JP5179007B2 - X-ray CT apparatus and X-ray CT image reconstruction method thereof - Google Patents

Description

  The present invention relates to an X-ray CT (Computed Tomography) apparatus using a two-dimensional area X-ray detector and an X-ray CT image reconstruction method thereof.

  Conventional X-ray CT apparatuses using a multi-row X-ray detector X-ray CT apparatus or a two-dimensional X-ray area detector typified by a flat panel planar X-ray detector are shown in FIGS. 15 (a) and 15 (b). ), As shown in FIG. 15C, data is collected by arranging X-ray detector channels of equal size in both the channel direction and the column direction (see, for example, Patent Document 1). For example, in the case of the column width d, the X-ray detector width in the column direction is 8d in the X-ray detector configuration of 8 columns × d, and in the case of the column width 2d, the portion of the column width d of the center 8 columns is Two rows are bundled into one row to form a 4 column × 2d X-ray detector configuration, and the entire X-ray detector has an 8 column × 2d X-ray detector configuration. In this way, the following two X-ray detector configurations were switched and used.

1． X-ray detector mode 1: 8 rows x d X-ray detector row width 8d
2． X-ray detector mode 2: 8 rows x 2d X-ray detector row width 16d
One reason is that the number of channels and columns that the data collection device DAS can handle are limited.

  In this way, when both the X-ray detector with the width d in the column direction and the X-ray detector with the width 2d in the column direction depend on each other as shown in FIG. 16, two X-ray detectors with the width d are bundled together. This is a problem from the viewpoint that the spatial resolution capability of the X-ray detector having the width d cannot be fully utilized because the data is collected with the width 2d in the column direction.

For example, in the X-ray detector shown in FIG. 16, the following X-ray detector mode is desired.
3． X-ray detector mode 3: 8 rows x d + 4 rows x 2d X-ray detector row width 16d
However, there are problems with restrictions on the number of channels and columns that can be handled by the data acquisition device DAS, and X-ray detector apertures with a width of d and 2d are mixed in the column direction, and image reconstruction cannot be performed in image reconstruction. There was a problem.
Japanese Patent Laid-Open No. 2004-073360

  However, in the X-ray CT apparatus using a multi-row X-ray detector X-ray CT apparatus or a two-dimensional X-ray area detector represented by a flat panel planar X-ray detector, the width in the column direction of the X-ray detector is It becomes smaller and almost equal to the width in the channel direction, and isotropic resolution is required as a tomographic image or a tomographic image continuous in the z direction, that is, isotropicity between the column direction width of the X-ray detector and the channel direction width. Is coming.

  Therefore, an object of the present invention is to use a multi-row X-ray detector or a two-dimensional X-ray area detector having a matrix structure represented by a flat panel planar X-ray detector. X-ray CT system or X-ray CT image reconstruction that realizes tomographic images with sufficient image quality by correcting the difference in detector channel aperture even if a line detector channel exists It is to provide a method.

  If the data acquisition device (DAS) can handle a sufficient number of channels and columns, or the data acquisition device (DAS) that can switch each channel of the X-ray detector, the channels of the X-ray detector that can be handled As long as there are no restrictions on the number and the number of columns, it is only necessary that the X-ray projection data obtained can be reconstructed even if the aperture of each channel of the X-ray detector has a different size.

  The present invention provides a plurality of X-ray projection data passing through a pixel having a tomographic image when reconstructing projection data in a three-dimensional image, and each projection data has different X-ray detector aperture areas. By applying a weighting factor proportional to the area ratio of the X-ray detector aperture to the X-ray projection data and performing weighted addition to reconstruct the 3D image, the multi-row X-ray detector or flat panel X-ray detector can be used. In a two-dimensional X-ray area detector with a representative matrix structure, conventional scan (axial scan), cine scan, helical scan, or variable pitch helical scan of an X-ray CT apparatus having X-ray detector channels with different aperture areas Provided is an X-ray CT apparatus or an X-ray CT image reconstruction method characterized in that X-ray projection data can be reconstructed.

  In a first aspect, the present invention relates to an X-ray generator and a two-dimensional X-ray area detector having a matrix structure that detects X-rays relative to each other while rotating about a rotation center between them. , X-ray data collection means for collecting X-ray projection data transmitted through the subject in between, image reconstruction means for reconstructing the projection data collected from the X-ray data collection means, and image-reconstructed tomogram In an X-ray CT apparatus comprising an image display means for displaying an image, when reconstructing projection data acquired by X-ray data, there is a plurality of projection data passing through pixels with a tomographic image, Image reconstruction means for reconstructing an image by weighting and adding a weighting factor depending on the area ratio of the X-ray detector aperture to each projection data when the X-ray detector aperture area of the projection data is different X-ray CT system characterized by having Subjected to.

  In the X-ray CT apparatus according to the first aspect described above, three-dimensional image reconstruction is performed by correcting the projection data having different X-ray detector aperture areas by weighted addition with a weighting factor that depends on the X-ray detector aperture area ratio. Therefore, image reconstruction can be performed without contradiction, and a tomographic image with good image quality can be obtained.

  In a second aspect, the present invention relates to an X-ray generator and a two-dimensional X-ray area detector having a matrix structure that detects X-rays relative to each other while rotating around the center of rotation. , X-ray data collection means for collecting X-ray projection data transmitted through the subject in between, image reconstruction means for reconstructing the projection data collected from the X-ray data collection means, and image-reconstructed tomogram In an X-ray CT apparatus comprising an image display means for displaying an image, when reconstructing projection data acquired by X-ray data, there is a plurality of projection data passing through pixels with a tomographic image, When the area of the X-ray detector aperture of the projection data is different, the effective aperture area of the X-ray detector aperture is expanded after superimposing a filter in the z-axis direction on each projection data. Weighting factor depending on the ratio of the X-ray detector aperture area And weighted addition over to provide an X-ray CT apparatus characterized by an image reconstruction means for the image reconstruction.

  In the X-ray CT apparatus according to the second aspect, a filter in the z-axis direction is superimposed on projection data having different X-ray detector aperture areas to expand the effective X-ray detector aperture area of the X-ray detector aperture. After that, since the three-dimensional image is reconstructed by performing weighted addition with a weighting factor depending on the area ratio of the effective aperture area of the X-ray detector, the image reconstruction can be performed with good image quality.

