory 102. Next, the computing means A calculates the average value of the air data for all frames and stores it in the discrete air data table 108. The series of operations from the start of imaging of the air data to the storage of the average air data in the discrete air data table 108 is performed on the discrete tube voltage values designated in advance, and the above operation is performed on all tube voltages. When the series of work is finished, the air data shooting is finished. In the following description, it is assumed that five types of tube voltages of 90 kV, 100 kV, 110 kV, 120 kV, and 130 kV are designated as the discrete tube voltage values. However, the present invention is limited to this example. It is not a thing. When the imaging of the air data is completed, the calculation unit B109 creates air data for a continuous tube voltage using a method described later, and stores the result in the continuous air data table 110. In the present embodiment, a continuous tube voltage is defined as a tube voltage in 1 kV increments in a section of 90 kV to 130 kV. If the continuous air data table 110 is used, it is possible to perform imaging at an arbitrary tube voltage in increments of 1 kV. The data structures of the discrete air data table 108 and the continuous air data table 110 will be described later.

  At the time of linearity correction data measurement, cylindrical water phantoms having different diameters as shown in Document 1 are arranged as subjects, and the respective shooting data are stored in the frame memory 102. The calculation means C111 creates a linearity correction table, which will be described later, using a known method disclosed in Document 1, and stores the result in the discrete linearity correction table 112. The series of operations from the photographing of the water phantom to the storage of the linearity correction table in the discrete linearity correction table 112 is performed for all the discrete tube voltage values. When the linearity correction table has been stored in the discrete linearity correction table 112 for all the discrete tube voltage values, the computing means C111 creates the discrete μd table 115 using a method described later. Subsequently, the calculation means D113 creates a μd table for the continuous tube voltage by referring to the discrete μd table 115 by a method described later, and stores the result in the continuous μd table 116. Further, the calculation means D113 creates a linearity correction table for the continuous tube voltage by a method described later with reference to the continuous μd table 116, and stores the result in the continuous linearity correction table 114. The data structures of the discrete linearity correction table 112, the continuous linearity correction table 114, the discrete μd table 115, and the continuous μd table 116 will be described later.

In FIG. 1, each of the air calibration means 103, linearity correction means 104, image reconstruction means 105, calculation means A107, calculation means B109, calculation means C111, and calculation means D113 is a dedicated calculator or a known general-purpose calculation. A vessel or the like is used.
FIG. 2 is a diagram for explaining the data structure of the discrete air data table 108. In FIG. 2, for simplicity, a table for one pixel (channel position i, slice position j) of the X-ray detector 2 is shown, but in reality, air data is stored for all pixels. The average value of the air data measured with the discrete tube voltage is stored at a corresponding position in the discrete air data table 108.

FIG. 3 is a diagram for explaining the data structure of the continuous air data table 110. As in the case of FIG. 2, only the table for one pixel of the X-ray detector 2 is shown for simplicity. The average value of the air data calculated for the continuous tube voltage by the calculation means B109 is stored in the corresponding position of the continuous air data table 110. However, the data in FIG. 2 is copied for air data that matches the discrete tube voltage in FIG.
FIG. 4 is a diagram for explaining a method of calculating air data for a continuous tube voltage from air data measured at a discrete tube voltage in the arithmetic means B109. The plot in FIG. 4 (a) shows the air data signal value Io measured for a discrete tube voltage. On the other hand, the solid line in FIG. 4 (a) is the result of calculating the theoretical value of air data using the method described later. This theoretical value is calculated for continuous tube voltage. The plot of FIG. 4 (b) shows the difference value between the actual measurement value and the theoretical value shown in FIG. 4 (a). Such an error between the actually measured value and the theoretical value is caused by distortion of the input / output characteristics of the X-ray detector 2 or the like. The broken line in FIG. 4B is a polynomial approximation of the plot in FIG. 4B. Usually, the following cubic equation is used as the approximate polynomial.

However, a0 to a3 are approximate coefficients, and V is the tube voltage. Now, assuming that the theoretical value of air data calculated in FIG. 4 (a) is J (V), the value of air data for an arbitrary tube voltage V can be calculated by the following equation.

The calculation means B109 calculates the theoretical value of the air data for the continuous tube voltage, performs the approximation of Equation 3 and the calculation of Equation 4, and stores the result in the corresponding position of the continuous air data table 110.

  FIG. 5 is a diagram for explaining a correction method in the linearity correction means 104. In the measurement of Document 1, photographing is performed on cylindrical water whose subject thickness d is known in advance, and air calibration is performed. At this time, from Equation 2, the signal value after air calibration corresponds to μd. In the linearity correction, the actually measured value of μd is converted into an ideal value. Here, according to the definition of the CT value, normalization is performed so that the water absorption coefficient is 1000, so the ideal value of μd is 1000 d. FIG. 5 shows the result of plotting such measured values and ideal values of μd as values on the X and Y axes. The linearity correction is a conversion of a value from X to Y, and such conversion is usually realized by approximating the plot with the following polynomial.

