JP2002058666A - X-ray computer-aided tomograph - Google Patents

Abstract

PROBLEM TO BE SOLVED: To improve the quality of an area where an X-ray beam of k-th rotation image and an X-ray beam of k+1-th rotation image overlap. SOLUTION: This X-ray computer-aided tomograph is equipped with an X-ray source 3 to generate X-rays with a width in the axis direction of a subject's body, a multichannel type X-ray detector 5, and a reconstruction processing part 12 to spirally scan by relative movement and relative rotation movement of the X-ray-source and the subject to reconstruct the image. Spiral scanning is executed making the pitch of the spiral scanning non-integer times as much as the basic slice thickness.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to an X-ray computed tomography apparatus (X-ray CT apparatus) for helically scanning an object with cone-beam X-rays.

[0002]

2. Description of the Related Art As shown in FIG. 16, an X-ray tube for generating an X-ray flux and an X-ray detector at a position facing and sandwiching the subject rotate around the subject as shown in FIG. While collecting projection data from various angles. Conventionally, the X-ray flux has been a fan beam-shaped X-ray flux, and the detector has been a one-dimensional array type detector.

[0003] There are two types of scan systems, a conventional scan system and a helical scan system. The conventional scan method is defined as a scan method in which the X-ray tube orbits the same circular orbit as shown in FIG. Figure 17 shows the helical scan method.
As shown in (b), the X-ray source and the detector rotate continuously around the subject, and the bed on which the subject is placed moves along the body axis in synchronization with the rotation. Given a moving coordinate system that is defined and moves with the subject, this name is used because the X-ray tube draws a spiral trajectory. In the moving coordinate system, the distance in the body axis direction, that is, the Z-axis direction, which is displaced during one rotation of the X-ray tube is defined as a helical pitch.

[0004] In recent years, in the third or fourth generation CT, an X-ray tube which generates a cone beam X-ray beam which spreads in the body axis direction instead of a fan beam X-ray beam, and an X-ray detector Has a two-dimensional array detector in which detector elements are arranged in a matrix such that a plurality of one-dimensional array detectors are stacked in a plurality of rows, for example, N rows in the Z-axis direction.
T is known and is referred to as cone beam CT.

Here, as shown in FIGS. 18 and 19,
Consider an X-ray beam incident on a certain detector row, and define the X-ray beam as the thickness in the Z-axis direction when the X-ray beam passes through the center of rotation (Z-axis, ie, the center of the imaging region FOV). The imaging region in the third-generation Cornbee CT is defined as a cylinder having a radius ω centered on the Z axis.
Several image reconstruction methods are known when scanning is performed by a conventional scan method using a cone beam CT.

(Reference 1) "Practical cone-beam algori
thm ”LAFeldkamp, LCDavis, and JWKress J.Opt.Soc.Am.A / Vol.1, No.6, pp.612-619 / June1984 This reconstruction method is called the Feldkamp reconstruction method from the developer name. This is an approximate three-dimensional reconstruction algorithm obtained by extending the fan-beam (in a two-dimensional plane) reconstruction algorithm, which is a mathematically exact reconstruction method, in the Z-axis direction. (1) Correction and weighting of projection data The projection data is multiplied by a weight corresponding to the Z coordinate to correct the effect of beam hunting (2) Convolution (convolution operation, Convolutio)
n) Convolution operation of the data of (1) and the same reconstruction function as the fan beam reconstruction is performed. (3) Back projection (Back projection operation, Back Proj
section) The data of (2) is added to a point (voxel) on the path (from the focal point to the detector) through which the X-ray has passed.

When the above (1) to (3) are repeated over a predetermined angle (360 ° or 180 ° + fan angle), an image is reconstructed.

After all of the above mathematical discussions, simulation results are described. Since this reconstruction method is originally an approximate reconstruction method, the image quality of the reconstructed image deteriorates as the beam spread in the Z-axis direction, that is, the cone angle increases. Therefore, the cone angle is practically limited in medical equipment and the like.

Regarding these Feldkamp reconstructions, the results of many computer simulations and experiments using II imaging plates have been reported.

In recent years, a relatively large photographing area FO in the Z-axis direction
A combination with a helical scan method using a cone beam CT has been considered as a method for collecting three-dimensional data related to V with high resolution and at high speed. This combination method is introduced in the following literature. (Reference 2) "CT apparatus" Toshiba Hiroshi Aratate, Kyojiro Nambu Japanese Patent Application Laid-Open No. 4-224736, filed on December 25, 1990 (Reference 3) "3D helical scan CT using conical beam projection" Tohoku University Hiroyuki Kudo, Tsukuba University Tsuneo Saito Transactions of the Institute of Electronics, Information and Communication Engineers DII Vol.J74-D-II, No.8, pp.11
08-1114, August 1994 These documents disclose a reconstruction method similar to the Feldkamp reconstruction method. This outline is based on the Fe
Similar to the ldkamp reconstruction method, weighting of projection data, convolution operation with a function, and weighted back projection onto an X-ray path over 360 °. Considering this back projection focusing on a certain reconstruction point (voxel), the data obtained through the path connecting the focal point and the reconstruction point in each view is weighted and added to the reconstruction point, and the 360 ° When the data of the X-ray paths corresponding to the number of views are added, the point is reconstructed.

[0011] As shown in FIG. 20, the helical pitch is described in the document 3 as "the upper end (straight line A) of the X-ray beam emitted from the position a (β) of the X-ray source and the X-ray after one rotation.
The intersection C with the lower end (straight line B) of the X-ray beam radiated from the position b (β + 2π) of the radiation source must be located outside the subject existence area (subject P) from the X-ray source under necessary and sufficient conditions. is there. There is a description. In addition, this is the angle required for the reconstruction point required to perform the above-described image reconstruction, for example, 360 degrees.
It can be easily inferred from the necessity of an X-ray beam of °.