  In a third aspect, the present invention relates to an X-ray generator and a two-dimensional X-ray area detector having a matrix structure that detects X-rays relative to each other while rotating around the center of rotation. , X-ray data collection means for collecting X-ray projection data transmitted through the subject in between, image reconstruction means for reconstructing the projection data collected from the X-ray data collection means, and image-reconstructed tomogram In an X-ray CT apparatus comprising an image display means for displaying an image, when reconstructing projection data acquired by X-ray data, there is a plurality of projection data passing through pixels with a tomographic image, When the X-ray detector aperture areas of the projection data are different, the z-axis direction filter is superimposed on each projection data, and the effective aperture area of the X-ray detector apertures is made approximately the same to reconstruct the image It has an image reconstruction means to perform An X-ray CT system is provided.

  In the X-ray CT apparatus according to the third aspect, the filter in the z-axis direction is superimposed on projection data with different areas of the X-ray detector aperture, and the effective aperture area of the X-ray detector aperture is substantially equalized. Since the three-dimensional image reconstruction is performed, the image reconstruction can be performed with good image quality.

  In a fourth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to third aspects, wherein the two-dimensional X-ray area detector is a multi-row X-ray detector typified by an arc shape. There is provided an X-ray CT apparatus characterized by having an X-ray data collection means and an image reconstruction means for performing a three-dimensional image reconstruction.

In the X-ray CT apparatus according to the fourth aspect, image reconstruction can be performed with good image quality even when an arc-shaped multi-row X-ray detector is used as the two-dimensional X-ray area detector.
In a fifth aspect, the present invention provides an X-ray CT apparatus according to any one of the first to third aspects, wherein the two-dimensional X-ray area detector is a planar type represented by a flat panel X-ray detector. X-ray data collection means that is a two-dimensional X-ray area detector, or a two-dimensional X-ray area detector in combination with a planar two-dimensional X-ray area detector, and image reconstruction means for performing three-dimensional image reconstruction An X-ray CT apparatus is provided.

  In the X-ray CT apparatus according to the fifth aspect, the two-dimensional X-ray area detector is combined with a flat-type two-dimensional X-ray area detector or a two-dimensional X-ray area detector. Similarly, an area detector can perform image reconstruction with good image quality.

  In a sixth aspect, the present invention provides an X-ray CT apparatus according to any one of the first to fifth aspects, wherein a plurality of projections passing through pixels having a tomographic image are obtained when reconstructing projection data into a three-dimensional image. If there is data and the area of the X-ray detector aperture of each projection data is different, each projection data is weighted and added with a weighting factor proportional to the area ratio of the X-ray detector aperture to reconstruct the 3D image. Provided is an X-ray CT apparatus characterized by having an image reconstruction means for configuring.

  In the X-ray CT apparatus according to the sixth aspect described above, three-dimensional image reconstruction is performed by correcting the projection data having different X-ray detector aperture areas by weighted addition with a weighting factor proportional to the X-ray detector aperture area ratio. Therefore, image reconstruction can be performed with good image quality.

  In a seventh aspect, the present invention is a conventional scan (axial scan) or cine scan data collection in which the z-axis coordinate is one position in the X-ray CT apparatus according to any one of the first to sixth aspects Provided is an X-ray CT apparatus characterized by having X-ray data collection means.

  In the X-ray CT apparatus according to the seventh aspect, image reconstruction can be performed with good image quality in the same manner by conventional scanning (axial scanning) or cine scanning at one z-axis coordinate position.

  In an eighth aspect, the present invention relates to an X-ray CT apparatus according to any one of the first to sixth aspects, wherein the X-axis coordinate is data acquisition of a conventional scan (axial scan) or cine scan having a plurality of z-axis coordinates. An X-ray CT apparatus characterized by having a line data collection means is provided.

  In the X-ray CT apparatus according to the eighth aspect described above, the X-ray projection data of adjacent scans can be obtained even if the apertures are different as in the conventional scan (axial scan) or cine scan of the z-axis coordinates at a plurality of locations. A weighted addition of the aperture areas is performed, and the images are reconstructed by mixing them to obtain a tomographic image with good image quality.

  In a ninth aspect, the present invention relates to an X-ray CT apparatus according to the eighth aspect, wherein X-axis scanning of different z-axis coordinates is performed in data acquisition of a conventional scan (axial scan) or cine scan with a plurality of z-axis coordinates. The aperture area of the 2D area X-ray detector is different for the line projection data, and each X-ray projection data is weighted and added with a weighting factor that depends on the aperture area ratio of the 2D area X-ray detector, and then 3D Provided is an X-ray CT apparatus characterized by having an image reconstruction means for performing image reconstruction.

  In the X-ray CT apparatus according to the ninth aspect described above, similarly to conventional scans (axial scans) or cine scans of z-axis coordinates at a plurality of locations, projection data having different X-ray detector aperture areas can be obtained. Since the three-dimensional image is reconstructed by weighting and adding with a weighting factor proportional to the area ratio, the image can be reconstructed with good image quality.

  In a tenth aspect, the present invention provides an X-ray CT apparatus characterized by having an X-ray data collection means for collecting helical scan data in any of the first to sixth X-ray CT apparatuses. provide.

In the X-ray CT apparatus according to the tenth aspect, image reconstruction can be performed with good image quality even in helical scanning.
In an eleventh aspect, the present invention provides an X-ray CT apparatus according to any one of the first to tenth aspects, wherein X-ray data collection means having different X-ray detector aperture areas in the X-ray detector row direction are provided. An X-ray CT apparatus characterized by having it is provided.

  In the X-ray CT apparatus according to the eleventh aspect, even if the areas of the X-ray detector openings are different in the column direction, the image is reproduced after an area ratio corresponding to the area of each X-ray detector opening is applied in the column direction. By performing the configuration, image reconstruction can be performed without contradiction, and a tomographic image with good S / N and resolution can be obtained.

  In a twelfth aspect, the present invention provides an X-ray CT apparatus according to any one of the first to tenth aspects, wherein X-ray data collection means having different X-ray detector aperture areas in the X-ray detector channel direction is provided. An X-ray CT apparatus characterized by having it is provided.

  In the X-ray CT apparatus according to the twelfth aspect, even if the area of the X-ray detector opening is different in the channel direction, image reconstruction is performed after applying an area ratio corresponding to the area of each X-ray detector opening in the channel direction. By performing the configuration, image reconstruction can be performed without contradiction, and a tomographic image with good S / N and resolution can be obtained.

  According to the X-ray CT apparatus or X-ray CT image reconstruction method of the present invention, a multi-row X-ray detector or a two-dimensional X-ray area detector having a matrix structure typified by a flat panel planar X-ray detector , Even if there are X-ray detector channels with different X-ray detector channel apertures, the difference in detector channel aperture is corrected and image reconstruction is performed to obtain a tomographic image with sufficient image quality. An X-ray CT apparatus or an X-ray CT image reconstruction method can be realized.

