JP2009142578A - X-ray ct apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide an X-ray CT apparatus capable of easily executing a process using an image reconstruction variable corresponding to multiple different imaging conditions such as a change in X-ray tube voltage conditions in an image reconstruction process of the X-ray CT apparatus. <P>SOLUTION: This X-ray CT apparatus 100 is provided with: an imaging condition setting means capable of setting a plurality of conditions on predetermined imaging conditions; an X-ray data collection means scanning a subject and collecting X-ray projection data while rotating a rotary section having an X-ray generator and an X-ray detector using the imaging conditions; a storage section storing image reconstruction variables different according to the plurality of conditions and at least corresponding to a typical condition, incidentally to a data set collected by the X-ray projection data collection means; and an image reconstruction section including an image reconstruction process of the X-ray projection data based on the image reconstruction variable corresponding to the imaging condition used for the collection of the X-ray data including at least the conditions other than the typical condition. <P>COPYRIGHT: (C)2009,JPO&INPIT

Description

  The present invention relates to an X-ray CT (Computed Tomography) apparatus for taking a tomographic image of a subject such as a patient.

  In the X-ray CT apparatus, for each X-ray tube voltage that can be set as an imaging condition, there are variables (hereinafter referred to as image reconstruction variables) used for image reconstruction such as correction data used in preprocessing of image reconstruction processing. It is generally memorized. In addition, the format of the image reconstruction variable is configured in a predetermined format.

The X-ray projection data includes all or a part of the image reconstruction variable data as supplementary information of the X-ray projection data, creates a data file, and media such as a DVD or other network or network as independent X-ray projection data And transferred to another X-ray CT apparatus.
Incidentally, the X-ray CT apparatus has imaged a CT value tomogram showing the distribution of the X-ray absorption coefficient, but the demand for the function of the X-ray CT apparatus has been diversified. Recently, a technique for imaging the distribution of each element by using the X-ray tube voltage dependence of the X-ray absorption coefficient of each element, a technique called dual energy scanning has been desired. (For example, refer to Patent Document 1).
JP 20066531 A

For example, in the case where the X-ray tube voltage condition changes during imaging as in the above-described dual energy imaging, an optimal method for handling data of image reconstruction variables is required. That is, it is assumed that it is necessary to have a lot of image reconstruction variable data because there are many states in the transitional state where the shooting conditions change, and it is necessary to deal with such many image reconstruction variables. However, there is a problem in that it is necessary to make a major software change in handling the image reconstruction variable.
Further, when the types of correction data used for each photographing condition increase, there is a problem that calibration work for collecting daily correction data and calibration work at the time of maintenance inspection become longer.

  Therefore, an object of the present invention is to deal with a number of different imaging conditions such as changes in X-ray tube voltage conditions in the image reconstruction processing of the X-ray CT apparatus, and can easily perform processing using image reconstruction variables. An object is to provide an X-ray CT apparatus.

According to the X-ray CT apparatus of the first aspect of the present invention,
An X-ray CT apparatus for imaging a tomographic image of a subject, an imaging condition setting unit capable of setting a plurality of conditions for a predetermined imaging condition, and an X-ray generator and an X-ray using the imaging conditions X-ray data collection means for collecting X-ray projection data by scanning a subject while rotating a rotating unit having a line detector, and image reconstruction variables that differ depending on the plurality of conditions, at least representative X-ray data collection including an image reconstruction variable corresponding to various conditions attached to and stored in a set of data collected by the X-ray projection data collection means, and at least a condition other than the representative condition X-ray comprising image reconstruction means for performing processing including an image reconstruction unit including image reconstruction processing of the X-ray projection data based on an image reconstruction variable corresponding to an imaging condition used in CT Location, is that.
According to the X-ray CT apparatus of the second aspect of the present invention, in the first aspect,
The image reconstruction unit is configured to perform image reconstruction of the X-ray projection data based on an image reconstruction variable corresponding to an imaging condition used for the X-ray data collection including the representative condition and a condition other than the representative condition. An X-ray CT apparatus that performs processing including an image reconstruction unit including configuration processing.
Moreover, according to the X-ray CT apparatus of the 3rd viewpoint of this invention, in the 1st or 2nd viewpoint,
The predetermined imaging condition is an X-ray tube voltage condition, and the X-ray data collection unit collects X-ray projection data using a plurality of X-ray tube voltage conditions, and the representative condition is An X-ray CT apparatus characterized by a plurality of X-ray tube voltage conditions used for the X-ray data collection.
According to the X-ray CT apparatus of the fourth aspect of the present invention, in the X-ray CT apparatus of the second aspect,
The predetermined imaging condition is an X-ray tube voltage condition, and the X-ray data collection unit collects X-ray projection data using a plurality of X-ray tube voltage conditions, and the representative condition is The plurality of X-ray tube voltage conditions used for the X-ray data collection, and the conditions other than the representative conditions include a condition of a transient section accompanying switching of the plurality of X-ray tube voltage conditions. X-ray CT apparatus.
According to the X-ray CT apparatus of the fifth aspect of the present invention, in the X-ray CT apparatus of any one of the first to fourth aspects,
The image reconstruction unit calculates or selects an image reconstruction variable corresponding to the representative condition, obtains an image reconstruction variable corresponding to a condition other than the representative condition, and determines the image reconstruction variable. And an X-ray CT apparatus that performs processing using an image reconstruction variable corresponding to a condition other than the representative condition.
According to the X-ray CT apparatus of the sixth aspect of the present invention, in the X-ray CT apparatus of the fifth aspect,
The X-ray CT apparatus, wherein the image reconstruction unit adds the obtained image reconstruction variable to the data set and stores it in a storage unit.
According to the X-ray CT apparatus of the seventh aspect of the present invention, in the X-ray CT apparatus of any one of the first to fourth aspects,
The image reconstruction unit calculates or selects X-ray projection data obtained by processing using an image reconstruction variable corresponding to the representative condition, and corresponds to a condition other than the representative condition. An X-ray CT apparatus that performs processing using an image reconstruction variable corresponding to a condition other than the representative condition by obtaining X-ray projection data obtained by processing using a reconstruction variable That's it.
According to the X-ray CT apparatus of the eighth aspect of the present invention, in the X-ray CT apparatus of any of the fifth to seventh aspects,
The storage unit corresponds to a condition other than the representative condition, and is an X-ray CT apparatus that stores the calculation condition of the calculation.
According to the X-ray CT apparatus of the ninth aspect of the present invention, in the X-ray CT apparatus of the first to fourth aspects,
The image reconstruction unit stores the range of the imaging condition and an image reconstruction variable corresponding to the range of the imaging condition, and performs image reconstruction corresponding to a condition other than the representative condition included in the imaging condition range. An X-ray CT apparatus is characterized in that a configuration variable is obtained from an image reconstruction variable corresponding to the stored range.
According to a tenth aspect of the present invention, in the X-ray CT apparatus according to any one of the first to ninth aspects,
The X-ray CT apparatus is characterized in that the image reconstruction variable is correction data used for preprocessing of image reconstruction.
According to the X-ray CT apparatus of the eleventh aspect of the present invention, in the X-ray CT apparatus of the first or second aspect, the imaging conditions include an X-ray tube voltage, an X-ray focal position, a gantry rotation speed, And an X-ray CT apparatus selected from the opening width of the collimator.
According to the X-ray CT apparatus of the twelfth aspect of the present invention,
A data file including X-ray projection data used in the X-ray CT apparatus according to the first aspect, including the X-ray projection data and image reconstruction variables corresponding to the plurality of imaging conditions. Is a data file.

  According to the X-ray CT apparatus of the present invention, it is possible to realize an X-ray CT apparatus that can easily deal with a number of different imaging conditions such as changes in X-ray tube voltage conditions and can easily perform processing using image reconstruction variables. There is.