[0012] Even in these documents, there is no description about the effect of discretization of the X-ray beam, that is, each detector channel has a finite and certain size in the channel direction and the Z-axis direction. In addition, although the result of simulation performed by a computer is described, there is no description about the detailed contents, such as an interpolation method.

[0013]

(1) Problems related to reconstruction errors due to the fact that the detection element of one channel has a constant size in the Z-axis direction and in the channel direction. After the above discussion, the results of the computer simulation are merely presented. In the reconstruction processing, the cylindrical imaging region FOV is defined as a group of a plurality of voxels. On a computer, it is possible to arbitrarily generate data at the point where a straight line connecting the voxels to be reconstructed from the focal point reaches the detector plane, so that the effects of these discretizations do not appear. In detectors such as II used in various experiments,
The detection element is formed of a square of 0.5 mm square or less, and its size is practically negligible in practical use, and its vertical and horizontal sizes are isotropic.

On the other hand, in an actual cone beam CT or the like, the size of the detecting element in the channel direction and the Z-axis direction is 1 unit.
It is about mm × 2 mm. Since the X-ray path of the actually measured projection data is recognized as a straight line connecting the X-ray focal point and the center of the channel, in many cases, the actually measured X-ray path is defined by the X-ray focal point and the reconstruction processing. The calculated X-ray path connecting to the center of the voxel is shifted. This deviation induces an error. If the size of the detection element is small enough to be regarded as a point, such as II, there is no problem in disregarding this deviation and the measured X-ray at the position closest to the calculated X-ray path is ignored. Although there is no problem in practical use of the data of the line path, this error cannot be ignored when the size of the detecting element is large as in the detector in the cone beam CT. If the size of the detection element in the Z-axis direction is further increased, this error will be further increased.

(2) Problems Regarding Handling of Data in Overlapping Area When the helical pitch is determined as described in Document 3, as shown by the oblique lines in FIG. , And the (k + 1) -th rotation of the X-ray beam. This means that the information of the hatched overlap area is included in the two projection data. However, in Literature 3, this overlapping area is ignored, and backprojection is performed by selectively using only the X-ray beam having a small intersection angle with the target reconstruction plane, that is, a cone angle, and the information is sufficiently used. I haven't.

(3) Problems related to streaks occurring in the direction of the joint between the reconstruction start angle and the end angle Also, as described in Reference 3, a target reconstruction plane is reconstructed using projection data with a cone angle as small as possible. X of time
FIG. 21B schematically shows the cone angle from the line focal point to the rotation center of the reconstruction plane. Although the reconstruction start angle and the reconstruction end angle are adjacent to each other, the reconstruction start angle is + 10 °, whereas the reconstruction end angle is −10 °, so that the joint direction (broken line) Shows that there is a large gap due to the discontinuity of the cone angle. In addition, there is also the influence of the movement of the subject due to the shift in the data collection time, and a clear streak artifact occurs in this direction.

(4) Problem of small helical pitch Reference 3 describes that "the intersection between the upper end and the lower end of the X-ray beam must be located outside the subject existing area". There is no description about the channel size of the vessel or other discretization. In other words, this is "it is necessary and sufficient that the intersection of the X-ray path passing through the center of the channel in the uppermost row of the detector and the X-ray path passing through the center of the lowermost row of the detector is located outside the subject existence area" (FIG. 22), which means that the helical pitch is considerably reduced. For example, the number of detector rows is N, and the basic slice thickness is Th.
If ick, the distance between the focal point and the center of rotation is FCD, and the effective visual field diameter is FOV, the helical pitch P is: P ≦ Thick × (N / 2−0.5) × (FCD-FOV / 2) /FCD×2.0 (2) ). For example, N = 10, Thick = 2mm, FCD = 500mm, FOV =
If it is 240mm, P = 13.68mm / REV. This is quite small, meaning that the shooting time is very long.

An object of the present invention is to provide a computed tomography apparatus capable of improving the image quality of an overlapping area between an X-ray beam at the k-th rotation and an X-ray beam at the k + 1-th rotation. Another object of the present invention is to provide an X-ray computed tomography apparatus capable of reducing the occurrence of streak artifacts due to discontinuity in cone angle.

[0019]

According to the present invention, there is provided an X-ray source for generating X-rays having a spread in the body axis direction of a subject, and a plurality of rows arranged along the body axis direction of the subject to transmit the X-ray. Multi-channel X-ray detection means for collecting X-rays, helical scanning by relative movement and relative rotation between the X-ray source and the subject, processing the collected data, and back-projecting the processed data An X-ray computed tomography apparatus having means for reconstructing an image by performing the spiral scan with the pitch of the spiral scan being a non-integer multiple of a basic slice thickness. The present invention relates to an X-ray generating X-ray having a spread in a body axis direction of a subject.
X-ray source, multi-channel X-ray detecting means arranged in a plurality of rows along the body axis direction of the subject to collect transmitted X-rays of the subject, relative movement and relative rotational movement of the X-ray source and the subject A reconstructing means for performing a helical scan, processing the collected data, and reconstructing an image by backprojecting the processed data, wherein the reconstructing means comprises a plurality of rotations. The method is characterized in that backprojection data is obtained from a plurality of data having the same phase and different rotations and backprojection is performed on a part or all of the region irradiated with the beam. The present invention provides an X-ray source that generates X-rays that spread in the body axis direction of a subject, and a multi-channel type that is arranged in a plurality of rows along the body axis direction of the subject and collects transmitted X-rays of the subject. X-ray detection means, reconstruction for performing helical scanning by relative movement and relative rotation between the X-ray source and the subject, processing collected data, and back-projecting the processed data to reconstruct an image X-ray computed tomography apparatus having means, wherein the reconstruction means, according to the position of the voxel to be back-projected with respect to the X-ray focus,
The projection data of a predetermined detection element row is selected from the projection data of the plurality of arrayed detection element rows, and back-projection data of the voxel is created from the projection data.