Hereinafter, the present invention will be described in more detail with reference to embodiments shown in the drawings. Note that the present invention is not limited thereby.
FIG. 1 is a configuration block diagram of an X-ray CT apparatus 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 collects X-ray detector data collected by the scanning device gantry 20 and an input device 2 that receives input from the operator, a central processing device 3 that performs preprocessing, image reconstruction processing, post-processing, and the like. Data acquisition buffer 5, monitor 6 that displays tomograms reconstructed from projection data obtained by preprocessing X-ray detector data, program, X-ray detector data, projection data, and X-ray tomogram And a storage device 7 for storing.

The photographing condition is input from the input device 2 and stored in the storage device 7. FIG. 14 shows an example of the shooting condition input screen.
The imaging table 10 includes a cradle 12 on which a subject is placed and taken in and out of the opening of the scanning gantry 20. The cradle 12 is moved up and down and linearly moved by the motor built in the imaging table 10.

  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, a DAS (Data Acquisition System) 25, A rotation unit controller 26 that controls the X-ray tube 21 rotating around the body axis, and a control controller 29 that exchanges control signals and the like with the operation console 1 and the imaging table 10 are provided. 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, making it possible to absorb more X-rays. It is an X-ray filter. For this reason, exposure of the body surface of the subject having a cross-sectional shape close to a circle or an ellipse can be reduced. The scanning gantry tilt controller 27 can tilt the scanning gantry 20 forward and backward in the z direction by about ± 30 degrees.

  The X-ray tube 21 and the multi-row X-ray detector 24 rotate around the rotation center IC. When the vertical direction is the y direction, the horizontal direction is the x direction, and the table and cradle traveling direction perpendicular to these are the z direction, the rotation plane of the X-ray tube 21 and the multi-row X-ray detector 24 is the xy plane. is there. The moving direction of the cradle 12 is the z direction.

2 and 3 are explanatory views of the geometric arrangement of the X-ray tube 21 and the multi-row X-ray detector 24 as viewed from the xy plane or the yz plane.
The X-ray tube 21 generates an X-ray beam called a cone beam CB. When the direction of the central axis of the cone beam CB is parallel to the y direction, the view angle is 0 degree.

The multi-row X-ray detector 24 has, for example, 256 X-ray detector rows in the z direction. Each X-ray detector array has, for example, 1024 X-ray detector channels in the channel direction.
In FIG. 2, the X-ray beam emitted from the X-ray focal point of the X-ray tube 21 is irradiated by the beam forming X-ray filter 28 so that more X-rays are generated at the center of the reconstruction area P and more at the periphery of the reconstruction area P. After the X-ray dose is spatially controlled so that a small amount of X-rays are irradiated, the X-rays are absorbed by the subject existing inside the reconstruction area P, and the transmitted X-rays are detected in the multi-row X-ray detector 24. Collected as X-ray detector data.

  In FIG. 3, the X-ray beam emitted from the X-ray focal point of the X-ray tube 21 is controlled in the slice thickness direction of the tomogram by the X-ray collimator 23, that is, the X-ray beam width becomes D at the rotation center axis IC. Thus, X-rays are absorbed by the subject existing in the vicinity of the rotation center axis IC, and the transmitted X-rays are collected as X-ray detector data by the multi-row X-ray detector 24.

  Projection data collected by irradiation with X-rays is A / D converted from the multi-row X-ray detector 24 by the DAS 25 and input to the data collection buffer 5 via the slip ring 30. The data input to the data collection buffer 5 is processed by the central processing unit 3 according to the program in the storage device 7, reconstructed into a tomographic image, and displayed on the monitor 6.

FIG. 4 is a flowchart showing an outline of the operation of the X-ray CT apparatus of the present embodiment.
In step P1, the subject is placed on the cradle 12 and aligned. The subject placed on the cradle 12 aligns the center position of the slice light of the scanning gantry 20 with the reference point of each part.

  In step P2, scout image collection is performed. Scout images are usually taken at 0 and 90 degrees, but depending on the part, for example, the head may be a 90-degree scout image only. Details of scout image shooting will be described later.

  In step P3, shooting conditions are set. The normal photographing condition is to perform photographing while displaying the position and size of the tomographic image to be photographed on the scout image. In this case, the entire X-ray dose information for one helical scan, variable pitch helical scan, helical shuttle scan, conventional scan (axial scan) or cine scan is displayed. In the cine scan, when the number of rotations or time is entered, X-ray dose information for the input number of rotations or the input time in the region of interest is displayed.

In step P4, tomographic imaging is performed. Details of tomographic imaging will be described later.
FIG. 5 is a flowchart showing an outline of tomographic and scout image capturing operations of the X-ray CT apparatus 100 of the present invention.

  In step S1, in the helical scan, the X-ray detector data is obtained by rotating the X-ray tube 21 and the multi-row X-ray detector 24 around the subject and moving the cradle 12 on the imaging table 10 linearly on the table. The X-ray detector data D0 (view, j, i) represented by the view angle view, detector row number j, and channel number i is moved to the table linear movement z direction position Ztable (view) To collect X-ray detector data. In a helical scan, a certain range of data is collected. In the variable pitch helical scan or helical shuttle scan, in addition to collecting data in a constant speed range, data collection is also performed during acceleration and deceleration.

  In conventional scan (axial scan) or cine scan, X-ray detector data is collected by rotating the data acquisition system one or more times while the cradle 12 on the imaging table 10 is fixed at a certain z-direction position. Do. If necessary, after moving to the next position in the z direction, the data acquisition system is rotated once or more times to collect data of X-ray detector data.

  In scout imaging, the X-ray tube 21 and the multi-row X-ray detector 24 are fixed, and the data collection operation of the X-ray detector data is performed while the cradle 12 on the imaging table 10 is moved linearly. .

  In step S2, the X-ray detector data D0 (view, j, i) is preprocessed and converted into projection data. As shown in FIG. 6, the preprocessing includes step S21 offset correction, step S22 logarithmic conversion, step S23 X-ray dose correction, and step S24 sensitivity correction.

  In the case of scout image capture, the preprocessed X-ray detector data can be displayed as a scout image by displaying the pixel size in the channel direction and the pixel size in the z direction, which is the cradle linear movement direction, in accordance with the display pixel size of the monitor 6. Completion.

  In step S3, beam hardening correction is performed on the preprocessed projection data D1 (view, j, i). In step S3 beam hardening correction, the projection data subjected to the sensitivity correction S24 in the pre-processing S2 is D1 (view, j, i), and the data after step S3 beam hardening correction is D11 (view, j, i). Then, the step S3 beam hardening correction is expressed, for example, in a polynomial form as shown in the following (Formula 1).