<Overall configuration of X-ray CT apparatus>
FIG. 1 is a configuration block diagram of an X-ray CT apparatus 100 according to an embodiment of the present invention. The X-ray CT apparatus 100 includes an operation console 1, an imaging table 10, and a scanning gantry 20.

  The operation console 1 includes an input device 2 such as a keyboard or a mouse that receives input from the operator, a central processing unit 3 that executes setting of imaging conditions, image reconstruction processing, and the like, and X-ray projection data collected by the scanning gantry 20. And a data collection buffer 5 for collecting data. Furthermore, the operation console 1 includes a monitor 6 that displays a tomographic image reconstructed from projection data, and a storage device 7 that stores a program, X-ray projection data, or X-ray tomographic image. The imaging table 10 includes a cradle 12 on which the subject HB is placed and taken in and out of the opening of the scanning gantry 20. The cradle 12 is moved up and down and moved linearly by a motor built in the imaging table 10.

  The gantry rotating unit 15 of the scanning gantry 20 includes an X-ray tube 21, an X-ray controller 22, a collimator 23, a beam forming X-ray filter 28, a multi-row X-ray detector 24, and a data acquisition device (DAS: Data Acquisition System) 25. The gantry rotating part 15 is rotatable via a bearing. When a rotation motor (not shown) rotates, the rotation is transmitted to the gantry rotation unit 15 via a belt (not shown), and the gantry rotation unit 15 rotates. Further, the scanning gantry 20 includes a rotating unit controller 26 that controls the gantry rotating unit 15 rotating around the body axis of the subject HB, and a control controller 29 that exchanges control signals and the like with the operation console 1 and the imaging table 10. It is equipped with. The beam forming X-ray filter 28 has the thinnest filter thickness in the X-ray direction toward the center of rotation, which is the imaging center, and the filter thickness increases toward the periphery so that X-rays can be absorbed more. X-ray filter.

The central processing unit 3 includes an imaging condition setting unit 31 and an image reconstruction unit 32.
The imaging condition setting unit 31 sets various imaging conditions (X-ray tube current, X-ray tube voltage, collimator aperture width, gantry rotation speed, various image reconstruction conditions, etc.) of the X-ray CT apparatus 100. The photographing condition is input from the input device 2. In addition, a signal based on the condition set by the imaging condition setting unit 31 is transmitted to the controller 29, and the image reconstruction condition is transmitted to the image reconstruction unit 32.
The storage device 7 reconstructs the X-ray projection data collected by the X-ray projection data collection means and the set of data accompanied by image reconstruction variables corresponding to typical conditions, and X-ray projection data. The obtained tomographic image data and the like are stored.

As preprocessing, the image reconstruction unit 32 performs, for example, offset correction, logarithmic conversion of X-ray projection data, X-ray scattering correction, X-ray detector sensitivity correction, beam hardening correction, and the like on X-ray projection data. Process that includes. The correction data used for these corrections is different data depending on, for example, the difference in X-ray tube voltage. Offset correction is correction for removing noise caused by DAS. X-ray scattering correction is correction for reducing the influence of scattered X-rays among detected X-rays. Sensitivity correction is correction for reducing variations in sensitivity that occur in the plane of the X-ray detector. Beam hardening correction is a phenomenon in which X-ray absorption changes depending on the transmission thickness even with the same material, and the CT value (brightness) on the CT image changes. In particular, the energy distribution of the radiation transmitted through the subject HB is high energy. The correction of the effect of so-called beam hardening, which is biased to the side, is performed for the slice direction and the channel direction of the X-ray projection data.
In the present embodiment, the image reconstruction unit 32 performs the X based on the image reconstruction variable corresponding to the imaging condition used for the X-ray data collection including the representative condition and the condition other than the representative condition. Image reconstruction processing of line projection data is performed.

  The image reconstruction unit 32 receives the projection data subjected to the beam hardening process, and reconstructs an image based on the projection data. The projection data is subjected to a fast Fourier transform (FFT) to be converted into the frequency domain, and a reconstruction function Kernel (j) is superimposed on the projection data to perform an inverse Fourier transform. Then, the image reconstruction unit 32 performs a three-dimensional backprojection process on the projection data obtained by superimposing the reconstruction function Kernel (j), and obtains a tomogram (for each body axis direction (Z direction) of the subject HB. xy plane). The image reconstruction unit 32 stores this tomographic image in the storage device 7.

<Operation flowchart of X-ray CT apparatus>
FIG. 2 is a flowchart showing an outline of the operation of the X-ray CT apparatus 100 of the present embodiment.
In step P1, the subject HB is placed on the cradle 12 and aligned. Here, the subject HB placed on the cradle 12 aligns the slice light center position of the scanning gantry 20 with the reference point of each part. Then, scout image collection is performed. In scout imaging, the X-ray tube 21 and the multi-row X-ray detector 24 are fixed, and the X-ray projection data is collected while the cradle 12 is moved linearly. Here, the scout image is usually photographed at view angle positions of 0 degrees and 90 degrees. The right side in FIG. 2 is an example of a scout image 41 obtained by photographing the vicinity of the chest at 0 degrees. From this scout image 41, it is possible to plan a tomographic image capturing position.

  In step P2, the photographing condition is set while displaying the position and size of the tomographic image to be photographed on the scout image 41. The dotted line shown in the scout image 41 is the position of the tomographic image. In the present embodiment, a plurality of scan patterns such as a conventional scan, a helical scan, a variable pitch helical scan, and a helical shuttle scan are provided. The conventional scan is a scan method in which the X-ray tube 21 and the multi-row X-ray detector 24 are rotated to acquire X-ray projection data every time the cradle 12 is moved at a predetermined interval in the z-axis direction. The helical scan is an imaging method for collecting X-ray projection data by moving the cradle 12 at a constant speed while the X-ray tube 21 and the multi-row X-ray detector 24 rotate. The variable pitch helical scan is an imaging method for collecting X-ray projection data by changing the speed of the cradle 12 while rotating the X-ray tube 21 and the multi-row X-ray detector 24 as in the helical scan. In the helical shuttle scan, the cradle 12 is accelerated and decelerated while rotating the X-ray tube 21 and the multi-row X-ray detector 24 as in the helical scan, and reciprocally moves in the positive direction of the z-axis or the negative direction of the z-axis. This is a scanning method for collecting X-ray projection data. When these plural radiographs are set, X-ray dose information as a whole is displayed.

  In setting the tomographic imaging conditions, the exposure of the subject HB can be optimized by using the automatic exposure mechanism of the X-ray CT apparatus 100. Further, in this tomographic imaging condition setting, for the so-called dual energy imaging tomographic imaging, the imaging condition of the X-ray tube 21 with a low X-ray tube voltage, for example, 80 kV, and the high X-ray tube voltage, for example, 140 kV, Shooting conditions can be set.

  In step P3 to step P9, tomographic imaging is performed. In step P3, X-ray data collection is performed. Here, when collecting data by helical scanning, the data collection operation of X-ray projection data is performed while rotating the gantry rotating unit 15 around the subject HB and moving the cradle 12 on the imaging table 10 linearly. I do. Then, the z-direction coordinates of the X-ray projection data D0 (view, j, i) (j = 1 to ROW, i = 1 to CH) represented by the view angle view, the detector row number j, and the channel number i A position Ztable (view) is added. As described above, in the helical scan, X-ray projection data collection within a constant speed range is performed.

In Steps P4 to P9, image reconstruction processing using the image reconstruction unit 32 is performed.
In step P4, the X-ray projection data D0 (view, j, i) is preprocessed and converted into projection data. Specifically, offset correction is performed, logarithmic conversion is performed, X-ray dose correction is performed, and sensitivity correction is performed.
In Step P5, beam hardening correction is performed. Here, beam hardening correction is performed on the preprocessed projection data D1 (view, j, i). At this time, since independent beam hardening correction can be performed for each j column of the detector, if the tube voltage of the X-ray tube 21 varies depending on the imaging conditions, the difference in the X-ray energy characteristics of the detector for each column. Can be corrected. In the present embodiment, the beam hardening correction process is changed according to the profile area, ellipticity, etc. of the subject HB.

