JP2000298105A - Computed tomography device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a computed tomography device which can contribute to making an image high quality. SOLUTION: An X-ray beam 3 emitted from an X-ray tube 1 and applied to a subject 4 to be detected is two dimensionally detected by X-ray I, I, 2, an interpolating function which has an approximately trapezoidal shape acts on a data matrix detected in a slice extractive part 11. Interpolating values on a slice line in the data matrix, which corresponds to a line in which a tomographical face 73 crossing at right angles with a rotation axis 72 is intersected with the X-ray I, I, 2, are precisely extracted. Furthermore, a distortion by a deviation of the detection position is corrected by applying a cosine function in a distortion correction part 14. Then, reconstruction of an image is processed in a reconstruction part 15.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a computer tomography apparatus for obtaining a tomographic image of a subject without destruction.

[0002]

2. Description of the Related Art FIG. 43 shows a general configuration of a conventional computed tomography apparatus. X-rays generated by the X-ray tube 201 become X-ray beams 203 on a fan by the collimator 202, pass through the subject 204, and several hundred detection elements are arranged in an arc at equal angles around the focal point F. Detector 2
06. When the fan beam has only a narrow spread, the scanning mechanism 205 collects transmission data by linearly moving (translating) the subject in a direction parallel to the output surface of the collimator 202, and then rotates the subject by a certain angle (rotation). ), And a translate / rotate scan (hereinafter referred to as “TR scan”) is repeated over an angle of 180 °. In addition, when the fan beam has a spread covering the imaging region, a rotation / rotate scan (hereinafter referred to as “RR scan”) that gives only a rotation operation to the subject is performed.
To do). The transmission data detected in this way is
The data is converted into a digital signal by the data collection unit 207 and sent to the data processing unit 208.

In a data processing unit 208, first, a difference in gain for each detection element is corrected by an air correction unit 212, and a logarithmic (LOG) conversion unit 213 converts the object 204 into an object.
Projection data corresponding to the line integral of the absorption coefficient of the X-rays within is calculated. Subsequently, the reconstruction unit 215 mainly performs a filtered back projection method (Filtered Back Projecti) from the projection data.
The image reconstruction is calculated using the on method (hereinafter, referred to as the “FBP method”), and the obtained image is displayed on the display 209. The transmission data is corrected by offset correction, source intensity correction, etc. in addition to air correction, but is not shown in the figure. The FBP method is described in “CT Scanner” (edited by Yoshinori Iwai: Corona, first edition, 1979). As shown in the figure, there are two types of detectors: one that detects X-rays one-dimensionally and one that detects X-rays two-dimensionally. There are detectors combining a multiplier tube (image intensifier) and a television camera, and a flat solid state detector in which a scintillator is attached to a photodiode array.

[0004]

In recent years, there has been a strong demand for an examination using a CT scanner, and there have been many demands for obtaining tomographic images of various subjects. For this reason, the demand for higher quality images tends to increase.

[0005] The present invention has been made in view of the above,
An object of the present invention is to provide a computed tomography apparatus which can contribute to high quality of an image.

[0006]

In order to achieve the above object, according to the present invention, a rotation of a pair of a radiation source and a detector arranged opposite to each other and an object located therebetween is controlled. Given, in a computer tomography apparatus to obtain a tomographic image of the subject using a data sequence obtained by detecting the radiation from the radiation source transmitted through the subject by the detector,
The gist of the present invention is to include a data interpolating unit that obtains an interpolated value between data of the data sequence by applying a cosine function to the data sequence.

In the present invention, when a cosine function is applied to the data sequence, noise may be canceled when data is interpolated, as in conventional linear interpolation. An interpolated value obtained by using two adjacent data values and an interpolated value that uses one data value in which noise remains as it is in the data sequence are mixed so as to prevent non-uniform noise. I have.

According to a second aspect of the present invention, radiation from a radiation source that transmits a subject is given relative rotation between a set of a radiation source and a detector that are arranged to face each other and a subject located therebetween. In a computed tomography apparatus that obtains a tomographic image of a subject using a data matrix obtained by two-dimensionally detecting the radiation source with a detection surface of a detector, a cosine function is applied to the data matrix to generate a radiation source. The gist of the present invention is to have a data interpolating means for obtaining an interpolated value on a line in the data matrix corresponding to an intersection line in which the imaging tomographic plane orthogonal to the axis of rotation intersects the detection plane.