  At this time, since independent beam hardening correction can be performed for each j column of the detector, if the tube voltage of each data acquisition system differs depending on the imaging conditions, the X-ray energy characteristics of the detector for each column Differences can be corrected.

In step S4, z filter convolution processing for applying a filter in the z direction (column direction) to the projection data D11 (view, j, i) subjected to beam hardening correction is performed.
That is, the projection of the multi-row X-ray detector D11 (view, j, i) (i = 1 to CH, j = 1 to ROW) subjected to beam hardening correction after preprocessing in each view angle and each data acquisition system For example, a filter with a column direction filter size of 5 columns as shown in (Formula 2) and (Formula 3) below is applied to the data in the column direction.

  The corrected detector data D12 (view, j, i) is as shown in (Formula 4) below.

It becomes. If the maximum value of the channel is CH and the maximum value of the column is ROW,
The following (Formula 5) and (Formula 6) are obtained.

  Further, when the column direction filter coefficient is changed for each channel, the slice thickness can be controlled in accordance with the distance from the image reconstruction center. In general, in slice images, the slice thickness is thicker in the periphery than in the reconstruction center, so the column direction filter coefficient is changed between the center and the periphery, and the column direction filter coefficient is changed in the column direction near the center channel. If the width of the filter coefficient is changed widely, the slice thickness can be made uniform in the peripheral part and the image reconstruction center part by changing the width of the column direction filter coefficient in the vicinity of the peripheral channel.

  In this way, by controlling the column direction filter coefficients of the central channel and the peripheral channel of the multi-row X-ray detector 24, the slice thickness can also be controlled at the central portion and the peripheral portion. When the slice thickness is slightly reduced with the row direction filter, both artifacts and noise are greatly improved. Thereby, artifact improvement and noise improvement can also be controlled. That is, it is possible to control the tomographic image reconstructed, that is, the image quality in the xy plane. As another embodiment, a thin slice thickness tomogram can be realized by using a deconvolution filter with column direction (z direction) filter coefficients.

  In step S5, reconstruction function superimposition processing is performed. That is, the Fourier transform is performed, the reconstruction function is multiplied, and the inverse Fourier transform is performed. In reconstruction function superimposition processing S5, assuming that the data after z filter convolution processing is D12, the data after reconstruction function convolution processing is D13, and the reconstruction function to be superimposed is Kernel (j), the reconstruction function convolution processing is as follows: (Equation 7).

In other words, since the reconstruction function kernel (j) can perform the reconstruction function superimposing process independently for each j column of the detector, the difference in noise characteristics and resolution characteristics for each column can be corrected.
In step S6, three-dimensional backprojection processing is performed on the projection data D13 (view, j, i) subjected to reconstruction function superimposition processing to obtain backprojection data D3 (x, y, z). The image to be reconstructed is a three-dimensional image reconstructed on the plane perpendicular to the z axis and the xy plane. The following reconstruction area P is assumed to be parallel to the xy plane. This three-dimensional backprojection process will be described later with reference to FIG.

In step S7, 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 D31 (x, y).
In post-processing image filter superimposition processing, the tomographic image after three-dimensional backprojection is D31 (x, y, z), the data after image filter superimposition is D32 (x, y, z), and the tomographic image plane xy If the two-dimensional image filter superimposed on the plane is Filter (z), the following (Formula 8) is obtained.

That is, since independent image filter convolution processing can be performed for each j column of the detector, the difference in noise characteristics and resolution characteristics for each column can be corrected.
Alternatively, the image filter Filter (z) may be a three-dimensional image filter.

Alternatively, the image space z-direction filter convolution process may be performed after the two-dimensional image filter convolution.
The obtained tomographic image is displayed on the monitor 6.

FIG. 7 is a flowchart showing details of the three-dimensional backprojection process (step S6 in FIG. 5).
In this embodiment, the image to be reconstructed is reconstructed into a three-dimensional image on a plane perpendicular to the z axis and on the xy plane. The following reconstruction area P is assumed to be parallel to the xy plane.

  In step S61, attention is paid to one view in all views necessary for image reconstruction of the tomogram (that is, a view of 360 degrees or a view of “180 degrees + fan angle”). Projection data Dr corresponding to each pixel is extracted.

  As shown in FIGS. 8 (a) and 8 (b), a 512 × 512 pixel square region parallel to the xy plane is used as a reconstruction region P, and pixel rows L0, y parallel to the x axis where y = 0 = 63 pixel column L63, y = 127 pixel column L127, y = 191 pixel column L191, y = 255 pixel column L255, y = 319 pixel column L319, y = 383 pixel column L383, y = 447 9 is projected onto the surface of the multi-row X-ray detector 24 in the X-ray transmission direction as shown in FIG. If projection data on .about.T511 are extracted, they become projection data Dr (view, x, y) of the pixel columns L0 to L511. However, x and y correspond to each pixel (x, y) of the tomographic image.

  The X-ray transmission direction is determined by the X-ray focal point of the X-ray tube 21 and the geometric position of each pixel and the multi-row X-ray detector 24, but z of the X-ray detector data D0 (view, j, i). Since the coordinate z (view) is attached to the X-ray detector data as the table linear movement z-direction position Ztable (view), the X-ray detector data D0 (view, j, i) during acceleration / deceleration is also known. The X-ray transmission direction can be accurately determined in the data acquisition geometric system of the X-ray focus and multi-row X-ray detector.

  For example, a part of the line goes out of the channel direction of the multi-row X-ray detector 24, such as a line T0 in which the pixel row L0 is projected on the surface of the multi-row X-ray detector 24 in the X-ray transmission direction. In this case, the corresponding projection data Dr (view, x, y) is set to “0”. Further, if the projection is out of the z direction, the projection data Dr (view, x, y) is extrapolated.

In this way, as shown in FIG. 10, projection data Dr (view, x, y) corresponding to each pixel of the reconstruction area P can be extracted.
Returning to FIG. 7, in step S62, the projection data Dr (view, x, y) is multiplied by the cone beam reconstruction weighting coefficient to create projection data D2 (view, x, y) as shown in FIG.

  Here, the cone beam reconstruction weighting coefficient w (i, j) is as follows. In the case of fan beam image reconstruction, in general, when view = βa, a straight line connecting the focal point of the X-ray tube 21 and the pixel g (x, y) on the reconstruction area P (on the xy plane) is the center of the X-ray beam. When the angle formed with respect to the axis Bc is γ and the opposite view is view = βb, the following (Formula 9) is obtained.

  If the angles between the X-ray beam passing through the pixel g (x, y) on the reconstruction area P and the opposite X-ray beam and the reconstruction plane P are αa and αb, the cone beam reconstruction weighting coefficient depending on them Multiply and multiply by ωa and ωb to obtain backprojection pixel data D2 (0, x, y). In this case, (Formula 10) is obtained.