[0020]

Embodiments of the present invention will be described below with reference to the drawings. (First Embodiment) FIG. 1 shows X according to a first embodiment.
It is a lineblock diagram of a computed tomography apparatus. FIG. 2 is an external view of the gantry of FIG. FIG. 3 is a perspective view of the two-dimensional array type detector of FIG. A gantry (also referred to as a gantry) 1 as a projection data measurement system includes an X-ray source 3 for generating a cone beam X-ray flux approximated to a cone, and a two-dimensional array in which a plurality of detection elements are two-dimensionally arranged. The X-ray detector 5 is accommodated. The X-ray source 3 and the two-dimensional array type X-ray detector 5 are mounted on the rotating ring 2 in a state where the X-ray source 3 and the two-dimensional array type X-ray detector 5 face each other with the subject placed on the slide top of the bed 6 therebetween. As the two-dimensional array type X-ray detector 5, a one-dimensional array type detector in which a plurality of detection elements are arranged one-dimensionally is mounted on the rotating ring 2 in a state where a plurality of rows are stacked. here,
One detecting element is defined as corresponding to one channel. X-rays from the X-ray source 3 are applied to the subject via the X-ray filter 4. X-rays that have passed through the subject are detected as electrical signals by the two-dimensional array X-ray detector 5.

The X-ray controller 8 supplies a trigger signal to the high voltage generator 7. The high voltage generator 7 applies a high voltage to the X-ray source 3 at the timing of receiving the trigger signal. As a result, X-rays are emitted from the X-ray source 3.

The gantry bed controller 9 controls the rotation of the rotating ring 2 of the gantry 1 and the slide of the slide top plate of the bed 6 in synchronization. The system controller 10 as a control center of the entire system controls the X-ray controller 8 and the gantry bed controller 9 so as to execute a so-called helical scan in which the X-ray source 3 moves in a spiral trajectory when viewed from the subject. I do. Specifically, the rotating ring 2 continuously rotates at a constant angular velocity, the slide top plate moves at a constant speed, and X-rays are continuously or intermittently emitted from the X-ray source 3 at every constant angle. .

The output signal of the two-dimensional array type X-ray detector 5 is amplified by the data collection unit 11 for each channel and converted into a digital signal. The projection data output from the data collection unit 11 is taken into the reconstruction processing unit 12. The reconstruction processing unit 12 calculates X for each voxel based on the projection data.
Back-projection data reflecting the line absorptivity is obtained. In the helical scan method using cone beam X-rays as in the present embodiment, the imaging area (effective field of view) has a cylindrical shape with a radius ω about the rotation center axis, and the reconstruction processing unit 1
Reference numeral 2 defines a plurality of voxels (three-dimensional pixels) in this photographing area, and obtains back projection data for each voxel. The three-dimensional image data or tomographic image data created based on the back projection data is sent to the display device 14 and visually displayed as a three-dimensional image or tomographic image.

Next, the operation of the present embodiment will be described. Here, the geometry of the system is shown in FIG.
As shown in (b), the number of detector rows N and the basic slice thickness Thic
k, the focal-rotation center distance FCD, the diameter ω of the imaging region (effective field of view), and the helical pitch P. For example, N = 1
0, Thick = 2mm, FCD = 500mm, FOV = 240mm. The basic slice thickness is defined as the thickness of the X-ray beam applied to the detection element for one channel near the imaging region FOV. The helical pitch P is defined as the interval between the spiral orbits of the X-ray source 3, specifically, the moving distance of the slide top that moves during one rotation of the X-ray source 3.

(1) Method of Creating Back Projection Data First, the apparatus side, that is, the reconstruction processing unit 12 is operated by the X-ray focus F
Calculation that the projection data has been obtained along a straight line (X-ray path) connecting the X-ray focal point F and the center of the detection element. Obtained along the connecting X-ray path, in other words, the X-ray path often has a center of sensitivity. X in this calculation
A deviation between the line path and the actual X-ray path can be an error factor that degrades image quality. I want to confirm this point.

FIG. 6 is a schematic diagram showing the relationship between an X-ray beam and a reconstructed voxel in a certain view I (for example, the rotation angle of the X-ray source 3 when the vertex position is set to 0 °). Here, the back projection of the projection data of this view on the voxel V indicated by oblique lines is considered. A straight line FVC connecting the X-ray focal point F and the center of the voxel V is extended, and a point that intersects the detector plane is defined as a point C. FIG. 7 shows the relationship between the point C and the detection element.
Point C is the center point of each of the (n, m) detecting element, (n, m + 1) detecting element, (n + 1, m) detecting element, and (n + 1, m + 1) detecting element. Shall exist between them. If the “center position of each channel” is defined as “the center of gravity of a rectangular channel”, the point C is out of the center of any channel. The data of the nearest detector row and the nearest channel are converted to the desired straight line
If the data is approximated to the data that has passed through VC, the above-described large error occurs in the case of CT. Therefore, in the present embodiment, as shown in Expression (3), the projection data actually measured along four actual X-ray paths existing around the calculated X-ray path FVC is subjected to the above-described processing. From the data, that is, the processed data in the four channels around the point C, linear interpolation is performed by the inverse ratio of the distance between the point C and the center position of each channel, and the obtained interpolation data is converted into the calculated X-ray path. The data is back-projected along a straight line FVC shown in the figure, and is back-projected with a predetermined weight.

[0027]

(Equation 1)

Here, the method of linearly interpolating the four data with the inverse ratio of the distance has been described. However, the four data are nonlinearly interpolated, or the data of 3, 4,. ,, 2N data, or non-linear interpolation of 6,8,, 2N data using data of 3,4,, N columns in the Z-axis direction, or channel direction and Z-axis direction May be used to change the interpolation function.