  In step P6, z filter convolution processing is performed. Here, a z-filter convolution process for applying a filter in the z direction (column direction) is performed on the projection data D11 (view, j, i) subjected to beam hardening correction.

In step P7, reconstruction function superimposition processing is performed. That is, the Fourier transform (Fourier Transform) for transforming the X-ray projection data into the frequency domain is performed, the reconstruction function is multiplied, and the inverse Fourier transform is performed.
In step P8, a three-dimensional backprojection process is performed. Here, three-dimensional backprojection processing is performed on the projection data D3 (view, j, i) subjected to the reconstruction function superimposition processing to obtain backprojection data D3 (x, y, z). The image to be reconstructed is a plane perpendicular to the z axis. A three-dimensional image is reconstructed on the xy plane. The following reconstruction area P is assumed to be parallel to the xy plane.

In step P9, post-processing is performed. Post-processing such as image filter superimposition and CT value conversion is performed on the backprojection data D3 (x, y, z) to obtain a tomographic image G.
In step P10, the tomographic image reconstructed is displayed. As an example of the tomographic image, a tomographic image G is shown on the right side of FIG.

Hereinafter, a method for obtaining an image reconstruction variable corresponding to a condition other than a typical condition will be described in detail using an embodiment.
(Example 1)

In the first embodiment, X-ray projection data obtained by performing pre-processing such as beam hardening correction, which differs depending on each X-ray tube voltage, on X-ray projection data acquired using a plurality of X-ray tube voltages in so-called dual energy imaging. This is an example in which preprocessing using correction data different from the correction data for correction stored in the storage unit is performed by weighted addition. Here, the correction data for preprocessing is an example of the “image reconstruction variable” in the present invention.
In the present embodiment, in the above-described X-ray CT apparatus, the central processing unit 3 further including a dual energy image reconstruction unit (not shown) can be used. The dual energy image reconstruction unit is a so-called dual energy imaging, which is a two-dimensional distribution tomographic image of X-ray tube voltage-dependent information related to the distribution of atoms, from projection data or tomographic images with a low X-ray tube voltage and a high X-ray tube voltage. The tomogram of the image is reconstructed.
First, in the dual energy imaging, a flow in the case where imaging is continuously performed with two types of X-ray tube voltages is shown.
FIG. 3A shows “dual energy imaging by conventional scanning at a plurality of coordinate positions in the z direction.
In this imaging method, at a certain z-axis direction coordinate position z = z1, X-ray data is collected for 360 degrees and one rotation at a low X-ray tube voltage LkV, and then 360 degrees and one rotation is performed at a high X-ray tube voltage HkV. X-ray data collection. This is similarly repeated at the z-axis direction coordinate positions z = z2 and z = z3.

FIG. 4 is a flowchart for explaining the X-ray data collection in FIG.
In step D1, scanning is performed at an X-ray tube voltage of 80 kV (kV1), and X-ray data collection is started.
In step D2, it is determined whether the X-ray tube voltage is 80 kV and has reached a constant view angle (360 * t7 / t0 degrees). If YES, go to step D3, and if NO, repeat step D2.
In step D3, X-ray data collection is performed while increasing at an X-ray tube voltage increase rate (kV2-kV1) / t6.
In step D4, it is determined whether the X-ray tube voltage has reached 140 kV (kV2). If YES, the process goes to step D5, and if NO, step D4 is repeated once more.
In step D5, scanning is performed at an X-ray tube voltage of 140 kV (kV2), and X-ray data is collected.
In step D6, it is determined whether the view angle is 720 degrees (two rotations). If YES, the process ends. If NO, step D5 is repeated once more.

  In this dual energy imaging, each tomogram G is imaged at t0 seconds per rotation, for example, when 0.35 seconds per rotation, 2 * t0 seconds = 0.70 seconds at 2 rotations, 0.7 seconds Thus, X-ray tube voltages kV1 and k2V are imaged. If it is 0.7 seconds, the body movement is considerably suppressed, and the alignment of the tomographic image G and the alignment of the projection data can be expected to be performed without any problem.

  In this case, the time t6 until the X-ray tube voltage rises from 80 kV (kV1) to 140 kV (kV2) is considered to be, for example, about t6 = 0.1 seconds. It depends on the device because it depends.

Here, the portions where the X-ray tube voltage is constant 80 kV (kV1) and 140 kV (kV2) may be processed with correction data of 80 kV and 140 kV, respectively, but continuously change from 80 kV to 140 kV. In the transition period, there is no preprocessing correction data corresponding to each X-ray tube voltage. However, in order to perform accurate image reconstruction, it is necessary to correct a portion where the X-ray tube voltage continuously changes from 80 kV (kV1) to 140 kV (kV2) using appropriate correction data.
In this embodiment, correction data corresponding to the X-ray tube voltage 80 kV (kV1) and correction data corresponding to the X-ray tube voltage 140 kV (kV2) are attached to the X-ray projection data file obtained above. The
The X-ray projection data from the X-ray tube voltage of 80 kV (kV1) to the X-ray tube voltage of 140 kV (kV2) corresponds to the correction data corresponding to the X-ray projection data and the transient section. For the X-ray projection data, an X-ray tube voltage threshold (for example, 110 kV) is set, and for X-ray projection data having an X-ray tube voltage lower than the X-ray tube voltage threshold, the X-ray tube voltage is 80 kV. For correction data corresponding to (kV1) and X-ray projection data having an X-ray tube voltage higher than the threshold value of the X-ray tube voltage, correction processing using correction data corresponding to the X-ray tube voltage 140 kV (kV2) is performed. Do.
The correction data used in the preprocessing is greatly influenced by the X-ray tube voltage, and has data for each X-ray tube voltage. The data for beam hardening correction, the data for X-ray scattering correction, and the X-ray detection For example, the sensitivity correction data.
(Example 2)

Usually, in the X-ray CT apparatus, correction data for preprocessing depending on the X-ray tube voltage is not included in all X-ray tube voltages, but correction data for about three to four types of X-ray tube voltages in advance. have.
For example, in the case of FIG. 3B, details of the scan timing at one coordinate position in the z-axis direction when correction data is provided at 80 kV, 100 kV, 120 kV, and 140 kV are shown.
In this dual energy imaging, the time required for one scan is divided into the time from t0 to t7 and determined as follows.
t0: Time for one rotation of the data acquisition system, t0 = t7 + t2 t1: Time until the X-ray tube voltage rises from 80 kV (kV1) to 100 kV t2: The X-ray tube voltage rises from 80 kV (kV1) to 110 kV T3: Time until the X-ray tube voltage increases from 100 kV to 120 kV t4: Time until the X-ray tube voltage increases from 110 kV to 140 kV (kV2) t5: X-ray tube voltage from 120 kV to 140 kV ( t6: Time until the X-ray tube voltage increases from 80 kV (kV1) to 140 kV (kV2) t7: Time from the start of scanning to the X-ray tube voltage of 80 kV (kV1) Therefore, in this embodiment, correction data of 80 kV and 140 kV are used for the transition section in the first embodiment. Replacing the law, between 80kV and 140 kV, using the correction data corresponding to 100kV and 120 kV.
That is, in this embodiment, correction data corresponding to an X-ray tube voltage of 80 kV (kV1), correction data corresponding to an X-ray tube voltage of 100 kV (kV1), and an X-ray tube for the X-ray projection data file. Correction data corresponding to the voltage 120 kV (kV1) and correction data corresponding to the X-ray tube voltage 140 kV (kV2) are attached.
The X-ray projection data from the X-ray tube voltage of 80 kV (kV1) to the X-ray tube voltage of 140 kV (kV2) corresponds to the correction data corresponding to the X-ray projection data and the transient section. For the X-ray projection data, an X-ray tube voltage threshold (for example, 110 kV) is set, and for X-ray projection data with an X-ray tube voltage lower than the X-ray tube voltage threshold, the X-ray tube voltage is 100 kV. For correction data corresponding to (kV1) and X-ray projection data having an X-ray tube voltage higher than the threshold of the X-ray tube voltage, correction processing using correction data corresponding to the X-ray tube voltage 120 kV (kV2) is performed. Do.
(Example 3)