In the present invention, a cosine function is applied to the data matrix to obtain an interpolation value corresponding to a line on which the imaging tomographic plane intersects the detector, so that the imaging tomographic plane can be used as a detection surface. Even if the data matrix is inclined and intersected, the data sequence on the tomographic plane which is originally required for the calculation of the reconstruction is accurately obtained.

According to a third aspect of the present invention, a pair of a radiation source and a detector arranged opposite to each other and a subject located therebetween are rotated relative to each other and a displacement in the direction of the axis of rotation is provided. A tomographic image of a subject is obtained by using a data matrix obtained by two-dimensionally detecting radiation from a radiation source that has passed through the detection surface of the detector. Within an extraction range surrounded by two lines that are arranged substantially symmetrically and in parallel with respect to an intersection line in which the imaging tomographic plane orthogonal to the axis intersects the detection plane,
A straight line parallel to the intersection line is caused to follow the movement of the target tomographic plane of the subject due to the displacement, and an interpolated value on a line in the data matrix corresponding to the straight line is obtained by applying a cosine function to the data matrix. The gist of the present invention is to have data interpolating means.

In the present invention, a straight line parallel to a line where the imaging tomographic plane intersects the detection plane within the extraction range is caused to follow the movement of the target tomographic plane of the subject, and the interpolation value corresponding to the straight line Is obtained, in tomographic imaging by helical scanning, a data sequence for the target tomographic plane of the subject can be obtained for all angles of the rotation.

According to a fourth aspect of the present invention, there is provided a radiation source which imparts relative rotation to a set of a radiation source and a detector which are arranged to face each other and a subject located therebetween, and transmits the subject. In a computer tomography apparatus that obtains a tomographic image of a subject by using a data matrix obtained by two-dimensionally detecting the subject by a detection surface of a detector, the displacement of the rotation in the axial direction relative to the subject is determined. The gist of the present invention is to have a displacement means for changing the tomographic imaging position by giving the same and a data interpolation means similar to the third aspect.

In the present invention, a straight line parallel to a line where the imaging tomographic plane intersects the detection plane within the extraction range follows the movement of the target tomographic plane of the subject, and the interpolation value corresponding to the straight line Is obtained, a scan image (scanogram image) of the subject along the target tomographic plane can be obtained in scanogram imaging by scanning the displacement.

The present invention according to claim 5 provides the invention according to claims 1 to 4
In the computer tomography apparatus described above, the data interpolating means may act on the data matrix with a trapezoidal function or a function obtained by replacing a trapezoidal inclined portion with a cosine function.

In the present invention, by applying a function of trapezoidal function or a function in which a trapezoidal slope portion is replaced by a cosine curve to the data matrix, the same effect as in the case of applying the cosine function is obtained. Is to be obtained.

According to a sixth aspect of the present invention, there is provided a radiation source which transmits a subject through relative rotation of a set of a radiation source and a detector which are arranged to face each other and a subject located therebetween. In a computer tomography apparatus for obtaining a tomographic image of a subject by using a data matrix obtained by two-dimensionally detecting a phantom with a detection surface of a detector, a phantom having a planar gap, By giving a relative displacement in the direction of the axis to the phantom installed perpendicular to the axis, the center line of the gap when the transmission image obtained for the planar gap of the phantom becomes clearest And crossing line setting means for obtaining the position of the line in the data matrix corresponding to.

In the present invention, the position of the gap orthogonal to the rotation axis in the direction of the rotation axis is adjusted, and the position of the center line of the gap when the gap on the transmission image becomes clearest is obtained. By setting the position of the intersection line where the tomographic plane intersects the detection plane, the imaging tomographic plane is accurately perpendicular to the rotation axis.

The present invention according to claim 7 provides the invention according to claims 2 to 6
In the computer tomography apparatus described above, the gist of the invention is to include a light projecting unit that projects a planar visible light beam at the position of the intersection line.

In the present invention, by projecting a planar visible light beam on the intersection line, the line where the imaging tomographic plane intersects the subject is indicated by the visible light beam. The target tomographic plane is easily aligned with the imaging tomographic plane.

The present invention according to claim 8 provides the invention according to claims 2 to 7
In the computer tomography apparatus described above, a gist of the present invention is to include a display unit for displaying a transmission image of the subject and the intersection line on a display device.

In the present invention, by displaying the transmission image of the subject and the intersection line, the target tomographic plane of the subject can be easily adjusted to the imaging tomographic plane by referring to these displays.