  Note that the sum of the cone beam reconstruction weighting coefficients between the opposed beams is expressed by (Formula 11).

Cone angle artifacts can be reduced by multiplying and adding cone beam reconstruction weighting coefficients ωa and ωb.
For example, the cone beam reconstruction weighting coefficients ωa and ωb can be obtained by the following equations. Note that ga is a weighting coefficient for the view βa, and gb is a weighting coefficient for the view βb.

  When 1/2 of the fan beam angle is γmax, the following (Expression 12) to (Expression 17) are obtained.

(For example, q = 1)
For example, as an example of ga and gb, when max [] is a function that takes the larger value, the following (Formula 18) and (Formula 19) are obtained.

In the case of fan beam image reconstruction, each pixel on the reconstruction area P is further multiplied by a distance coefficient. For the distance coefficient, the distance from the focus of the X-ray tube 21 to the detector row j and channel i of the multi-row X-ray detector 24 corresponding to the projection data Dr is r0, and the distance from the focus of the X-ray tube 21 corresponds to the projection data Dr. When the distance to the pixel on the reconstruction area P to be set is r1, (r1 / r0) ² .

In the case of parallel beam image reconstruction, each pixel on the reconstruction area P may be multiplied by only the cone beam reconstruction weight coefficient w (i, j).
Further, S ₁ is the X-ray detector aperture of the plurality of projection data passing through a certain pixel of the tomographic image, S _2, ... If it is S _n, the projection data P ₁ to be backprojected to the pixel, P ₂ ... P _n Is backprojected as follows: If the backprojected data is D ₃ (x, y), the result is as follows.

  Even if the above correction is not performed, especially when the X-ray detector opening width is different in the z direction as in Examples 1, 2, and 3 to be described later, that is, the X-ray detector row width in the z direction. If they are different, the z-filter superimposing process in step S4 can change the z-filter superimposed on each column to make the widths of the respective columns in the z direction uniform. For example, if each detector row width and row position in the original z direction is as shown in FIG. 20, the slice thickness of the 8A-1A row and the 1B-8A row is increased in the z direction by z filter convolution processing. Thus, it can be as shown in FIG. Further, as shown in FIG. 22, the slice thickness can be made uniform by reducing the slice thickness by using the z filter coefficients of the 12A to 9A rows and the 12B to 9B rows as deconvolution filters.

In this manner, the X-ray detector aperture width in the z direction may be controlled using the z filter convolution process.
In step S63, as shown in FIG. 12, the projection data D2 (view, x, y) is added in correspondence with the pixels to the backprojection data D3 (x, y) that has been cleared in advance.

  In step S64, steps S61 to S63 are repeated for all views necessary for image reconstruction of tomographic images (ie, views for 360 degrees or views for “180 degrees + fan angle”), as shown in FIG. Then, back projection data D3 (x, y) is obtained.

As shown in FIGS. 13 (a) and 13 (b), the reconstruction area P may not be a square area of 512 × 512 pixels, but a circular area having a diameter of 512 pixels.
As described above, in the two-dimensional X-ray area detector having a matrix structure typified by a multi-row X-ray detector or a flat panel planar X-ray detector, the X-ray detector channel openings are different from each other. An example has been described in which image reconstruction is performed by correcting the X-ray detector channel aperture even if a line detector channel exists.

In the following embodiment, an example will be described in which image reconstruction is performed by correcting the X-ray detector channel aperture even under different imaging conditions under the following imaging conditions.
(Example 1) When the aperture width of the X-ray detector differs in the z direction in a conventional scan (axial scan) or cine scan with one z-direction coordinate.

(Example 2) When the X-ray detector aperture width is different in the z direction in a conventional scan (axial scan) or cine scan with two or more z-direction coordinates.
(Example 3) Helical scan when the X-ray detector aperture width differs in the z direction.

  (Example 4) Conventional scan (axial scan) or cine scan and helical scan when the X-ray detector channel opening size differs in the channel direction.

  In the first embodiment, image reconstruction will be described in the case where the X-ray detector aperture widths in the z direction (column direction) in one conventional scan (axial scan) or cine scan in which the z direction coordinate is z = z0 are different.

  FIG. 17 shows a conventional scan (axial scan) or a cine scan at a certain z-direction coordinate z = z0. At this time, the multi-row X-ray detector 24 has 2n rows (16 rows) of X-ray detector rows of width d in the center and n / 2 rows (8 rows) of X-ray detector rows of width 2d on the outside. There is one by one.

  At this time, in the pixel G (x1, y1) on the tomographic image plane (xy plane), the X-ray beam in the 0-degree direction has an X-ray detector channel aperture in the column direction of 2d, and the X-ray beam in the 180-degree direction The X-ray detector channel aperture in the column direction is d.

  To perform this image reconstruction, the X-ray detector aperture in the column direction is superimposed in the z direction (column direction) inverse weight (deconvolution) filter on the 2d column before the 3D back projection. Perform 3D backprojection with narrowing. The slice thickness of the tomographic image obtained in this way can be a tomographic image having a slice thickness slightly thicker than d, for example, a slice thickness of about 1.5d.

  Alternatively, both the x-ray detector channel aperture in the column direction and the x-ray detector row with 2d in the column direction and the x-ray detector column in the column direction x-ray detector channel aperture with the z-direction (column direction) filter and z Direction (column direction) without superimposing (deconvolution) filters, the aperture size ratio r, that is, in this case, the ratio r = 2d: d = 2, or the size ratio of this aperture By adding the ratio of the function f (r) depending on r to the 2d X-ray detector array, a tomogram having a slice thickness of approximately 2d can be obtained.

  In the second embodiment, images in the case where the X-ray detector aperture width in the z direction (column direction) in the conventional scan (axial scan) or the cine scan in the z-direction coordinate z = z0 or z1 is different. The reconstruction will be described.

FIG. 18 shows a conventional scan (axial scan) or a cine scan at two z-direction coordinate positions z = z0 or z1.
At this time, in the multi-row X-ray detector 24, there are 2n rows (16 rows) of X-ray detector rows having a width d in the center, and n / 2 rows (8 rows) of X-ray detector rows having a width 2d on the outside. ) One by one.

  At this time, in the pixel G (x2, y2) on the tomographic image plane (xy plane), the X-ray beam of z = z0 in the 180 degree direction is d in the column direction X-ray detector channel aperture, and z = z1 The X-ray beam in the 0 degree direction of the X-ray detector channel opening in the column direction is 2d. In this way, even between X-ray projection data with different z-direction coordinate positions, image reconstruction can be performed by correcting the difference in aperture of the X-ray detector channel.