In the present embodiment, a method of obtaining corrected X-ray projection data corresponding to a transient section by weighted addition of X-ray projection data corrected using correction data corresponding to kV1 and kV2 It is.
The X-ray projection data of the channel ch, column row, and view view collected at the X-ray tube voltage kV is D (kV, ch, row, view), and the result of preprocessing this X-ray projection data is Prep_D (kV, ch, row, view). At this time, in the storage device, correction data for the X-ray tube voltage kV does not exist in the storage device, and correction data exists in the storage device for the X-ray tube voltages kV1 and kV2. . Further, it is assumed that kV1 <kV <kV2. In addition, X-ray projection data that has been pre-processed with correction data of the X-ray tube voltages kV1 and kV2 is represented by Prep_D (kV1, ch, row, view) and Prep_D (kV2, ch, row, view).

In this case, the preprocessed X-ray projection data Prep_D (kV, ch, row, view) of the X-ray tube voltage kV is expressed by the following (Formula 1).
... (Formula 1)
In this way, by performing weighted addition processing of the respective X-ray projection data corrected by the correction data of the X-ray tube voltages kV1 and kV2, the pre-processed X-ray projection data corresponding to the X-ray tube voltage kV Prep_D (kV, ch, row, view) is obtained. This pre-processed X-ray projection data Prep_D (kV, ch, row, view) can be regarded as pre-processing using correction data corresponding to kV.
In addition, the example of the weighting coefficients w1 and w2 at this time is as the following (Formula 2) and (Formula 3).
... (Formula 2)
... (Formula 3)

X-ray projection data collected at the X-ray tube voltage kV is corrected by performing weighted addition processing on the X-ray projection data collected at the X-ray tube voltages kV1 and kV2 near kV and corrected by the respective correction data. can do. An example of the processing is shown below. The processing positions (SH1 to SH3) are shown in FIG.
Process 1 (SH1): When the X-ray tube voltage is between 80 kV and 100 kV, the preprocessed X-ray projection data of 80 kV and 100 kV is subjected to weighted addition processing.
Process 2 (SH2): When the X-ray tube voltage is between 100 kV and 120 kV, the preprocessed X-ray projection data of 100 kV and 120 kV is subjected to weighted addition processing.
Process 3 (SH3): When the X-ray tube voltage is between 120 kV and 140 kV, the pre-processed X-ray projection data of 120 kV and 140 kV is subjected to weighted addition processing.

Specifically, in the case of the X-ray tube voltage of 105 kV, the preprocessed X-ray projection data Prep_D of 105 kV is obtained from the X-ray projection data preprocessed with the correction data of 100 kV and 120 kV (Equation 4) It is required as follows.
... (Formula 4)

FIG. 5 is a flowchart showing image reconstruction processing including preprocessing of X-ray projection data for which correction data is not stored in the storage device, using the image reconstruction unit 32, corresponding to the X-ray tube voltage kV.
In step M1, the view number j = 1.
In step M2, the X-ray projection data of j view is read, and the X-ray tube voltage kV is read from the incidental information of the X-ray projection data.
In step M3, it is determined whether the correction data corresponding to the X-ray tube voltage kV is in the storage device. If YES, the process goes to step M4, and if NO, the process goes to step M10.
In step M4, preprocessing is performed from correction data of the X-ray tube voltage kV.
In step M5, it is determined whether j = N view. If YES, the process goes to step M6, and if NO, the process goes to step M9.
In step M6, reconstruction function superimposition processing is performed.
In step M7, a three-dimensional backprojection process is performed.
In step M8, post-processing is performed.

In step M9, j = j + 1 is set, and the process returns to step M2.
In Step M10, it is determined whether there is correction data of the X-ray tube voltage kV1 lower than the X-ray tube voltage kV. If YES, the process goes to Step M11, and if NO, the process goes to Step M15.
In step M11, correction data for the low X-ray tube voltage kV1 is read, and pre-processing of j-view X-ray projection data is performed with the low X-ray tube voltage kV1.
In Step M12, it is determined whether there is correction data for the X-ray tube voltage kV2 higher than the X-ray tube voltage kV. If YES, the process goes to Step M13, and if NO, the process goes to Step M15.
In step M13, the correction data of the high X-ray tube voltage kV2 is read, and the pre-processing of the j-view X-ray projection data is performed with the high X-ray tube voltage kV2.
In step M14, the preprocessing result Prep1 corrected with the X-ray tube voltage kV1 and the preprocessing result Prep2 corrected with the X-ray tube voltage kV2 are weighted and added to obtain the preprocessing result Prep_D with the X-ray tube voltage kV. . After this, go to step M5. That is, the processes shown in the above (Formula 2), (Formula 3), and (Formula 4) are performed.
In step M15, it is determined that the data collection is abnormal, and the process ends.

  As shown in step M2, in the image reconstruction unit 32, information on the X-ray tube voltage kV is included in the incidental information of the X-ray projection data. Since the X-ray tube voltage of each view is recorded in real time at the time of data collection, correct preprocessing can be performed.

As described above, the X-ray projection data Prep_D (kV1) pre-processed at the low X-ray tube voltage kV1 by searching for the nearby X-ray tube voltages kV1 and kV2 having correction data for the increasing X-ray tube voltage kV. , Ch, row, view) and X-ray projection data Prep_D (kV2, ch, row, view) preprocessed at a high X-ray tube voltage kV2. Then, the image reconstruction variable processing unit 39 can obtain the preprocessed X-ray projection data of the X-ray tube voltage kV only by changing the weighting coefficients w1 and w2 according to (Equation 2) and (Equation 3). .
Example 4

  In this embodiment, an image reconstruction variable corresponding to an X-ray tube voltage different from an image reconstruction variable corresponding to a plurality of X-ray tube voltages stored in the storage unit is obtained, and X It is an example which processes a line projection data.

FIG. 6A conceptually shows the image reconstruction variable group Par (kVi) at each X-ray tube voltage.
The correction data processing unit 39 at this time has correction data Par (kV1), Par (kV2i), Par (kV3), and Par (kV4) of the X-ray tube voltages kV1, kV2, kV3, and kV4 in advance.

On the other hand, when the X-ray tube voltage changes during imaging, the image quality of the tomographic image is improved by increasing the correction data corresponding to the X-ray tube voltage as necessary. At this time, the image reconstruction unit 32 includes an image reconstruction variable addition tool or an image reconstruction variable structure editing tool, and obtains correction data corresponding to the added X-ray tube voltage. For example, when the configuration of the correction data is changed, the correction data of the added X-ray tube voltage is acquired from the X-ray projection data or calculated from the existing correction data and selected to obtain the correction data. In this way, optimal correction data collection can be performed by appropriately changing the points of the correction data from time to time.
In addition, when using X-ray projection data with an X-ray CT apparatus having the same software, if all the correction data used in the pre-processing is added to the X-ray projection data, the pre-processing is performed under optimum correction data conditions. It can be performed.
For example, as shown in FIG. 6B, when a new X-ray tube voltage kV5 is added, the correction data of the new X-ray tube voltage is added to the image reconstruction variable group.

FIG. 6C shows a case where common correction data exists across the X-ray tube voltage.
When the correction data of the X-ray tube voltage kV1 is common to the X-ray tube voltage kV2, the X-ray tube voltage kV3, and the X-ray tube voltage kV4, the image reconstruction unit 32 displays “X-ray tube” on the software setting file. “Same as voltage kV1”. As described above, the correction data setting can be simplified if the software of the correction data processing unit 39 can interpret the setting file.