According to a ninth aspect of the present invention, there is provided a radiation source from a radiation source which transmits a subject by imparting a relative rotation to a set of a radiation source and a detector which are arranged to face each other and a subject located therebetween. Distance changing means for arbitrarily changing a distance (FCD) between a radiation source and the rotation axis in a computer tomography apparatus which obtains a tomographic image of a subject by using a data sequence obtained by detecting an image by a detector. FCD calibration means for determining a magnification of the dimension of the subject on the tomographic image obtained using the input FCD value with respect to the dimension of the subject, and dividing the FCD value by the magnification to correct the FCD value. Is the gist.

In the present invention, the size of the subject is compared with the size of the subject on a tomographic image obtained by imaging the subject, and the distance between the radiation source and the rotation axis used for calculation is calculated. By making the correction, a tomographic image with high dimensional accuracy can be easily obtained at an arbitrary position from the radiation source.

According to a tenth aspect of the present invention, there is provided a radiation source from a radiation source which imparts a relative rotation to a pair of a radiation source and a detector which are arranged to face each other and a subject located therebetween, and transmits the subject. In a computer tomography apparatus that obtains a tomographic image of a subject by using a data sequence obtained by detecting an image with a detector. The point of the invention is to have a rotation center finding means for finding a position on the data string corresponding to the center of the above data and using the calculated position as the center of rotation.

According to the present invention, the position on the data sequence corresponding to the center of the symmetrical data obtained by adding or averaging the data sequence for 360 ° with respect to an arbitrary subject is obtained. It is possible to calculate the image reconstruction by grasping the position on the data sequence corresponding to the position on the detector where the radiation passing through is incident, eliminating the need to mechanically adjust the displacement of the rotation axis. In addition, a high-quality tomographic image can be easily obtained.

According to the eleventh aspect of the present invention, a relative rotation is given to a set of a radiation source and a detector which are arranged to face each other and a turntable on which a subject placed between them is mounted so as to allow the subject to pass therethrough. In a computer tomography apparatus for obtaining a tomographic image of a subject using a data matrix obtained by two-dimensionally detecting radiation from a radiation source using a detection surface of a detector, the rotation provided on the turntable A plane gap perpendicular to the axis of, and intersection line setting means for determining the position of a line on the data matrix corresponding to the line at the center of the gap when the transmission image obtained for the gap is the clearest, A phantom installation / removal means capable of installing / removing a predetermined phantom in front of a detector, and calibrating distortion on a tomographic image due to a shift in a detection position using a data matrix obtained for the predetermined phantom. And distortion calibration means that the distance measuring means for measuring a distance to the axis of the rotation from the radiation source,
The gist is to have.

According to a twelfth aspect of the present invention, a relative rotation is given to a set of a radiation source and a detector which are arranged to face each other and a turntable on which a subject placed between them is mounted so as to allow the subject to pass therethrough. 12. A computer tomography apparatus for obtaining a tomographic image of a subject using a data matrix obtained by two-dimensionally detecting radiation from a radiation source using a detection surface of a detector, wherein the same gap and intersection as those in claim 11 are provided. Line setting means, a pin provided on the turntable and parallel to the axis of rotation, and
Using the data sequence obtained for the pin, a position on the data sequence corresponding to a position on the detection surface where the radiation passing through the axis of rotation is incident is determined, and a rotation center is calculated as the center of rotation. The gist of the present invention is that it has an output means and the same distance measuring means as described in claim 11.

According to the present invention, a phantom for obtaining the intersection line, a phantom for obtaining the position of the rotation axis, and the predetermined phantom for distortion correction are previously set or installed.・ Be able to separate and measure the distance from the radiation source to the axis of rotation to determine the position of the intersection line and the axis of rotation, and to calibrate the distortion or calculate the distance between the source and the axis of rotation. By setting automatically, installation of a dedicated phantom, calibration of the position where the tomographic plane intersects the detection surface, calibration of the angle of rotation axis installation, calibration of the distance from the radiation source to the rotation axis , Can be avoided by manually adjusting each mechanism.

According to a thirteenth aspect of the present invention, a relative rotation is given to a set of a radiation source and a detector which are arranged to face each other and a turntable on which a subject placed between them is mounted so that the subject can pass therethrough. In a computed tomography apparatus that obtains a tomographic image of a subject using a data matrix obtained by two-dimensionally detecting radiation from a radiation source using a detection surface of a detector, the turntable includes the axis of rotation. A phantom having a pattern of concentric circles centered on a phantom, or a phantom having a grid arranged in parallel at a predetermined interval in a direction parallel to the detection surface, or a pin parallel to the axis of rotation at a position away from the axis Wherein the phantom is fixed, and distortion correction means for correcting distortion on a tomographic image due to a shift in a detection position using a data matrix obtained for the rotary table. To.