(2) Method of back-projecting two data on beam overlap region FIG. 8 shows an X-ray beam from the X-ray focal point F (k) at the k-th rotation sandwiching a certain reconstruction plane at the same phase (same view I). And the irradiation range of the X-ray beam from the X-ray focal point F (k + 1) at the (K + 1) th rotation viewed from a direction perpendicular to the Z axis. The irradiation range of the X-ray beam from the X-ray focal point F (k) at the k-th rotation partially overlaps with the irradiation range of the X-ray beam from the X-ray focal point F (k + 1) at the K + 1-th rotation. Helical scan is performed as follows. For voxel V1 outside the overlap region,
Since only the X-ray beam from the focal point F (k) is passing,
Data obtained by processing projection data obtained along the X-ray path from the focal point F (k) through the voxel V1 to a certain detection element is weighted according to the equation (4) and back-projected.

Back (I, k) = 1 / W ² · D (k) (4) where D (k) indicates projection data on which predetermined processing such as convolution operation has been performed. W indicates a distance between two points obtained by projecting the position F of the X-ray source and the reconstruction point V on the xy plane.

Since the X-ray beams from both focal points intersect voxel V2 in the overlapping area, focal point F
After predetermined processing of the projection data obtained with the X-ray beam from (k), data Back (I, k) is obtained by weighting according to equation (4) and back-projecting, and from the focal point F (k + 1) Similarly, data Back (I, k + 1) obtained by back-projecting data obtained by the X-ray beam by weighting according to equation (4) is obtained, and two data Back (I, k) are obtained according to equation (5). ) And Back (I, k + 1) are weighted and added to obtain back projection data Back (I) of the voxel V2.

Back (I) = α · Back (I, k) + (1−α) · Back (I, k + 1) (5) where α = | (Z coordinate of point CF) − ( The Z coordinate of the constituent point) | / (WZ) Then, the CT data V (x, y, z) of the voxel V2 is obtained according to the equation (6) as the integration of the back projection data Back (I) of all the views.

[0033] Here, CF is a point where two beams intersect, and WZ is a range in the Z-axis direction in which two data are back-projected. When the range is the overlapping region itself, the result is as shown in FIG. 8, and is determined by the positional relationship between the focal point and the voxel.

Here, two data Back (I, k) and Back (I, K)
+1) is weighted and added using the weighting coefficient α, but α is not necessarily specified by the proviso of equation (5), but is a weighting coefficient determined by the positional relationship between the focal position and the reconstruction point. Or the weight α may be fixed at a certain value such as 0.5 or 0.7.

In equation (5), two (back-projected) convolved data D (channels, columns) are back-projected to obtain Back (I, k) and Back (I, k + 1). Are obtained and then weighted and added, but backprojection may be performed after weighted addition according to Expressions (7) and (8).

D ′ (I) = α · D (k) + (1−α) · D (k + 1) (7) Back (I) = 1 / W ² · D ′ (I) (8) (3) Regarding extrapolation processing and helical pitch According to the latter document, the helical pitch is defined as “an X-ray beam with respect to the center of the channel in the top row of the detector and an X-ray beam with respect to the center of the channel in the bottom row of the detector. It is necessary and sufficient that the intersection of is located outside the subject existence area. "
Can be interpreted as Here, in order to project back projection data by the above-described four-data linear interpolation on voxels of the entire effective visual field, the above-mentioned intersection point may be present outside the center of the endmost voxel of the effective visual field. This alone increases the helical pitch slightly. For example, a pitch P = 13.72 mm can be obtained under the same conditions as in the past.

Of course, the area where the four data cannot be linearly interpolated is not back-projected, and the helical pitch can be increased by ignoring the image quality deterioration near the outside of the FOV.

However, the Z coordinate of the intersection between the X-ray focal point and the straight line (X-ray path) passing through the center of the voxel and the detector plane is outside (upper or lower) than the center of the channel in the uppermost or lowermost row of the detector. In the case of, the uppermost or lower two
Extrapolation (also referred to as extrapolation) is performed using the data of columns and two channels, and data to be back-projected is created. This has better image quality.

For example, FIG. 9A shows a relationship between an X-ray beam and a reconstructed voxel in a certain phase (a certain view I).
It is the figure seen from the direction perpendicular to an axis. Consider backprojection on voxel V. A straight line connecting the focal point F and the center of the voxel V is extended, and a point that intersects the detector plane is defined as a point C (FIG. 9).
(B)). The point C exists on the n-th channel and the (n + 1) -th channel in the channel direction and on the first column in the Z-axis direction. Point C is clearly above the center Z coordinate of the top detector row. Thus, according to equation (9), the first column and the second column
The back projection data Back (I) is obtained by extrapolating the four data of the n channel and the n + 1 channel in the column with the inverse ratio of the distance.

[0040]

(Equation 2)

When the intersection is below the center of the channel in the bottom row of the detector, extrapolation is similarly performed, and Z (1) → Z in equation (9).
(N), Z (2) → Z (N−1).

Thus, in the case of extrapolating by one detector row, the helical pitch P can be defined by equation (10), and P = 1
It can be 6.72mm, and can be increased by about 22%, shortening the shooting time.

P ≦ Thick × ((N + 2i) /2-0.5) × (FCD-FOV / 2) /FCD×2.0 (10) Here, i = 1 in the equation (10), and one row of the detector is outside. Although the case of performing interpolation is shown, the range of extrapolation is not limited to this. If the image quality deterioration near the end of the effective field of view FOV can be tolerated, the value of the underlined portion of Expression (10) may be increased to N + 2, N + 3,. However, excessively large extrapolation may degrade the image quality, so the range (boundary) of extrapolated backprojection data is determined in advance, and if the helical pitch exceeds the range, the boundary is outside the range. May be given. That is, in equation (9), Z (2), Z
A value for the boundary may be input in (1).

Here, a method of performing linear extrapolation of four data using the inverse ratio of distance has been described.
Linear, or non-linear extrapolation of 6,8,..., 2N data using 3, 4,..., N columns of data in the axial direction may be employed.

In the above extrapolation processing, extrapolation is used when the intersection is located outside the center of the outermost detector. However, the following method may be used.