For further convenience, X-ray tube voltage correction data may be added to the setting file in such a way that the image reconstruction variable does not depend on the X-ray tube voltage.
For example, FIG. 6D shows a case where the correction data is used as an image reconstruction variable that does not depend on the X-ray tube voltage, or correction data of the X-ray tube voltage. In this case, the added new correction is operated on the image reconstruction, and is first operated without depending on the X-ray tube voltage. If a phenomenon depending on the X-ray tube voltage is observed, it is sufficient to finely adjust the correction data different for each X-ray tube voltage.
A specific example using the correction data setting file CDR as described above and software that interprets the setting file will be described.

  For example, FIG. 7A shows a case where the X-ray tube voltage is changed during one scan in dual energy imaging. At this time, one scan is performed by collecting X-ray data of two rotations or less than two rotations. In this case, the X-ray tube voltage starts at a low tube voltage of 80 kV and continuously changes to a high X-ray tube voltage of 140 kV during the scan.

FIG. 7B is an example of the correction data setting file CDR used in this case. (* 1) in FIG. 7B indicates correction data used in normal photographing. For the 80 kV, 100 kV, 120 kV, and 140 kV portions where the X-ray tube voltage does not change, pre-processing is performed using (* 1).
For the transition period where the X-ray tube voltage changes, correction data of 90 kV, 110 kV, and 130 kV are added to the part (* 2). In addition, since the X-ray tube voltage continuously changes, values such as 81 kV, 82 kV, 83 kV,. For this reason, (* 3) shows an interpolation mode or a weighted addition mode (calculation condition) used for obtaining values between the added correction data by interpolation processing or weighted addition processing.
Thus, the accuracy of the preprocessing corresponding to each X-ray tube voltage is improved by performing interpolation processing or weighted addition processing in addition to the added correction data. Further, the software of the correction data processing unit 39 interprets the correction data setting file CDR, always understands the structure, recognizes the image reconstruction variable and the increase / decrease of the correction data, and is applied to which shooting condition. Let me recognize it.

FIG. 8 shows another correction data setting file CDR.
The structure of the X-ray projection data in the dual energy imaging is added to the X-ray projection data as shown in FIG. 8A, correction data corresponding to the X-ray tube voltages of 80 kV and 140 kV, and the added X-ray tube. What is necessary is just to have as an X-ray projection data file including the correction data from voltage kV1 to X-ray tube voltage kVN.
FIG. 8B shows a modification of the correction data setting file CDR.
In FIG. 7B, each X-ray tube voltage has a variable one by one in each column indicating each image reconstruction variable or each correction data. For example, in the case of a BHC coefficient indicating a beam hardening correction coefficient, beam hardening correction coefficients b10, b11, b12, and b13 are included at an X-ray tube voltage of 80 kV. At 100 kV, beam hardening correction coefficients b20, b21, b22, and b23 are included. On the other hand, in the case of [120, 130; b1i] in FIG. 8B, it is said that “beam hardening correction coefficients b10, b11, b12, b13 are used from 120 kV to 130 kV”. The rules and grammar on the symbol should be established. In this way, the correction data setting file CDR can also describe the correction data more easily.

  Although X-ray tube voltage correction data and image reconstruction variables are not obtained by weighted addition processing for X-ray tube voltages not collected in advance, some of them may be obtained by weighted addition processing. it can. An example is shown below.

FIG. 9 shows a beam hardening correction curve at each X-ray tube voltage.
For example, the beam hardening correction coefficient at an X-ray tube voltage of 90 kV is expected to be the one shown in FIG.
When this relationship is expressed in a polynomial form, the projection data corrected by the beam hardening is as shown in (Formula 5) when the X-ray tube voltage is 80 kV, and when the X-ray tube voltage is 100 kV, (Formula 6). It becomes like this.
... (Formula 5)
... (Formula 6)

The following (Formula 7) is predicted from (Formula 5) and (Formula 6).
... (Formula 7)
From the above relationship, it is predicted that the beam hardening correction coefficient can be expressed by (Formula 8), (Formula 9), and (Formula 10).
... (Formula 8)
... (Formula 9)
(Equation 10)
From these relational expressions, the interpolation value of the beam hardening correction coefficient can be predicted. As an example of k, k = 1/2 is considered and takes an intermediate value.
Such an interpolation method is a weighted addition process using only variables, such as a beam hardening correction coefficient, and correction data is obtained by an interpolation process, and the actual preprocessing is performed, and the result of the preprocessing is weighted addition processing. Or there is what is performed by interpolation processing.

FIG. 10 is an example of a flowchart for performing processing using both.
In step M21, the view number j = 1.
In step M22, the X-ray projection data of j view is read, and the X-ray tube voltage kV is read from the incidental information of the X-ray projection data.
In step M23, it is determined whether there is correction data corresponding to the X-ray tube voltage kV. If YES, the process goes to step M24, and if NO, the process goes to step M31.
In step M24, preprocessing is performed from correction data of the X-ray tube voltage kV.
In step M25, it is determined whether j = N view. If YES, the process goes to step M26, and if NO, the process goes to step M30. Here, N is the number of views of 360 degrees and one rotation.
In step M26, reconstruction function superimposition processing is performed.
In step M27, a three-dimensional backprojection process is performed.
In step M28, post-processing is performed.
In step M29, an image is displayed.

In step M30, j = j + 1 is set, and the process returns to step M22.
In step M31, it is determined whether the variable can be weighted and added to the X-ray tube voltage. If YES, the process goes to step M32, and if NO, the process goes to step M36.
In step M32, a weighted addition or interpolation method is read.
In step M33, a weighted addition process is performed based on a nearby X-ray tube voltage variable.
In step M34, it is determined whether there is a new variable in addition to the variable obtained in step M33. If YES, the process goes to step M35, and if NO, the process returns to step M25.
In step M35, the variable obtained in step M33 is registered.
In step M36, preprocessing is performed based on a variable of the nearby X-ray tube voltage, and the result is subjected to weighted addition processing. After this process, the process returns to step M25.

Pre-processing is performed using the correction data obtained as described above.
(Example 5)

In the fifth embodiment, when the correction data in the beam hardening correction is changed, the software automatically recognizes the change.
The beam hardening correction can be expressed by a cubic polynomial as follows.
... (Formula 11)
If this is subjected to beam hardening correction with a quartic or higher correction equation as in (Equation 12) below, the accuracy may be further improved.
... (Formula 12)
In the case of (Formula 11), as correction coefficient data, in the case of (B ₀ , B ₁ , B ₂ ), in the case of (Formula 12), in a vector format such as (B ₀ , B ₁ , B ₂ , B ₃ ) Will have.