In the present invention, by fixing one of the phantoms to the turntable, it is possible to avoid the trouble of manually installing the phantom and to provide the phantom having the concentric pattern or the grid. By using a transmission image from one direction for a phantom and a transmission image from a plurality of directions for a phantom having the pin, it is possible to calibrate distortion due to a shift in the radiation detection position.

According to a fourteenth aspect of the present invention, a relative displacement is provided between the radiation source and the detector arranged opposite to each other and the subject located therebetween, and the radiation from the radiation source transmitted through the subject is detected. In a computed tomography apparatus for obtaining a tomographic image of a subject using a data sequence obtained by detection by a detector, a phantom for evaluating the focus of a radiation source,
Focus evaluation means for evaluating the focus state of the radiation source using the data sequence obtained for the phantom, and radiation source focus adjustment means for adjusting the focus of the radiation source based on the result of the evaluation That is the gist.

According to a fifteenth aspect of the present invention, a relative displacement is provided between the radiation source and the detector arranged opposite to each other and the subject located therebetween, and the radiation from the radiation source transmitted through the subject is detected. In a computed tomography apparatus for obtaining a tomographic image of a subject using a data sequence obtained by detection by a detector, a phantom for evaluating the focus of a detector and detection using a data sequence obtained for the phantom The gist of the present invention is to include focus evaluation means for evaluating the state of the focus of the detector, and detector focus adjustment means for adjusting the focus of the detector based on the result of the evaluation.

According to the present invention, the focus state is evaluated from the ratio between the contrast value and the signal value of the transmitted image of the phantom for focus evaluation,
The focus of the radiation source or the focus of the detector can be adjusted so that this ratio is maximized, so that these focuses can be easily adjusted without depending on the skill of the operator. I can do it.

According to a sixteenth aspect of the present invention, a relative displacement is provided between a radiation source, a detector and an object located between the radiation source and the object, and the radiation from the radiation source transmitted through the object is detected. In a computer tomography apparatus for obtaining a tomographic image of a subject by logarithmically converting a data sequence obtained by a detector, a substantially constant value is added to each data of the data sequence before the logarithmic conversion. The gist of the present invention is to include a noise reduction unit that reduces noise after conversion.

According to a seventeenth aspect of the present invention, a relative displacement is provided between the radiation source and the detector arranged opposite to each other and the subject located therebetween, and the radiation from the radiation source transmitted through the subject is detected. In a computer tomography apparatus which obtains a tomographic image of a subject by logarithmically converting a data sequence obtained by detection by a detector, when detecting radiation transmitted through the subject, the detected value is increased by a substantially constant value. The gist of the present invention is to have a noise reduction unit that reduces noise after logarithmic conversion.

According to the present invention, a relative displacement is provided between the radiation source and the detector arranged opposite to each other and the subject located therebetween, and the radiation from the radiation source transmitted through the subject is detected. In a computer tomography apparatus that obtains a tomographic image of a subject by performing logarithmic conversion and other predetermined processing on a data sequence obtained by a detector, an input value is input to the data sequence obtained by performing the logarithmic conversion. The gist of the present invention is to have a noise compressing means for compressing noise by applying a function of decreasing the output value with an increase in the noise.

According to the nineteenth aspect of the present invention, a relative displacement is provided between the radiation source and the detector arranged opposite to each other and the subject located therebetween, and the radiation from the radiation source transmitted through the subject is detected. In a computer tomography apparatus that performs predetermined processing on a data sequence obtained by a detector and obtains a tomographic image of a subject, a logarithmic transformation is performed on the data sequence with an increase in an output value of the logarithmic transformation. The gist of the present invention is to provide a noise compression unit that simultaneously operates a function for compressing an output value.

According to the present invention, the logarithmic conversion is performed after adding or increasing a substantially constant value to the detected data, or the data sequence after the logarithmic conversion is performed. By applying the compression function to, or by applying the compression function simultaneously with the logarithmic conversion, to reduce noise that is relatively increased due to a small amount of radiation detection I try to make it.