On the other hand, if the conversion from Y to X is approximated by a polynomial such as the following equation, an actually measured value of μd can be calculated for an arbitrary subject thickness d.

  FIG. 6 is a diagram for explaining the data structure of the discrete linearity correction table 112. In FIG. 6, for simplicity, a table for one pixel (channel position i, slice position j) of the X-ray detector 2 is shown, but in reality linearity correction coefficients are stored for all pixels. The calculation means C111 approximates the relationship between the actually measured value of the obtained μd and its ideal value for each discrete tube voltage by Equation 5, and stores the obtained correction coefficients b2 and b1 in the corresponding table positions. To do.

  FIG. 7 is a diagram for explaining the data structure of the discrete μd table 115. As in the case of FIG. 6, only a table for one pixel of the X-ray detector 2 is shown for simplicity. If Equation 6 is used, an actual measurement value of μd can be calculated for an arbitrary subject thickness d. The computing means C111 calculates the actual measured value of μd at each subject thickness shown in the discrete μd table 115 for each discrete tube voltage, and stores the result in the corresponding table position.

  FIG. 8 is a diagram for explaining the data structure of the continuous linearity correction table 114. As in the case of FIG. 6, only a table for one pixel of the X-ray detector 2 is shown for simplicity. The linearity correction coefficient calculated for the continuous tube voltage by the calculation means D113 is stored in the corresponding position of the continuous linearity correction table 114. However, the data in FIG. 6 is copied for the linearity correction coefficient that matches the discrete tube voltage in FIG.

  FIG. 9 is a diagram for explaining the data structure of the continuous μd table 116. As in the case of FIG. 7, only the table for one pixel of the X-ray detector 2 is shown for simplicity. The value of μd calculated for the continuous tube voltage by the calculation means D113 is stored in the corresponding position of the continuous μd table 116. However, the data in FIG. 7 is copied for the value of μd that matches the discrete tube voltage in FIG.

FIG. 10 is a diagram for explaining a method of calculating the μd value for the continuous tube voltage from the μd value measured at the discrete tube voltage in the calculation means D113. The plot in FIG. 10A shows the relationship between the discrete tube voltage and the actually measured μd value for one subject thickness in the discrete μd table 115. On the other hand, the solid line in FIG. 10A shows the result of calculating the theoretical value of the μd value using the method described later. This theoretical value is calculated for continuous tube voltage.
The plot of FIG. 10 (b) shows the difference value between the actual measurement value and the theoretical value shown in FIG. 10 (a). Such an error between the actually measured value and the theoretical value is caused by the scattered X-ray, the distortion of the input / output characteristics of the X-ray detector 2, or the like. A broken line in FIG. 10B is a polynomial approximation of the plot in FIG. Usually, the following cubic equation is used as the approximate polynomial.

However, p0 to p3 are approximate coefficients. Assuming that the theoretical value of the μd value calculated in FIG. 10A is K (V), the μd value for an arbitrary tube voltage V can be calculated by the following equation.

The calculation means D113 calculates the theoretical value of μd for the continuous tube voltage, performs the approximation of Equation 7 and the calculation of Equation 8, and stores the result in the corresponding position of the continuous μd table 116. Subsequently, the computing means D113 refers to the continuous μd table 116, and approximates the relationship between the μd value (X axis) and the corresponding ideal μd value (Y axis) at each tube voltage by Equation 5. The linearity correction coefficients b2 and b1 obtained at this time are stored in the continuous linearity correction table 114.

  FIG. 11 is a diagram for explaining a calculation method of air data and the theoretical value of μd. The X-ray beam radiated from the X-ray generation point 1100 is partially attenuated by the intrinsic filtration 1103 of the X-ray tube 1, the quality filter 1104, the compensation filter 1105, the subject 1106, etc., and then the scintillator in the X-ray detector 2. The light is incident on 1102 and detected. Where X-ray energy is ε, X-ray energy spectrum at tube voltage V is Nv (ε), X-ray beam path length and total absorption coefficient in intrinsic filtration 1103 are λ1, μ1 (ε), and radiation quality, respectively. The path length and the total X-ray absorption coefficient of the X-ray beam in the filter 1104 are respectively λ2, μ2 (ε), and the path length of the X-ray beam and the total X-ray absorption coefficient in the compensation filter 1105 are respectively λ3 and μ3 (ε). , The path length of the X-ray beam in the subject 1106 and the total X-ray absorption coefficient are λw, d, the thickness of the scintillator 1102 and the energy absorption coefficient are λd, μd (ε), respectively, and the gain of the X-ray detector is α Then, the air data and the signal data after passing through the subject can be calculated by equations 9 and 10, respectively.