  To perform this image reconstruction, the X-ray detector aperture in the column direction is superimposed in the z direction (column direction) inverse weight (deconvolution) filter on the 2d column before the 3D back projection. Perform 3D backprojection with narrowing. The slice thickness of the tomographic image obtained in this way can be a tomographic image having a slice thickness slightly thicker than d, for example, a slice thickness of about 1.5d.

  Alternatively, both the x-ray detector channel aperture in the column direction and the x-ray detector row with 2d in the column direction and the x-ray detector column in the column direction x-ray detector channel aperture with the z-direction (column direction) filter and z Direction (column direction) without superimposing (deconvolution) filters, the aperture size ratio r, that is, in this case, the ratio r = 2d: d = 2, or the size ratio of this aperture By adding the ratio of the function f (r) depending on r to the 2d X-ray detector array, a tomographic image having a slice thickness of approximately 2d can be obtained.

  In the third embodiment, the X-ray detector aperture width in the z direction (column direction) is different between the two views where the z-direction coordinate is z = z0 or z1 and the opposite view showing the case of helical scanning. The image reconstruction in this case will be described.

In FIG. 19, the X-ray detector aperture width in the z direction (column direction) differs between a certain view and its opposite view at the z-direction coordinate position z = z0 or z1.
The X-ray beam in the 0-degree direction with z = z0 is d in the X-ray detector channel aperture in the column direction, and the X-ray beam in the column direction with z = z1 is d in the X-ray detector channel aperture in the column direction. 2d. In this way, it is possible to reconstruct an image by correcting the difference in the aperture of the X-ray detector channel even between X-ray projection data of views with different z-direction coordinate positions in the X-ray projection data of one rotation of the helical scan. .

  To perform this image reconstruction, the X-ray detector aperture in the column direction is superimposed in the z direction (column direction) inverse weight (deconvolution) filter on the 2d column before the 3D back projection. Perform 3D backprojection with narrowing. The slice thickness of the tomographic image obtained in this way can be a tomographic image having a slice thickness slightly thicker than d, for example, a slice thickness of about 1.5d.

  Alternatively, both the x-ray detector channel aperture in the column direction and the x-ray detector row with 2d in the column direction and the x-ray detector column in the column direction x-ray detector channel aperture with the z-direction (column direction) filter and z Direction (column direction) without superimposing (deconvolution) filters, the aperture size ratio r, that is, in this case, the ratio r = 2d: d = 2, or the size ratio of this aperture By adding the ratio of the function f (r) depending on r to the 2d X-ray detector array, a tomographic image having a slice thickness of approximately 2d can be obtained.

In the fourth embodiment, image reconstruction in the case where X-ray detector channel aperture widths differ in the channel direction in X-ray projection data of a certain view will be described.
An example is shown in FIGS. The opening width of the X-ray detector in FIGS. 23 and 24 differs in the channel direction.

However, in this case, image reconstruction can be performed by the following method.
Example 4-1: A channel whose aperture width is n times (na) in the channel direction is divided into n width a channel data having the same value, and X-ray projection data is corrected to reconstruct an image. (However, the normal channel X-ray detector channel aperture width is a)
Example 4-2: A channel whose aperture width is n times (na) in the channel direction is interpolated and divided into n pieces of data of width a, and X-ray projection data is corrected to reconstruct an image.

  Example 4-3: Projection data having different opening widths when back projection is performed by dividing projection data for each data having the same opening width in the channel direction, superimposing a reconstruction function separately for each projection data having different opening widths. All the back projections of are added to cause image reconstruction.

  FIG. 25 shows X-ray projection data having different X-ray detector channel opening widths for each channel position. The horizontal axis indicates the channel direction of the X-ray detector data or X-ray projection data of the multi-row X-ray detector 24, and the vertical axis indicates the view direction of the X-ray detector data or projection data.

  Channel opening width of X-ray detector data from channel 1 to C1-1 channel is a2, channel opening width of X-ray detector data from channel C1 to C2-1 channel is a1, channel C2 to C3-1 X-ray detector data channel aperture width is a, C3 channel to C4-1 channel channel aperture width is a1, C4 channel to N channel X-ray detector data channel aperture width Collects X-ray projection data at a2. However, the maximum channel is N channels. The channel opening width is a ≦ a1 ≦ a2, and the channel number is 1 ≦ C1 ≦ C2 ≦ C3 ≦ C4 ≦ N. For example, a1 is a channel opening width twice as large as a, and a2 is a channel opening width three times as large as a.

FIG. 30 shows the flow of image reconstruction processing in Example 4-1.
Data collection in step S11 performs the same processing as in step S1 in FIG.
The pre-processing in step S12 is the same as that in step S2 in FIG.

The beam hardening correction in step S13 performs the same process as in step S3 in FIG.
The z filter convolution process in step S14 performs the same process as in step S4 in FIG.
The reconstruction function superimposing process in step S16 performs the same process as in step S5 in FIG.

The three-dimensional backprojection process in step S17 performs the same process as in step S6 in FIG.
The post-process at step S18 is the same as step S7 in FIG.
In the channel direction data correction in step S15, the original X-ray projection data in FIG. 26 is corrected as shown in FIG. 26 and 27, the horizontal axis is the channel direction, and the vertical axis is the direction of the X-ray projection data value. In the original X-ray projection data of FIG. 26, the channel aperture width of the X-ray detector data from the C2 channel to the C3-1 channel is a, from the C2-1 channel to the C1 channel, and from the C3 channel to the C4-1 channel. X-ray detector data channel aperture width is a1 (for example, twice the aperture width of a), C1-1 channel to 1 channel, C4-1 channel to N channel X-ray detector data channel aperture width Is a2 (for example, an opening width three times larger than a).

  On the other hand, in the X-ray projection data corrected according to Example 4-1 in FIG. 27, the X-ray detector data from the C2 channel to the C3-1 channel remains unchanged, but from the C2-1 channel. The X-ray detector data from the C3 channel to the C4-1 channel up to the C1 channel is interpolated as follows. If the interpolated j-row i-channel X-ray detector data is d1 (i, j), the original X-ray detector data from the C2 channel to the C3-1 channel or from the C3 channel to the C4-1 channel is X-ray projection data exists only every 2 channels at intervals of 2a. Assuming that this X-ray projection data is d0 (2k, j), zero-order interpolation is performed as in the following (Expression 21) and (Expression 22).

  Similarly, in the X-ray detector data from channel 1 to C1-1 channel, C4 channel to N channel, the original X-ray detector data is X-ray detector data only every 3 channels at intervals of 3a. not exist. Assuming that this X-ray projection data is d0 (3k, j), zero-order interpolation is performed as in the following (Equation 23), (Equation 24), and (Equation 25).