(1) At the time of scanning, dummy data is provided one column above and below N columns of data in each view. (2) Extrapolation of the data in each of the upper and lower columns is performed using the data of N columns in each view.

For example, the first and second columns of the n-th channel,
From the data in the Nth column, virtual data in the 0th and N + 1th columns are extrapolated as in the following equation (11).

D (n, 0) = 2 × D (n, 1) −D (n, 2) D (n, N + 1) = 2 × D (n, N) −D (n, N−1 )… (11) (3) The virtual detector array calculated in (3) and (2) is regarded as data of N + 2 columns including the 0th and N + 1th columns, and the data of (1) is described. Data to be back-projected is obtained by performing interpolation.

According to the method in which the extrapolation and the interpolation are used together, the fixed processing in (2) can be performed at high speed by general-purpose hardware, and the interpolation in (3) can be performed. Is the same processing as (1), so that it can be realized with an inexpensive configuration without special hardware for extrapolation. In this example, N + 2 virtual detector rows are created by adding one row at a time to the upper and lower rows, but by further increasing the number of virtual detector rows, the equation (10) N + 1 can be changed to N + 2, N +3 ,,, can be supported. (1) is not always necessary.

According to the above (1) to (3), scanning is performed with a helical pitch being determined, and projection data is subjected to predetermined processing to interpolate and backproject to reconstruct an image.

(1) and (3) can also be applied to Feldkamp reconstruction in conventional scan instead of helical scan.

(Second Embodiment) Regarding the system configuration and (1) a method of creating backprojection data, (2) backprojection processing of two data on a beam overlapping area, (3) extrapolation interpolation processing and helical pitch , The same as the first embodiment,
Description will be omitted, and only portions different from the first embodiment will be described.

(4) Method of Determining Helical Pitch In the (2) two-data back-projection processing to the beam overlap region described in the first embodiment, the k-th rotation and the (k + 1) -th rotation
Two projection data from the focal point were back-projected to one voxel and weighted and added. Here, even when backprojection data is virtually created and backprojected by interpolation or extrapolation of four data, the projection data or backprojection data is
Slight artifacts occur due to errors in the interpolation itself and errors that occur when switching between detector rows or channels used for interpolation in adjacent views because they are created by interpolation or extrapolation. .
In this embodiment, the helical pitch P is devised in order to reduce artifacts caused by this switching.

Now, when determining the helical pitch P,
There is also a custom of single slice CT, and the determination is made so as to satisfy Expressions (10) and (12).

P = Thick × I (I is an integer) (12) and P ≦ Thick × ((N + 2i) /2-0.5) × (FCD-FOV / 2) /
FCD × 2.0 (10) Repeat For example, P = 16 mm under the conditions of the first embodiment. However, when the helical pitch is an integral multiple of the basic slice thickness as in equation (12), switching is performed at a timing obtained by dividing 360 ° per revolution into integers, so the same phenomenon always occurs at the same phase at any rotation. Would. For example, FIG. 10 shows the effect of a gap caused by switching detector rows used when obtaining projection data by back interpolation for four voxels for a certain voxel.
FIG. 10A shows the angle (direction) at which a gap occurs due to switching by the X-ray focal point F (N) below the reconstruction plane, and FIG. 10B shows the upper focal point F (N + 1) after one rotation.
Is the switching gap. Switching occurs at the same phase (rotation angle), and backprojection including an error is performed in the same direction. Therefore, artifacts are superimposed, and two data from two focal points are backprojected to one voxel and weighted and added. Despite doing so, the artifacts do not weaken. FIG. 10C shows the result of the weighted addition.

Therefore, in the second embodiment, the helical pitch is determined so as to satisfy the equations (10) and (13).

P = Thick × X (X is not an integer) (13) and P ≦ Thick × ((N + 2i) /2-0.5) × (FCD-FOV / 2) /
FCD × 2.0 (10) Repeat That is, the helical pitch is set to a non-integer multiple of the basic slice thickness.

At this time, the direction of the gap by the above-mentioned switching is as shown in FIG. FIG. 11A shows the result based on the X-ray focus below the reconstruction plane, FIG. 11B shows the result based on the X-ray focus above, and FIG. 11C shows the result of weighted addition. The direction in which the switching of the artifact occurs on the upper side and the lower side is offset from each other, and the artifact is obviously weakened.

The helical pitch is not limited to the above conditions. It is important to "determine the phase at which switching occurs so as to be shifted up and down across the reconstruction plane", and at least the condition that "the helical pitch is a non-integer multiple of the basic slice thickness" is important.

Equation (10) is not a requirement when combining the so-called half reconstruction of 180 ° + fan angle with this Feldkamp reconstruction, as is well known, for example, for fan beam reconstruction. The condition of equation (13) is essential.

Further, the effective visual field FO is usually used in a clinical setting.
V is M size for abdomen and S for head
Change the size, and so on. The FOV of the helical pitch conditional expression (10) is fixed at W = FOV / 2 = 120 mm in the latter document, but is not fixed but variable according to the FOV size of the imaging target site. Thus, when the photographing area is small, the image can be sent at a large helical pitch.

For example, the FOV size is changed from 240 mm to 1
In the case of 20 mm, in the case of 360 ° reconstruction, the expression (1)
According to 0), the upper limit of P is from 16.72 mm to 19.36.
mm, and if an appropriate value is selected from Equation (13), P becomes 1
It is changed from 6.5 mm to 18.5 mm. Or, conversely, the FOV size may be determined from the helical pitch.

A helical scan is performed at the helical pitch determined as described above, and an image is reconstructed in the same manner as in the first embodiment.

(Third Embodiment) System Configuration
The method of determining (3) extrapolation and helical pitch and (4) the method of determining helical pitch are the same as in the first and second embodiments.

(5) Method of Creating Back Projection Data The weighting of the projection data and the convolution with the function are the same as those in the literature described above.