The energy spectrum Nv (ε) can be calculated using the method described in Med. Phys. Vol. 15, No. 5, 713-720, Sep / Oct, 1988 (hereinafter referred to as Reference 2). When calculating the theoretical value of μd, Equations 9 and 10 may be substituted into Equation 2. At this time, assuming that the subject 1106 is homogeneous water, μd is calculated for various subject thicknesses d shown in the discrete μd table 115.

  If the X-ray CT apparatus shown in the present embodiment is used as described above, the air data and linearity correction coefficient for any continuous tube voltage based on the air data and linearity correction coefficient measured with a small number of discrete tube voltages. Therefore, it is possible to provide an X-ray CT apparatus that can reduce the measurement time of these correction data and can cope with a large number of tube voltages. In the above calculation, only the difference between the theoretical value and the actual measurement value is approximated by a polynomial having the tube voltage as a variable, so that the approximation can be performed with high accuracy without being influenced by the theoretical value component. As a result, the reduction in correction accuracy due to the approximation error can be reduced.

  As mentioned above, although the example of the X-ray CT apparatus according to the present invention has been described using the embodiment, the present invention is not limited to this example, and various modifications can be made without departing from the scope of the present invention. Not too long. For example, in the present embodiment, air data and linearity correction coefficients for continuous tube voltages are obtained in advance and recorded in the continuous air data table 110 and the continuous linearity correction table 114, respectively. May be calculated each time. In this case, a memory space for storing the continuous air data table 110 and the continuous linearity correction table 114 becomes unnecessary, and the apparatus cost can be reduced.

1 is a schematic front view of an X-ray CT apparatus according to an embodiment of the present invention. It is a figure for demonstrating the data structure of the discrete air data table. It is a figure for demonstrating the data structure of the continuous air data table. It is a figure for demonstrating the method to calculate the air data with respect to continuous tube voltage from the air data measured by discrete tube voltage in calculating means B109. It is a figure for demonstrating the correction method in the linearity correction | amendment means 104. FIG. It is a figure for demonstrating the data structure of the discrete linearity correction table. 6 is a diagram for explaining a data structure of a discrete μd table 115. FIG. It is a figure for demonstrating the data structure of the continuous linearity correction table. It is a figure for demonstrating the data structure of the continuous micro | micron | mud table. It is a figure for demonstrating the method to calculate the (micro | micron | mu) d value with respect to a continuous tube voltage from the (micro | micron | mu) d value measured with the discrete tube voltage in the calculating means D113. It is a figure for demonstrating the calculation method of the air data and the theoretical value of μd.

Explanation of symbols

1 ... X-ray tube,
2 ... X-ray detector,
3 ... Rotating plate,
4 ... Couch top,
5 ... Gantry,
6 ... opening,
10 ... Subject,
1100 ... X-ray generation point,
1102 .. scintillator,
1103 ... Inherent filtration,
1104 ... quality filter,
1105: Compensation filter,
1106: Subject.

Claims (4)

  An X-ray tube; an X-ray detector disposed opposite to the X-ray tube; and a rotation mechanism for rotating a pair of the X-ray tube and the X-ray detector, and a plurality of positions in the circumferential direction of the rotation. In the X-ray CT apparatus for generating a tomographic image of the subject based on the X-ray transmission image of the subject acquired in step N, the N types of the X-ray tube (where N is a natural number of 2 or more) without the subject being arranged Recording means for recording measured values of air data photographed at a tube voltage of), means for calculating theoretical values of air data for different M types of tube voltages (where M is a natural number satisfying N <M), and , Means for calculating a difference value between the measured value of the air data and the theoretical value of the air data for the N types of tube voltages, and means for approximating the difference value of the air data with an approximate expression using the tube voltage as a variable. And using the approximate expression for the M types of tube voltages An X-ray comprising means for calculating a corrected air data by calculating a difference value of the data and adding it to a theoretical value of the air data, and means for calculating an air calibration using the corrected air data CT device.   The X-ray CT apparatus according to claim 1, further comprising a second recording unit that records the corrected air data.   Third recording means for actually measuring and recording photographing data at the N types of tube voltages with respect to a simulated subject composed of a substantially uniform material, and photographing data of the simulated subject with respect to the M types of tube voltages Means for calculating a theoretical value of the imaging data, means for calculating a difference value between the measured value of the photographing data and the theoretical value of the photographing data with respect to the N types of tube voltages, and a difference between the photographing data as a variable of the tube voltage And means for approximating with the approximate expression, and calculating the difference value of the photographic data for the M types of tube voltages using the approximate expression, and then adding to the theoretical value of the photographic data to calculate the corrected photographic data. Means for calculating a conversion formula for converting the corrected shooting data so that a CT value of the simulated subject generated with respect to the corrected shooting data is substantially constant, and shooting of the subject based on the conversion formula Data X-ray CT apparatus according to claim 1 or 2, characterized in that it has a positive to means. The X-ray CT apparatus according to claim 3, further comprising a fourth recording unit that records a parameter value defining the conversion formula.

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