According to a twentieth aspect of the present invention, a relative displacement is provided between a radiation source and a detector arranged opposite to each other and a subject located therebetween, and the radiation from the radiation source transmitted through the subject is detected. A computer tomography apparatus for obtaining a tomographic image of a subject using a data sequence obtained by detection by a detector, wherein data identification means for identifying data in the data sequence corresponding to radiation that has passed through an area outside the subject. ,
An essential point of the present invention is to include an intensity variation correction unit that corrects the influence of the intensity variation of the radiation using the data identified by the data identification unit.

According to the present invention, by correcting the influence of the intensity variation of the radiation using the value of the data in the area, a high-quality image can be obtained without using a dedicated comparison detector for measuring the intensity variation of the radiation. A tomographic image is obtained.

According to the twenty-first aspect of the present invention, the radiation from a radiation source that transmits a subject by imparting a relative rotation to a set of a radiation source and a detector arranged opposite to each other and a subject located therebetween. In a computed tomography apparatus that obtains a tomographic image of a subject by using a data sequence obtained by detecting a detection channel sequence by a detection channel sequence of a detector, the detection channel sequence is divided by 1 / with respect to the path of radiation passing through the rotation axis. The gist of the present invention is to have a data insertion unit for inserting an interpolation value for interpolating between data into a data sequence obtained by shifting the data by four channels.

In the present invention, the detection channel sequence is set to 1 /
The number of data is substantially doubled to improve the spatial resolution by inserting an interpolated value between data in a data string obtained by detecting radiation transmitted through the subject in a state shifted by four channels. In addition, a high-quality tomographic image is obtained.

According to a twenty-second aspect of the present invention, radiation from a radiation source which transmits a subject by giving a relative rotation to a set of a radiation source and a detector which are arranged to face each other and a subject located therebetween. In a computed tomography apparatus for obtaining a tomographic image of a subject by using a data sequence obtained by detecting a position on a detector at which radiation passing through the axis of rotation is incident on the data sequence The point is that there is provided a window function calculating means for applying a window function having an inclined portion centered on a position on the data sequence corresponding to the above.

According to the present invention, by applying the window function to the data sequence, when imaging a large subject, one side of the subject is moved out of the imaging region by shifting the rotation axis to the outside. The data string obtained by detecting the radiation transmitted through the subject in the state can be smoothly changed from the detection value in the imaging region to the 0 value in the non-imaging region, and the ring-shaped false image on the tomographic image is reduced. I am trying to be.

According to the twenty-third aspect of the present invention, the radiation from the radiation source which transmits the object by giving a relative rotation to the set of the radiation source and the detector arranged opposite to each other and the object located therebetween. In a computed tomography apparatus that obtains a tomographic image of a subject by using a data sequence obtained by detecting a detection channel sequence that is linearly arranged at equal intervals in a detector, the data sequence is defined by the rotation axis. The gist of the present invention is to provide a backprojection unit that performs backprojection of image reconstruction by converting the data sequence into data strings at equal intervals on two-dimensional coordinates as an origin.

According to the present invention, the data train obtained by converting the data trains obtained at equal intervals in a straight line into two-dimensional coordinates having the rotation axis as the origin is reconstructed, whereby the data train is obtained. Does not need to be converted into a data string with equal angular intervals, and the occurrence of an error when converting into a data string with equal angular intervals is prevented.

[0047]

Embodiments of the present invention will be described below with reference to the drawings.