In this manner, the channel direction data correction makes it possible to make the data interval in the channel direction constant in all channels.
Thereafter, by performing the reconstruction function convolution process in step S16, the three-dimensional backprojection process in step S17, and the post-process in step S18, image reconstruction can be performed even when the X-ray detector aperture is different in the channel direction.

FIG. 30 shows the flow of image reconstruction processing in the embodiment 4-2.
Data collection in step S11 performs the same processing as in step S1 in FIG.
The pre-processing in step S12 is the same as that in step S2 in FIG.

The beam hardening correction in step S13 performs the same process as in step S3 in FIG.
The z filter convolution process in step S14 performs the same process as in step S4 in FIG.
The reconstruction function superimposing process in step S16 performs the same process as in step S5 in FIG.

The three-dimensional backprojection process in step S17 performs the same process as in step S6 in FIG.
The post-process at step S18 is the same as step S7 in FIG.
In the channel direction data correction in step S15, the original X-ray projection data in FIG. 26 is corrected as shown in FIG. In FIG. 26 and FIG. 28, the horizontal axis is the channel direction, and the vertical axis is the direction of the X-ray projection data value. In the original X-ray projection data of FIG. 26, the channel aperture width of the X-ray detector data from the C2 channel to the C3-1 channel is a, from the C2-1 channel to the C1 channel, and from the C3 channel to the C4-1 channel. X-ray detector data channel aperture width is a1 (for example, twice the aperture width of a), C1-1 channel to 1 channel, C4-1 channel to N channel X-ray detector data channel aperture width Is a2 (for example, an opening width three times larger than a).

  On the other hand, in the X-ray projection data corrected by the embodiment 4-2 in FIG. 28, the X-ray detector data from the C2 channel to the C3-1 channel remains unchanged, but the C2-1 channel. The X-ray detector data from the C1 channel to the C1 channel and from the C3 channel to the C4-1 channel are interpolated as follows. If the interpolated j-row i-channel X-ray detector data is d1 (i, j), the original X-ray detector data from the C2 channel to the C3-1 channel or from the C3 channel to the C4-1 channel is X-ray projection data exists only every 2 channels at intervals of 2a. If this X-ray projection data is d0 (2k, j), linear interpolation is performed as in the following (Equation 26) and (Equation 27).

  Similarly, in the X-ray detector data from channel 1 to C1-1 channel, C4 channel to N channel, the original X-ray detector data is X-ray detector data only every 3 channels at intervals of 3a. not exist. If this X-ray projection data is d0 (3k, j), linear interpolation is performed as in the following (Equation 28), (Equation 29), and (Equation 30).

FIG. 31 shows the flow of image reconstruction processing in the embodiment 4-3.
Data collection in step S21 performs the same processing as in step S1 in FIG.
The pre-processing in step S22 performs the same processing as step S2 in FIG.

The beam hardening correction in step S23 performs the same process as in step S3 in FIG.
The z filter convolution process in step S24 performs the same process as in step S4 in FIG.
In the channel direction data division processing in step S25, the original X-ray projection data in FIG. 26 is converted to the original channel opening width of the X-ray detector from the C2 channel to the C3-1 channel as shown in FIG. X-ray projection data 1 with X-ray projection data remaining and 0 for the other part, channel aperture width of X-ray detector from C1 channel to C2-1 channel, or C3 channel to C4-1 channel is only 2a X-ray projection data 2 with the original X-ray projection data remaining and the other parts set to 0, and similarly, X-ray detector channels from 1 channel to C1-1 channel or C4 channel to N channel Only the aperture width 3a is left as the original X-ray projection data, and the other portions are divided into three X-ray projection data of X-ray projection data 3 where 0 is set.

  The reconstruction function superimposing process of step S26 is performed on the projection data of X-ray projection data 1, X-ray projection data 2, and X-ray projection data 3 divided in this way, and the reconstruction function superimposition of step S5 in FIG. 5 is performed. Repeat 3 times as in the process. However, the superposed reconstruction function superimposes the X-ray detector apertures of the X-ray projection data 1, 2, and 3, and the reconstruction functions considering a, 2a, and 3a as Nyquist frequencies. In the three-dimensional backprojection process of step S27, as shown in FIG. 32, a reconstruction function convolution process and a three-dimensional backprojection process are performed on the projection data on which these three reconstruction functions are superimposed, and image reconstruction is performed. Addition processing is performed on each configured tomographic image.

A tomogram reconstructed from projection data 1 of channel aperture width a in the channel range [C2, C3-1] is represented by G ₁ (x, y), channels [C1, C2-1], [C3, C4-1 ] A tomogram reconstructed from projection data 2 with channel opening width 2a of G ₂ (x, y), projection data with channel opening width 3a of channel range [1, C1-1], [C4, N] If the tomographic image reconstructed from ₃ is G ₃ (x, y) and the final tomographic image is G (x, y), the following equation is obtained.

After this, by performing post-processing at step S28, image reconstruction can be performed even when the X-ray detector aperture is different in the channel direction.
As described above, in any of the methods of Example 4-1, Example 4-2, and Example 4-3, image reconstruction can be performed even if the X-ray detector channel opening width is different in the channel direction.

  In the above X-ray CT apparatus 100, according to the X-ray CT apparatus or the X-ray CT imaging method of the present invention, a two-dimensional matrix structure represented by a multi-row X-ray detector or a flat panel X-ray detector In the X-ray cone beam that spreads in the z direction that existed at the beginning and end of conventional scan (axial scan) or cine scan or helical scan or variable pitch helical scan of X-ray CT apparatus with X-ray area detector, Efficiently uses X-ray projection data adjacent in the z direction to reduce exposure or reduce artifacts.

Note that the image reconstruction method may be a three-dimensional image reconstruction method by a conventionally known Feldkamp method. Furthermore, other three-dimensional image reconstruction methods may be used.
In the third embodiment, the helical scan is described, but it can be similarly applied to a variable pitch helical scan or a shuttle mode helical scan, and the same effect can be obtained.

  In addition, in this embodiment, column direction (z direction) filters with different coefficients are superimposed on each column to adjust differences in image quality due to differences in X-ray cone angle, especially in conventional scans (axial scans). In addition, a uniform slice thickness, artifact, and noise image quality are realized in each column, and various filter coefficients can be considered for this, but the same effects can be obtained in any case.

  Also, in this embodiment, column direction (z direction) filters with different coefficients are superimposed on each column to adjust the image quality variation and achieve uniform slice thickness, artifact, and noise image quality in each column. However, various z-direction filter coefficients are conceivable for this, and any of them can produce the same effect.