In (1) of the first embodiment, a method of generating backprojection data mainly by linear interpolation of four data has been described in detail. In the third embodiment, a method for creating backprojection data by nonlinear interpolation will be described in detail.

The weight coefficient curve of the nonlinear interpolation is a Gaussian function,
Although there are countless trigonometric functions and higher-order functions, a relatively simple method will be described here.

(1) Assume that the size of the voxel V to be reconstructed in the Z-axis direction is Hv (see FIG. 12). (2) Consider J points arranged at equal intervals ΔZ up and down in the Z-axis direction from the center Vc of the voxel V. Here, it is assumed that ΔZ = Hv / J and J = 5. They are called MicroV (1), MicroV (2) ,, from the top.

(3) Consider MicroV (1) at the top of the J points assumed in (2). In one view I, X
Extend the straight line connecting the line focal point and the point MicroV (1), perform linear interpolation of 4 data using a total of 4 data of 2 rows and 2 channels closest to the point of intersection with the detector surface, and backproject to MicroV (1) To obtain Back-MicorV (I, 1). (4) Backprojection is performed on all MicroV (J) in the same manner as above to obtain Back-MicroV (I, J).

(5) According to equation (14), all Micro
The back-projected value Back-MicroV (I, J) is averaged with respect to V (J) to obtain a back-projected value Back (I) to voxel V. As described above, it is possible to perform complicated non-linear interpolation with an inexpensive system configuration and hardware configuration capable of performing only linear interpolation. In addition, according to this interpolation method, a beam passing through the entire voxel is back-projected instead of back-projecting data on a straight line passing through the center of the voxel. I can say.

As a result, as shown in FIG. 13, two to four rows of projection data are appropriately used depending on the spread of the X-ray beam passing through the voxel, that is, the position of the voxel (distance from the focal point). One voxel back projection data is created.

In the above (2), the voxels V are divided at equal intervals, but they need not necessarily be at equal intervals. Although the number of divisions J is set to 5, the number of divisions is arbitrary. In (3), linear interpolation of four data is used. However, the present invention is not limited to this, and any interpolation method may be used. In (5), the averaging is performed when obtaining the back projection value to the voxel V. However, weighted addition or the like may be used. By changing the division interval and the number of divisions in (2) and performing weighted addition in (5), more complicated nonlinear interpolation can be performed.

Further, in the above, back projection is performed by performing linear interpolation on each of the four points with respect to the J point. Non-linear interpolation may be performed using a function that considers the beam spread due to position dependence. Equation (15) of the weight Wt (a certain detector) for a certain detector row is a function of the position of the intersection and the position of the voxel, and is given as follows.

Wt (a certain detector row) = F (intersection point, voxel position) (15) Alternatively, the following equation (16) may be applied by simplifying by ignoring the position dependency. At this time, if the weight is set to be equal to the weight of Expression (15) at the rotation center, the image quality deterioration will not be noticeable.

Wt (a certain detector row) = F (intersection position) (16) For example, when back-projecting onto a certain voxel, the weights from the intersection position to the first to Nth columns are expressed by the following equation (16). Let's say

Wt (1) = 0.0, Wt (2) = 0.0, Wt (3) = 0.3, W
t (4) = 0.6, Wt (5) = 0.1, Wt (6) = Wt (7) = ... Wt (N) =
0.0 When the weight in the channel direction is Wt-CH, the data to be back-projected is given by the following equation (17).

[0077] In the first method, since the weight of each point in the back projection is almost the same, the data may be averaged before the back projection. That is, after performing the processes of (1) and (2) in the same manner, the MicroV at the top of the J points assumed in (3) and (2)
Consider (1). In view I, focus and point MicroV
A straight line connecting (1) is extended to perform four-point linear interpolation using a total of four data of two rows and two channels closest to the point of intersection with the detector surface to obtain Pre-Back-MicroV (I, 1). (4) Four-point linear interpolation is performed for all MicroV (J) in the same manner as above to obtain Pre-Back-MicroV (I, I).

(5) Pre-Back for all MicroV (J)
-Value to add and average MicroV (I, J) and backproject to voxel
Pre-Back (I). (6) The above Pre-Back (I) is back-projected to a value Back (I) back-projected to voxel V. As a result, the calculation time for backprojection can be significantly reduced, and backprojection is performed for all MicroV
The error that can be done only once without performing (J) is almost negligible. However, even in the above, it is not always necessary to perform the averaging, and a weighted addition may be used.

(6) Two-data back-projection method 2 Since there is a large gap in the direction of the reconstruction start angle and the end angle for 360 °, there is a problem that a clear streak is generated on the image. Also, when the two-data back-projection method onto the beam overlapping area is performed, an image having a low density and a good density resolution is obtained, but the skewed data having a large cone angle results in a somewhat thicker effective slice thickness. In clinical practice, thinner effective slice thickness may be prioritized over noise, and the method will be described.

If only the angle of the X-ray beam with respect to the XY plane, that is, only the X-ray beam having a small cone angle can be back-projected, the effective slice thickness becomes small. Therefore, a very limited area (Border Area) in the overlapping area of the X-ray beams from the X-ray focal points F (k) and F (k + 1) at the K-th rotation and the K + 1-th rotation, that is, the K-th rotation Two data are back-projected to the voxels in the region where the cone angles of the X-ray beams from each of the X-ray focal points F (k) and F (k + 1) at the (K + 1) th rotation are substantially the same, and overlap. For voxels outside the region, even if two X-ray beams overlap,
Back projection is performed by selectively using only the projection data of the X-ray beam having a small cone angle.

FIG. 15A shows an example of a Border Area. At this time, the focus F (k) having a small cone angle is applied to the voxel in one area A in the overlapping area and outside the border area.
+1) is back-projected from the projection data obtained by the X-ray, and the other area B in the overlapping area and outside the Border Area
Projection data obtained by X-rays from F (k) is back-projected to the voxels inside. However, voxels in a Border Area having substantially the same cone angle from the two focal points are back-projected from the two focal points F (k) and F (k + 1) to the respective reconstruction points, and Back (I, k), Back (I, k + 1) are obtained, and weighted addition or averaging is performed on each of them to obtain a back projection value Back (I) to the reconstruction point.