FIG. 11 shows the flow of processing for recognizing the order of beam hardening correction.
The flow of each step process is as follows.
In step B1, the correction data processing unit 39 counts the number of beam hardening correction coefficient data.
In step B2, the correction data processing unit 39 determines whether there are three beam hardening correction coefficient data. If YES, the process goes to step B3, and if NO, the process goes to step B4.
In step B3, the beam hardening processing unit 33 performs beam hardening correction of a third-order polynomial.
In step B4, the correction data processing unit 39 determines whether there are four correction coefficient data. If YES, the process goes to step B5, and if NO, the process goes to step B6.
In step B5, the beam hardening processing unit 33 performs beam hardening correction of a fourth-order polynomial.
In step B6, error processing is performed as a beam hardening correction coefficient error.
In this case, it is assumed that a third-order polynomial and a fourth-order polynomial are used. However, when beam hardening correction coefficient data of 5th order or higher is given, correction can be performed using a polynomial of 5th order or higher. Even more versatility can be obtained.
As described above, the correction data processing unit 39 can correctly perform the correction by automatically recognizing the correction that the correction data means.
(Example 6)

In the sixth embodiment, an example in which imaging is performed using a gantry rotation speed without correction data as the gantry rotation speed of the scanning gantry 20 and image reconstruction is performed using correction data of the nearest gantry rotation speed is shown. .
FIG. 12 is a diagram showing the relationship between the gantry rotation speed s1 ′ without correction data and the gantry rotation speeds s1 and s2 with correction data. As shown in FIG. 12, corresponding to the gantry rotation speeds s1 and s2, each X-ray detector configuration, each X-ray beam aperture, the size of the X-ray focal point, and a correction for each of a plurality of beam forming filters. It is assumed that correction data such as phantom X-ray data, correction air X-ray projection data, and beam hardening correction data are collected in advance in the correction data collection mode. In contrast, s1 ′, which is the gantry rotation speed between s1 and s2, has no correction data. At this time, it is assumed that the number of X-ray projection data views is the same at the gantry rotation speed s1 and at the gantry rotation speed s1 '.
In this case, since the gantry rotation speeds s1's1 and s2 are different from each other, the rotation gantry 20 is rotated or the rotation of the X-ray tube 21 and the multi-row X-ray detector 24 is deformed due to deflection. The correction data is different.

Considering the difference in deformation due to the rotation vibration of the scanning gantry 20 or the deflection of the gantry rotating part 15 that supports the X-ray tube 21 and the multi-row X-ray detector 24, correction data at the gantry rotation speed s1 ' Will have to collect. However, this difference in gantry rotation speed has little effect on image quality. For this reason, if it is judged clinically that the difference in image quality is within the allowable range, the correction data of the gantry rotation speed s1 ′ can be substituted with the correction data of the gantry rotation speed s1.
The correction data corresponding to the gantry rotation speed is obtained by evaluating the image quality of the tomographic image G between the gantry rotation speeds s1 and s2, so that up to which range the correction data of the gantry rotation speed s1 is substituted. Determines whether the correction data of the gantry rotation speed s2 can be substituted. In this way, the range in which one piece of correction data is used can be as wide as possible, and the number of types of correction data can be reduced.
The X-ray projection data in this case can be reconstructed in the image reconstruction process by adding correction data for the gantry rotation speed s1 to the X-ray projection data for the gantry rotation speed s1 ′.

FIG. 13 shows a flow of processing for performing image reconstruction of the gantry rotation speed s1 ′ using the correction data of the gantry rotation speed s1.
In step V1, the gantry rotation speed s1 is set on the photographing condition setting screen.
In step V2, scanning is started.
In step V3, the correction data processing unit 39 attaches correction data for the gantry rotation speed s1 to the X-ray projection data for the gantry rotation speed s1 '.
In step V4, the scan ends and X-ray projection data collection is completed.
In step V5, image reconstruction is started.
In step V6, correction is performed using correction data for the gantry rotation speed s1, and image reconstruction is performed.
In step V7, the tomographic image G is displayed after the completion of image reconstruction.
In this case, since it is known that the gantry rotation speed is different from s1 and s1 ′, the gantry rotation speed s1 correction data is partially changed and corrected for the gantry rotation speed s1 ′ correction data. It is technically possible to perform the correction.
(Example 7)

In the seventh embodiment, the gantry rotation speed changes during the scan, and the X-ray projection data is collected and image reconstruction is performed under the imaging conditions without such changing gantry rotation speed correction data. Show.
In FIGS. 14A and 14B, it is assumed that correction data is collected in advance in the correction data collection mode at gantry rotation speeds s3 and s4. FIG. 14A shows an example in which the gantry rotation speed changes with a sine wave as EX1 and EX2. FIG. 14B shows an example in which the gantry rotation speed increases.

For example, when the heart is imaged, the heart rate often rises as the imaging progresses compared to the time when the imaging starts since the subject normally holds his breath. For example, the heart rate rises from 60 bpm (beat per minute) to 70 bpm.
In this case, the optimum gantry rotation speed increases as the heart rate increases. A synchronized tomographic image G with good image quality can be easily taken by changing the gantry rotation speed as shown in FIG.
In the X-ray projection data when the gantry rotation speed changes between s3 ′ and s4 ′, as shown in FIG. 14C, correction data for the gantry rotation speed s3 and the X of the gantry rotation speed s4. For line projection data, correction data for the gantry rotation speed s3 can be selected and used as long as it is from the gantry rotation speed s3 to the gantry rotation speed s3 ′. Further, correction data for the gantry rotation speed s4 can be selected from the gantry rotation speed s4 to the gantry rotation speed s4 ′.
On the other hand, from the gantry rotation speed s3 ′ to the gantry rotation speed s4 ′, correction data for the gantry rotation speeds s3 ′ to s4 ′ is obtained based on correction data for the gantry rotation speed s3 and the gantry rotation speed s4. Based on the correction data, the image reconstruction unit 34 performs image reconstruction. A method for obtaining the correction data will be described later.

FIG. 15 shows a flow of processing for reconstructing an image by obtaining correction data for the gantry rotation speeds s3 ′ to s4 ′ using the correction data for the gantry rotation speeds s3 and s4.
In step V11, the gantry rotation speed is set on the photographing condition setting screen.
In step V12, scanning is started.
In step V13, correction data for the gantry rotation speeds s3 and s4 is attached to the X-ray projection data.
In step V14, the scan is finished and the X-ray projection data collection is completed.
In Step V15, image reconstruction is started.
In step V16, correction data for the gantry rotation speeds s3 'to s4' is obtained from the correction data for the gantry rotation speeds s3 and s4.
In step V17, image reconstruction is performed using correction data for the gantry rotation speeds s3 and s4 and correction data for the gantry rotation speeds s3 'to s4'.
In step V18, the tomogram G is displayed after the image reconstruction is completed.

Next, how to obtain the correction data corresponding to the gantry rotation speed without the correction data will be described using the correction data of the deflection of the gantry part as an example.
First, the deflection of the gantry will be described. In FIG. 16A, when the X-ray tube 21 and the multi-row X-ray detector 24 which are X-ray generators in the case of the gantry rotating portion having the structure shown in the figure are rotated in the rotating unit 15 of the scanning gantry 20. Shows the deflection.
In this case, the gantry rotating unit 15 rotates in the xy plane perpendicular to the z direction around the rotation axis in the z direction. As a result of this rotation, the X-ray tube 21 and the multi-row X-ray detector 24 are deflected in the direction outside the rotational operation by centrifugal force.
The X-ray tube 21 and the multi-row X-ray detector 24 generate a displacement Ez due to deflection in the z direction and a displacement Ex due to deflection in the xy plane due to centrifugal force. Any deflection is approximated by a sine function (sin function).

  In step V16 described above, correction data for the gantry rotation speeds s3 'to s4' can be obtained from the correction data for the gantry rotation speeds s3 and s4. For example, as shown in FIG. 16B, the deflection of the gantry rotating unit 15 is the deflection E3 of the gantry rotating unit 15 at the gantry rotating speed s3 and the deflection E4 of the gantry rotating unit 15 at the gantry rotating speed s4. At this time, the deflection E of the gantry rotating portion 15 at the gantry rotating speeds s3 'to s4' is obtained as follows.