This embodiment can be used not only for medical X-ray CT apparatuses but also for industrial X-ray CT apparatuses or X-ray CT-PET apparatuses, X-ray CT-SPECT apparatuses combined with other apparatuses.
In the present embodiment, an example in which the channel opening of the X-ray detector is different twice or several times in the z direction or the channel direction has been described, but it is also possible to deal with a difference in opening area smaller than that. . For example, it is possible to perform correction corresponding to a difference in opening area of several percent or less that may occur in the manufacturing process of the X-ray detector.

  In the present embodiment, a multi-row X-ray detector is described. However, in a planar two-dimensional X-ray area detector such as a flat panel X-ray detector or a plurality of planar two-dimensional X-ray detectors. Can produce the same effect.

  In this embodiment, a conventional scan (Axia scan), a cine scan, a helical scan, a variable-pitch helical scan or a helical shuttle scan is described. Can be put out.

It is a block diagram which shows the X-ray CT apparatus concerning one Embodiment of this invention. It is explanatory drawing which looked at the X-ray generator (X-ray tube) and the multi-row X-ray detector in the xy plane. It is explanatory drawing which looked at the X-ray generator (X-ray tube) and the multi-row X-ray detector on the yz plane. It is a flowchart which shows the flow of subject imaging | photography. It is a flowchart which shows schematic operation | movement of the X-ray CT apparatus which concerns on one Embodiment of this invention. It is a flowchart which shows the detail of pre-processing. It is a flowchart which shows the detail of a three-dimensional image reconstruction process. It is a conceptual diagram which shows the state which projects the line on a reconstruction area | region to a X-ray transmissive direction. It is a conceptual diagram which shows the line projected on the detector surface. It is a conceptual diagram which shows the state which projected the projection data Dr (view, x, y) on the reconstruction area. It is a conceptual diagram which shows the backprojection pixel data D2 of each pixel on a reconstruction area. It is explanatory drawing which shows the state which obtains backprojection data D3 by adding all the views to backprojection pixel data D2 corresponding to a pixel. It is a conceptual diagram which shows the state which projects the line on a circular reconstruction area | region to a X-ray transmissive direction. It is a figure which shows the imaging condition input screen of an X-ray CT apparatus. (A) It is a figure which shows a circular arc type | mold multi row X-ray detector.

(B) It is a figure which shows a planar 2D X-ray area detector.
(C) is a diagram showing a two-dimensional X-ray area detector using a plurality of planar X-ray detectors.
It is a figure which shows the X-ray detector from which X-ray detector width differs in a column direction. It is a figure which shows the case where an image is reconstructed using the projection data of the different X-ray detector opening in a conventional scan (axial scan) or a cine scan. It is a figure which shows the case where an image is reconstructed using projection data of different X-ray detector apertures in conventional scan (axial scan) or cine scan at two z-axis coordinate positions. It is a figure which shows the case where an image is reconstructed using the projection data of different X-ray detector apertures in a helical scan. It is a figure which shows each detector row width and row position of the original z direction. FIG. 10 is a diagram illustrating Example 1 in which slice thicknesses are aligned by z-direction filter superimposition processing. FIG. 10 is a diagram illustrating Example 2 in which slice thicknesses are aligned by z-direction filter superimposition processing. FIG. 6 is a diagram showing an example 1 in which X-ray detector channel openings are different in the channel direction. FIG. 10 is a diagram showing an example 2 in which X-ray detector channel openings are different in the channel direction. It is a figure which shows the X-ray projection data from which a channel opening width differs for every channel position. It is a figure which shows the X-ray projection of the original data of a certain column of a certain view. It is a figure which shows the data correction by Example 4-1. It is a figure which shows the data correction by Example 4-2. It is a figure which shows the data division | segmentation by Example 4-3. FIG. 10 is a flowchart showing the flow of image reconstruction processing in Example 4-1 or Example 4-2. FIG. 10 is a flowchart showing a flow of image reconstruction processing in Example 4-3. It is a figure which shows the image reconstruction from the division | segmentation projection data.

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 10 Imaging table 12 Cradle 15 Rotating part 20 Scanning gantry 21 X-ray tube 22 X-ray controller 23 Collimator 24 Multi-row X-ray detector or two-dimensional X-ray area detector 25 DAS (data collection device)
26 Rotating section controller 27 Scanning gantry tilt controller 28 Beam forming X-ray filter 29 Control controller 30 Slip ring dP X-ray detector plane P Reconstruction area PP Projection plane
IC rotation center (ISO)
CB X-ray beam
BC beam center axis
D Multi-row X-ray detector width on the rotation axis

Claims (5)

An X-ray generator and a two-dimensional X-ray area detector with a matrix structure that detects X-rays relative to each other are rotated around the center of rotation between them and transmitted through the subject between them. X-ray data collection means for collecting projection data based on a line;
Image reconstruction means for reconstructing an image based on projection data collected from the X-ray data collection means to obtain a tomographic image;
In an X-ray CT apparatus provided with an image display means for displaying the obtained tomographic image,
The image reconstruction means, when reconstructing the projection data collected by the X-ray data collection means, in the three-dimensional image reconstruction, the same pixel position of the tomographic image to be reconstructed irradiated by the X-ray generator from different positions a plurality of projection data based on X-rays passing through to each of the plurality of projection data area of the X-ray detector aperture different, by applying a load coefficient depending on the area ratio of the X-ray detector aperture, they Weighted addition or z-axis direction which is the body axis direction of the subject in at least one of the projection data so that the X-ray detector openings of the plurality of projection data having different areas of the X-ray detector openings are aligned. X-ray CT apparatus characterized by filtering by weighted addition of them after the one in which obtaining the tomographic image by the image reconstruction based on the projection data after the weighted addition. The X-ray CT apparatus according to claim 1, wherein the weighting factor depending on the area ratio of the X-ray detector opening is a weighting coefficient proportional to the area ratio of the X-ray detector opening. The X-ray passing through the same pixel position of the tomographic image to be reconstructed irradiated from a different position by the X-ray generator is a conventional scan (axial scan) at one position in the z-axis direction that is the body axis direction of the subject. The X-ray CT apparatus according to claim 1, wherein the X-ray CT apparatus is an X-ray irradiated in a cine scan. The X-ray passing through the same pixel position of the tomographic image to be reconstructed irradiated from different positions by the X-ray generator is a conventional scan (axial scan) at a plurality of positions in the z-axis direction that is the body axis direction of the subject. The X-ray CT apparatus according to claim 1, wherein the X-ray CT apparatus is an X-ray irradiated in a cine scan. The X-ray that passes through the same pixel position of the tomographic image to be reconstructed irradiated by the X-ray generator from different positions is an X-ray irradiated in a helical scan. X-ray CT system.

Images