As shown in FIG. 15B, considering a certain reconstruction plane, when the difference between the cone angles of the upper and lower focal points sandwiching the reconstruction plane is large, only the beam with a small cone angle is back-projected, and the difference is reduced. When the distance is small, back projection is performed from both the upper and lower focal points and weighted addition is performed. That is, instead of reconstructing a single tomographic image with 360 ° projection data, the overlapping area is weighted and added using slightly overlapping 360 ° + α data to create 360 ° reconstruction data. , 1
The tomographic images are reconstructed. When performing weighted addition, the weights may be changed linearly or non-linearly in accordance with this angle.

Here, the weighting corresponding to the angle will be described. FIG. 23 is a diagram showing that data collected at angles on the horizontal axis is back-projected with weights on the vertical axis in FIG. FIG. 23 (a) is a diagram shown for reference, and shows a method of selecting and backprojecting data having a smaller cone angle even in an overlapping area by a conventional backprojection method. The angle of reverse shadowing is 2π, that is, 360 °. FIG. 23 (b) and FIG.
(C) is a method in which the overlap angle is α and the data obtained by two rotations are weighted and back-projected. FIG. 23B illustrates linear weighting, and FIG. 23C illustrates non-linear weighting. For the purpose of suppressing the influence of the gap, it is ideal to use a nonlinear weight function in which the weight itself is continuous, the first derivative of the weight is continuous, and the maximum value of the first derivative is minimized.

As a result, clear artifacts generated in the directions of the reconstruction start angle and the end angle can be eliminated or attenuated.

The shape of the Border Area and the range (angle) of the overlapping area are not limited to the above, and may have an in-plane position dependency as shown in FIG. Also, in the overlap region, the two beams are not limited to linear or non-linear weighted addition. Weighted addition with a constant weight or averaging may be used.

The above two (back-projected)
Convolved data D (channel, column)
Are back-projected to obtain Back (I, k) and Back (I, k + 1), and then weighted and added. However, backprojection may be performed after weighted addition. That is, equation (7) obtained by modifying equations (4) and (5)
Is adopted.

D ′ (I) = β · D (k) + (1−β) · D (k + 1) Back (I) = 1 / W ² · D ′ (I) (7) Repeated X-ray The difference between the two types of weighting methods for the region irradiated twice is described again.

(1) For data at a certain angle, the weight of each voxel in the slice plane depends on the position, and the method described in FIG. 14 will be described. For example, 2 in the same axial plane of FIG. 24A (that is, the Z coordinate is the same)
Consider two voxels V1 and V2. Voxels V1 and V2
Is irradiated with an X-ray beam from two focal points, F (k) and F (k + 1).

At this time, a straight line L (V1) passing through the voxel V1 and being parallel to the Z axis (rotation axis) is considered, and the straight lines L (V1) and F (K) are considered.
And D and C are the intersections with the ends of the X-ray beam emitted from F and (K + 1). At this time, the weight W (k) when back-projecting the data of each voxel focal point F (K) on this straight line
Is as shown in FIG. Similarly, focus F (k + 1)
Weight W (k + 1) when back-projecting the data of FIG.
(C).

Similarly, the weights W of the focal points F (k) and F (k + 1) on a voxel on a straight line passing through a voxel V2 which is in the same axial plane as the voxel V1 (ie, has the same Z coordinate).
(k) and W (K + 1) are as shown in FIG. FIG.
As shown in FIG. 24 (b) and FIG. 24 (d), the weights W (k) of V1 and V2 are different, that is, even when data obtained from the same focal point is reversely shadowed in the same axial plane. , The weight of each beam at the time of back projection depends on the positional relationship between the voxel and the focal point in the axial plane.

(2) Weights of All Voxels in Slice Plane are the Same for Data at a Certain Angle The method described in FIG. 15 will be described. FIG. 26 (a)
Considers voxels V1 and V2 on the same axial plane and voxel V3 on an axial plane different from them. All three voxels are irradiated with X-ray beams from two focal points F (k) and F (k + 1).

In this method, as described above, the two beams are reversely reflected only in the limited area of the overlapped area. Therefore, the voxel V3 is back-projected based on the projection data obtained by the focusing operation with the smaller cone angle. Only the data is reverse shadowed.

Next, considering a straight line L (V1) passing through the voxel V1 and parallel to the Z-axis, an area in which two data are overlapped and back-projected is set as a Border Area in FIG. The intersections between V1) and the Border Area are B1 and B2. At this time, the weight W (k) when the data of the focal point F (k) is reversely reflected on each voxel on this straight line is shown in FIG.
The weight W (k + 1) when back-projecting the data at the focal point F (k + 1) is as shown in FIG. Note that the region where the two data are weighted and backprojected is different from FIG.

The weight W of the focus F (k), F (k + 1)) on a voxel on a straight line passing through the voxel V2 in the same axial plane as the voxel V1 (that is, the same Z coordinate).
(k) and W (k + 1) are similarly calculated from the intersection with the Border Area.
The result is as shown in FIG. This is the same as the weighting diagram 25 (b) for the voxel V1. That is, the weight is (1)
Does not depend on the spread of the beam, that is, does not depend on the position of the voxel in the axial plane, the weight of the two beams when back-projecting all voxels in the same axial plane becomes the same.

That is, weight expression by angles as shown in FIG. 15B or FIG. 23 becomes possible. This indicates that it can be realized with a simple configuration when it is realized by hardware, and data with a large cone angle is not used for image reconstruction, so that the image quality is also improved.

In the system configuration described above, scanning is performed at the helical pitch determined by (3) extrapolation processing and helical pitch, and (4) helical pitch determination method 2. (1)
Create back projection data by the method of creating back projection data,
(6) An image is reconstructed by backprojection using the two-data backprojection method 2.