First, the deflection E3 of the gantry rotation unit 15 at the gantry rotation speed s3 is represented by (Formula 13).
... (Formula 13)
Further, the deflection E4 of the gantry rotating unit 15 at the gantry rotating speed s4 is represented by (Expression 14).
... (Formula 14)

When the gantry rotation speed is s3 ′, the constant k is determined as (Equation 15). However, 0 ≦ k ≦ 1.
... (Formula 15)
The deflection E of the gantry rotating unit 15 at this time is predicted as in the following (Equation 16).
... (Formula 16)
However, A, B, and cos α are represented by the following (Equation 17) to (Equation 19).
... (Formula 17)
... (Formula 18)
(Equation 19)
From the above, as can be seen from (Equation 16), the correction data processing unit 39 determines the deflection E of the gantry rotation unit 15 at the gantry rotation speeds s3 ′ to s4 ′ and the deflection of the gantry rotation unit 15 at the gantry rotation speeds s3 and s4. It can be obtained from the positional deviation correction data.
(Example 8)

  In the eighth embodiment, an embodiment related to correction data when using a technique for increasing the spatial resolution in the xy plane called by a name such as a flying X-ray focal point (focus) or wobble will be described. Flying X-ray focus (focus) imaging or wobble imaging is generally X-ray tomographic imaging by moving the X-ray focal position in the channel direction or Z direction within the xy plane.

  The X-ray focal position is controlled by switching the X-ray focal position for each view of the X-ray projection data, or switching the X-ray focal position for each of a plurality of views. At this time, in FIG. 17A, the X-ray focal position is quickly switched from the X-ray focal position F1 to the X-ray focal position F2, so the X-ray focal point control is rectangular. In FIG. 17B, when the X-ray focus is switched, each X-ray focus position is continuously moved in the channel direction, so that the trapezoidal control X-ray focus position control is performed. In FIG. 17C, when the X-ray focal point is switched, the X-ray focal point position control is a sine wave type control in the channel direction.

Thus, in any case, by moving the X-ray focal point in the channel direction, the X-ray beam passing through the image reconstruction center pixel is half the width of the X-ray beam at the X-ray focal point positions F0, F1, and F2. Each time, the position passing through the image reconstruction center pixel can be shifted. FIG. 18A shows this state.
Thereby, the image reconstruction center pixel in the image reconstruction area P can be decomposed to half the X-ray beam width by the sampling theorem, and high resolution can be realized. Further, when the X-ray focal position is moved, it is necessary to perform preprocessing and beam hardening correction.

For example, as shown in FIG. 18 (b), when the X-ray focal position moves in the channel direction from F1 to F0 to F2, the X-ray transmission path through the beam forming X-ray filter 28 is also different. The air correction data for correcting the sensitivity of each channel of the row X-ray detector 24 and the beam hardening correction coefficient are also different.
For example, when the air correction data of the X-ray focal positions F1 and F2 are known, the correction data processing unit 39 can obtain the air correction data of the X-ray focal position F0 based on these air correction data. .

Since the linearity can be maintained after logarithmic conversion of each X-ray projection data, the air correction data of the X-ray focal position F0 can be obtained as in the following (Equation 20). However, the air correction data distribution after logarithmic conversion of the X-ray focal position F1 is prfF1 (ch), the air correction data distribution after logarithmic conversion of the X-ray focal position F2 is prfF2 (ch), and the logarithmic conversion of the X-ray focal position F0. The subsequent air correction data distribution is assumed to be prfF0 (ch). The constant k is 0 ≦ k ≦ 1, and is defined by (Formula 21).
(Equation 20)
... (Formula 21)
In this way, the correction data processing unit 39 can predict and obtain correction data for each X-ray focal position from the correction data for the left and right end positions F1, F2 of the X-ray focal position.
In this case, in the three-dimensional backprojection processing, X-ray projection data corresponding to each pixel on the tomographic image is extracted in the X-ray transmission direction determined from each X-ray focal position, and the X-ray projection is performed in the X-ray East Asia k direction. It is preferable to perform a three-dimensional backprojection process that backprojects X-ray projection data onto each pixel on the tomographic image in the backprojection direction. In this case, image reconstruction processing is performed using each X-ray focal position information.

<Focus position measurement method>
Since each focal position can be corrected from the X-ray focal positions at the left and right ends, how to obtain the left and right end positions F1 and F2 will be described below.
In FIG. 19A, in order to make the positions of the X-ray focal positions F1 and F2 easy to understand, the positions are shifted excessively on the drawing. Actually, for example, when the X-ray focal point size is 0.5 mm in the channel direction × 0.5 mm in the z direction, the X-ray focal point positions F1 and F2 are shifted by, for example, ± 0.5 mm or ± 0.25 mm in the channel direction. It becomes.
The irradiation range at this time is shifted by d2 / d1 times as shown in FIG. 19A, and the surface of the multi-row X-ray detector 24 is irradiated. FIG. 19B shows a profile of the output of the X-ray detector in a certain column of the X-ray detector.
Here, it is shown that the deviation of the X-ray detector output profile at each X-ray detector position F1, F0, F2 is in the channel direction. FIG. 19C is an enlarged view of the left end portion of the X-ray detector profile of FIG. 19B.

The end point of the profile here is determined by its half value c0 / 2 obtained from the output value c0 of each X-ray detector. That is, the end point is determined by the full width half maximum (FWHM) of the peak value. xlF2 is the left end position of the FWHM determined from the X-ray detector profile in the case of the X-ray focal position F2, xlF0 is the left end position of FWHM determined from the X-ray detector profile in the case of the X-ray focal position F0, and xlF1 is the X-ray focal position. This is the left end position of the FWHM determined from the X-ray detector profile in the case of F1. In this case, the position of the X-ray focal point F0 is obtained from the following distance FD20 (Formula 22) from F2 to F0 and distance FD10 (Formula 23) from F1 to F0.
... (Formula 22)
(Equation 23)

In this way, the X-ray focal position can also be obtained from the shadow of the X-rays irradiated on the surface of the multi-row X-ray detector 24.

FIG. 20 is a flowchart showing how to obtain the X-ray focal position.
In step V31, scanning is started.
In Step V32, the correction data processing unit 39 determines whether there is a place where the left and right ends of the X-ray detector can measure the left and right ends of the X-ray detector. If YES, go to Step V33; Go to V37.
In step V33, the correction data processing unit 39 searches for the peak value c0 of the X-ray detector profile.
In Step V34, the half value c0 / 2 (FWHM) of the peak value of the X-ray detector profile is obtained.
In step V35, the correction data processing unit 39 obtains the FWHM positions xlF (left end position) and xrF (right end position) of the left and right ends of the X-ray detector profile.
In step V36, the correction data processing unit 39 obtains the distance between the X-ray focal points at that time from the X-ray focal point left and right ends F1, F2.
In step V37, the X-ray focal position is predicted from the front and rear views, or the X-ray focal position obtained from the X-ray generator is used.

In this flowchart, the X-ray focal position F0 is obtained from the X-ray detector profile being imaged. In this case, if the subject is too large and covers the entire X-ray detector surface, the left and right ends of the X-ray detector profile cannot be obtained correctly. As a countermeasure, the correction data processing unit 39 may predict the X-ray focal position in this case from the X-ray focal positions measured in the previous and next views.
As another method, in the X-ray tube capable of controlling the X-ray focal position, the X-ray focal position can be predicted from the control voltage value of the electron beam position control grid of the X-ray generator. Each X-ray focal position obtained in this way is also necessary in the three-dimensional backprojection process.
In the present embodiment, the example of air correction data has been described. However, as image reconstruction variables that change due to changes in the X-ray focal position, other preprocessing correction data such as beam hardening correction data is used. You may apply to.
Example 9

The ninth embodiment is an example in which the correction data is different depending on the opening width of the collimator for shaping the irradiated X-rays.
In this case, correction data 1 is stored as typical opening width correction data, and the data setting that the correction data 1 is commonly used in the range of the opening width 1 to the opening width 2 is [Opening 1, Opening 2; By setting as in [Data 1], correction data can be extracted for any aperture width including aperture width 1 to aperture width 2.
In addition, when there exists correction data for every opening width, it is preferable that it can change to the data setting which uses each correction data with respect to each opening width.

According to the above-described embodiment, it is possible to easily perform an image reconstruction process using an image reconstruction variable in correspondence with many different photographing conditions. In addition, the types of correction data used for each shooting condition are diversified, and in the current situation where calibration work for collecting daily correction data and calibration work during maintenance and inspection are taking a long time, various correction data can be used. Therefore, it is possible to perform image reconstruction suitable for various photographing conditions, and thus it is possible to prevent the calibration work from being further prolonged.
In the above embodiment, the medical X-ray CT apparatus is described as the original, but an industrial X-ray CT apparatus, or an X-ray CT-PET apparatus, an X-ray CT-SPECT apparatus combined with other apparatuses, etc. Can also be used.

1 is a configuration block diagram of an X-ray CT apparatus 100 according to an embodiment of the present invention. It is a flowchart which shows the outline | summary of operation | movement about the X-ray CT apparatus 100 of this embodiment. (A) is a figure which shows the dual energy imaging | photography by the conventional scan in several z direction coordinate position. (B) shows a time-resolved view of dual energy imaging by conventional scanning at one z-direction coordinate position. It is a flowchart of X-ray data collection in dual energy imaging. It is a flowchart of an image reconstruction process of X-ray projection data having a plurality of X-ray tube voltage values. (A) is a figure which shows the image reconstruction variable group in each conventional X-ray tube voltage. (B) is a diagram showing an image reconstruction variable group having a structure in which correction data and image reconstruction variables can be added in the direction of the X-ray tube voltage. (C) is a figure which shows the image reconstruction variable group which has the common variable which crossed the X-ray tube voltage. (D) is a diagram showing an image reconstruction variable group having a structure in which correction data and an image reconstruction variable can be added in the direction of the X-ray tube voltage variable. (A) is a figure which shows the imaging | photography in which the X-ray tube voltage changed. (B) is a figure which shows the example of the correction data setting file CDR. (A) is a figure which shows the format of X-ray projection data in case an X-ray tube voltage changes. (B) is a figure which shows the modification of the correction data setting file CDR. It is a figure which shows the correction curve in a some X-ray tube voltage value. It is a flowchart of the process in which the variable calculated | required by weighted addition process and the variable which is not calculated | required by weighted addition process are mixed. It is a flowchart of the process which recognizes the order of beam hardening correction. It is a figure which shows the relationship between gantry rotational speed s1 'without correction data, and gantry rotational speed s1, s2 with correction data. It is a flowchart which performs image reconstruction of gantry rotational speed s1 'using the correction data of gantry rotational speed s1. (A) is a figure which shows the relationship between gantry rotational speed s3 ', s4' without correction data, and gantry rotational speed s3, s4 with correction data. (B) is a figure which shows the case where a gantry rotational speed goes up. (C) is a diagram showing an X-ray projection data format when the gantry rotation speed changes. It is a flowchart which performs image reconstruction when a gantry rotational speed changes. (A) is a figure which shows the deflection | deviation during rotation of the gantry rotation part 15. FIG. (B) is a figure which shows the deflection | deviation E of the gantry rotation part 15 in the gantry rotation speed s3'-s4 'estimated from the deflection | deviations E3 and E4 of the gantry rotation part 15 of the gantry rotation speeds s3 and s4. (A) is a figure which shows X-ray focus position control of rectangular type control. (B) is a figure which shows the X-ray focus position control of trapezoid type | mold control. (C) is a figure which shows X-ray focus position control of sine wave type | mold control. (A) is a figure which shows high resolution of the tomogram G by a X-ray focus position. (B) is a figure which shows the positional relationship of the X-ray focus center position and X-ray focus F1, F2. (A) is a figure which shows each X-ray focal position in xy plane. (B) is a figure which shows the shift | offset | difference of the X-ray detector output profile in each X-ray focus position. (C) is a figure which shows the expanded X-ray detector output profile. It is a flowchart which performs image reconstruction in case an X-ray focal position changes.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 ... Operation console 2 ... Input device 3 ... Central processing unit 5 ... Data collection buffer 6 ... Monitor 7 ... Storage device 12 ... Cradle 15 ... Rotating part 20 ... Scanning gantry 21 ... X-ray tube 22 ... X-ray controller 23 ... Collimator 24 ... Multi-row X-ray detector 25 ... Data acquisition device (DAS)
26: Rotating unit controller 28: Beam forming X-ray filter 29 ... Controller 31 ... Imaging condition setting unit 32 ... Image reconstruction unit G ... Tomographic image

Claims (12)

An X-ray CT apparatus for taking a tomographic image of a subject,
Shooting condition setting means capable of setting a plurality of conditions for predetermined shooting conditions;
X-ray data collection means for scanning the subject and collecting X-ray projection data while rotating a rotating unit having an X-ray generator and an X-ray detector using the imaging conditions;
An image reconstruction variable that differs depending on the plurality of conditions, and at least an image reconstruction variable corresponding to a representative condition is stored in association with a set of data collected by the X-ray projection data collection unit; ,
A process including an image reconstruction unit including an image reconstruction process of the X-ray projection data based on an image reconstruction variable corresponding to an imaging condition used for the X-ray data collection including at least a condition other than the representative condition An X-ray CT apparatus comprising: an image reconstruction unit that performs the operation.   The image reconstruction unit is configured to perform image reconstruction of the X-ray projection data based on an image reconstruction variable corresponding to an imaging condition used for the X-ray data collection including the representative condition and a condition other than the representative condition. The X-ray CT apparatus according to claim 1, wherein a process including an image reconstruction unit including a configuration process is performed. The predetermined imaging condition is an X-ray tube voltage condition,
The X-ray data collection unit collects X-ray projection data using a plurality of X-ray tube voltage conditions,
The X-ray CT apparatus according to claim 1, wherein the representative condition is a plurality of X-ray tube voltage conditions used for the X-ray data collection. The predetermined imaging condition is an X-ray tube voltage condition,
The X-ray data collection unit collects X-ray projection data using a plurality of X-ray tube voltage conditions,
The representative condition is a plurality of X-ray tube voltage conditions used for the X-ray data collection,
The X-ray CT apparatus according to claim 2, wherein the condition other than the representative condition includes a condition of a transient section accompanying switching of the plurality of X-ray tube voltage conditions.   The image reconstruction unit calculates or selects an image reconstruction variable corresponding to the representative condition, obtains an image reconstruction variable corresponding to a condition other than the representative condition, and determines the image reconstruction variable. 5. The X-ray CT apparatus according to claim 1, wherein processing using an image reconstruction variable corresponding to a condition other than the representative condition is performed.   The X-ray CT apparatus according to claim 5, wherein the image reconstruction unit adds the obtained image reconstruction variable to the data set and causes the storage unit to store the variable.   The image reconstruction unit calculates or selects X-ray projection data obtained by processing using an image reconstruction variable corresponding to the representative condition, and corresponds to a condition other than the representative condition. 2. The process using an image reconstruction variable corresponding to a condition other than the representative condition is performed by obtaining X-ray projection data obtained by a process using a reconstruction variable. The X-ray CT apparatus according to any one of 4.   The X-ray CT apparatus according to claim 5, wherein the storage unit stores a calculation condition for the calculation corresponding to a condition other than the representative condition. The image reconstruction unit stores the range of the imaging condition and an image reconstruction variable corresponding to the range of the imaging condition, and performs image reconstruction corresponding to a condition other than the representative condition included in the imaging condition range. The X-ray CT apparatus according to claim 1, wherein the configuration variable is obtained from an image reconstruction variable corresponding to the stored range.   The X-ray CT apparatus according to claim 1, wherein the image reconstruction variable is correction data used for preprocessing of image reconstruction.   The X-ray CT apparatus according to claim 1, wherein the imaging conditions are selected from an X-ray tube voltage, an X-ray focal position, a gantry rotation speed, and an aperture width of a collimator.   A data file including X-ray projection data used in the X-ray CT apparatus according to claim 1, wherein the data file includes the X-ray projection data and image reconstruction variables corresponding to the plurality of imaging conditions. Data file to be used.