The present invention can be implemented in various modifications without being limited to the above embodiment.

[0098]

According to the present invention, it is possible to improve the image quality of the overlapping area between the X-ray beam at the k-th rotation and the X-ray beam at the (k + 1) -th rotation. ADVANTAGE OF THE INVENTION According to this invention, generation | occurrence | production of the streak artifact by discontinuity of a cone angle can be reduced.

[Brief description of the drawings]

FIG. 1 is a configuration diagram of an X-ray computed tomography apparatus according to a first embodiment.

FIG. 2 is an external view of the gantry of FIG.

FIG. 3 is a perspective view of the two-dimensional array type X-ray detector of FIG. 1;

FIG. 4 is an explanatory diagram of various parameters.

FIG. 5 is an explanatory diagram of a helical pitch.

FIG. 6 is an explanatory diagram of an X-ray path in calculation.

FIG. 7 is an explanatory diagram of an interpolation process.

FIG. 8 is a diagram showing an X-ray path intersecting with voxels in an overlapping area.

FIG. 9 is an explanatory diagram of extrapolation processing.

FIG. 10 is an explanatory diagram relating to generation of an artifact due to switching between the kth rotation and the (k + 1) th rotation of projection data used for conventional interpolation.

FIG. 11 is an explanatory diagram related to reduction of artifacts by switching between k-th rotation and k + 1-th rotation of projection data used for interpolation according to the second embodiment.

FIG. 12 is a diagram illustrating J points vertically arranged at equal intervals ΔZ in the Z-axis direction from a center Vc of a voxel V applied by nonlinear interpolation according to the third embodiment.

FIG. 13 is a diagram showing a relationship between projection data used for nonlinear interpolation and voxel positions according to the third embodiment.

FIG. 14 is a view showing a Border Area in an overlapping area.

FIG. 15 is an explanatory diagram of the shape of a Border Area and the dependence of the range of the overlapping area on the in-plane position.

FIG. 16 is a perspective view of a one-dimensional array type X-ray detector.

FIG. 17 is an explanatory diagram of a conventional scan and a helical scan.

FIG. 18 is an explanatory diagram of a basic slice thickness.

FIG. 19 is an explanatory diagram of cone beam scanning.

FIG. 20 is an explanatory diagram of conventional handling of projection data for voxels in an overlapping area.

FIG. 21 is a diagram showing an overlapping area.

FIG. 22 is an explanatory diagram of a conventional helical pitch determination method.

FIG. 23 is a diagram showing a weight function for an angle change.

FIG. 24 is a diagram showing a weighting function with respect to a change in the position of a voxel and a focal point in an axial plane.

FIG. 25 is a view showing a weight function corresponding to a Border Area.

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 ... Gantry, 2 ... Rotating ring, 3 ... X-ray source, 4 ... X-ray filter, 5 ... 2-dimensional array type X-ray detector, 6 ... Bed, 7 ... High voltage generator, 8 ... X-ray controller, 9 ... gantry bed controller, 10 ... system controller, 11 ... data collection unit, 12 ... reconstruction processing unit, 13 ... display device.

Claims (7)

[Claims] An X-ray source for generating X-rays extending in the body axis direction of a subject, and a multi-channel type for collecting transmitted X-rays of the subject arranged in a plurality of rows along the body axis direction of the subject. X
X-ray detection means, and means for performing a helical scan by the relative movement and relative rotational movement of the X-ray source and the subject, processing the collected data, and reconstructing an image by back-projecting the processed data. The X-ray computed tomography apparatus according to claim 1, wherein the spiral scan is performed by setting a pitch of the spiral scan to a non-integer multiple of a basic slice thickness. 2. The X-ray computed tomography apparatus according to claim 1, wherein the basic slice thickness is a thickness in an imaging region of an X-ray beam applied to a detection element for one channel. 3. The X-ray according to claim 1, wherein the pitch is determined such that a phase at which switching by an X-ray focal point required for reconstruction occurs is vertically shifted across a reconstruction plane. Computer tomography equipment. 4. An X-ray source for generating X-rays extending in the body axis direction of the subject, and a multi-channel type arranged in a plurality of rows along the body axis direction of the subject to collect transmitted X-rays of the subject. X
X-ray detection means and reconstruction means for performing helical scanning by relative movement and relative rotation between the X-ray source and the subject, processing collected data, and reprojecting the processed data to reconstruct an image In the X-ray computed tomography apparatus having, the reconstruction means, for some or all of the region irradiated with a plurality of rotation beams, to obtain back projection data from a plurality of data of the same phase and different rotations, An X-ray computed tomography apparatus characterized by performing back projection. 5. The reconstructing means obtains backprojection data for a part of the overlapping area using a plurality of data at the same phase and different focal positions, and obtains the partial projection data for the part of the overlapping area. 5. The X-ray computed tomography apparatus according to claim 4, wherein backprojection data is obtained using data at one focal position except for the region. 6. The reconstructing means weights and adds projection data of more than 360 degrees to the overlapping area to obtain 3
The X-ray computed tomography apparatus according to claim 4 or 5, wherein data for 60 degrees is obtained, and an image is reconstructed based on the data. 7. An X-ray source for generating X-rays extending in the body axis direction of the subject, and a multi-channel type arranged in a plurality of rows along the body axis direction of the subject to collect transmitted X-rays of the subject. X
X-ray detection means and reconstruction means for performing helical scanning by relative movement and relative rotation between the X-ray source and the subject, processing collected data, and reprojecting the processed data to reconstruct an image In the X-ray computed tomography apparatus having the above, the reconstruction means is configured to determine a predetermined detection element from among the projection data of the plurality of arrayed detection elements according to the position of the voxel to be back-projected with respect to the X-ray focus. An X-ray computed tomography apparatus, wherein projection data of a row is selected, and back-projection data of the voxel is created from the projection data.