FIG. 1 is a diagram showing a computer tomography apparatus (hereinafter, appropriately referred to as a “CT scanner”) according to a first embodiment of the present invention. FIG. 1A shows a front view, and FIG. 1B shows a side view. The CT scanner shown in FIG. 1 includes an X-ray tube 1 that generates an X-ray beam 3 in a cone shape,
Rotating table 5 on which rotary shaft 7 is mounted
A rotation mechanism 19 for giving a rotational movement about the center 2, and an imaging tomographic plane 73 (ideally, X) in which data of a desired target tomographic plane of the subject 4 is collected by moving the rotary table 5 up and down.
A tomographic plane feed mechanism 20 that adjusts to a plane passing through a line focus and orthogonal to the rotation axis), and a visible light 2 along the imaging tomographic plane 73.
A focal length changing mechanism 21 for changing the distance (FCD) from the focal point F to the rotation axis 72 to adjust the photographing magnification; a mechanism control unit 6 for controlling these mechanisms; An X-ray fluorescence intensifier (X-ray image intensifier; hereinafter, referred to as “X-ray II”) 2 that two-dimensionally detects the X-ray transmitted through the subject 4 and converts it into a visible image.
A television camera 7 for converting the visible image into a digital signal; and a data processing unit 8 for performing various processes described below on the digital signal transmitted from the television camera 7 to calculate image reconstruction. And a display 9 for displaying a reconstructed image, other processing results, and a transmission image before the processing. The X-ray II.2 does not have a detecting element, but a digital signal can be obtained by associating a surface for detecting X-rays with a matrix (hereinafter, referred to as a "detection channel matrix"). The data processing unit 8
It is a normal computer, has an arithmetic processing unit (not shown), a storage device that stores necessary programs and the like, a memory, an interface, a keyboard, and the like, and has a line (hereinafter referred to as a line) where the imaging tomographic plane 73 intersects the detection channel matrix. A slice extraction unit 11 for performing a process of extracting an interpolation value on a “slice line”, and an air correction unit 12 for correcting a difference in gain for each detection channel using data obtained by capturing air.
A LOG conversion unit 13 that calculates projection data corresponding to a line integral of an X-ray absorption coefficient using detected data (hereinafter, appropriately referred to as “transmission data”), and a distortion caused by a shift in a detection position. A distortion correcting unit 14 for correcting the image, a reconstructing unit 15 for performing image reconstruction, a slice line setting unit (not shown) for setting the position of the slice line, and displaying the slice line and the like so as to overlap the transmission image of the subject. And a FCD calibration unit (not shown) for correcting an FCD value used for reconstruction from the magnification of the tomographic image. The operator uses the data processing unit 8 to set parameters, select menus, start scanning, monitor the processing state, display and analyze tomographic images, and the like. here,
The slice extracting unit 11 constitutes data interpolating means according to claims 2 to 5, the distortion correction unit 14 constitutes data interpolating means according to claim 1, the light projector 28 constitutes light emitting means,
The slice line setting unit constitutes an intersection line setting unit, the slice line display unit constitutes a display unit, and the FCD calibration unit constitutes an FCD calibration unit.

In such a CT scanner, the operator places the subject 4 on the rotary table 5, sets the magnification with the focal length changing mechanism 21, and sets the target slice of the subject with the tomographic plane sending mechanism 20. The position of the plane is adjusted to the imaging tomographic plane 73.
Subsequently, when X-ray irradiation is performed to start imaging, the rotary table 5 is rotated by 360 degrees at a constant angle or continuously, and
The line II.2 collects transmission data at each position of rotation, and the television camera 7 converts this into digital data and sends it to the data processing unit 8. In the data processing unit 8, first, the slice extraction unit 11 transmits the transmission data d on the slice line.
(I) is obtained by interpolation, and the details of this processing will be described later.

Next, transmission data (air data) previously collected by the air correction unit 12 in the absence of the subject.
gain correction is performed by taking the ratio of d _a. Assuming that the number of the detection channel is i (i is a natural number), the transmission data d
The air correction of (i) is

H (i) = (d (i) −d _off (i)) / (d _a (i) −d _off (i)) (1) Here, d _off (i) is data detected by X-ray II.2 in a state where X-rays are not irradiated. Next, L
The OG conversion unit 13 performs logarithmic conversion and converts the projection data into projection data P (i) corresponding to the line integral of the absorption coefficient. LOG conversion is

P (i) = LOG (1 / h (i)) (2) Next, the distortion correction unit 14 corrects the distortion of the tomographic image due to the shift of the detection position, and the reconstruction unit 15 calculates the image reconstruction using the 360-degree projection data subjected to the above-described processing. Do.

The distortion of the tomographic image caused by the shift of the detection position is caused by the spherical shape of the photoelectric surface 10 of the X-ray II2 (and also by the distortion of the internal electron lens). That is, since the position where the X-rays enter the photocathode 10 spreads outward from the photocathode 10, the X-ray II.
When the image is reconstructed using the projection data P (i) detected by the above, distortion is generated on the image as it goes outward. Therefore, it is necessary to obtain projection data corresponding to a detection channel at a position to be detected by interpolation.

FIG. 2 shows detection channel numbers at actual detection positions. i and the channel No. at the position to be detected.
FIG. 9 is a diagram illustrating a distortion correction curve obtained from a positional relationship with n (n is a natural number). The correction curve is obtained in advance using a predetermined phantom as described later. With this correction curve, the position i of the detection channel corresponding to channel n
(N) can be obtained, and the projection data P (i (n)) at this i (n) (hereinafter, referred to as “P ′ (n)”)
Will be described below.

FIG. 3 is a diagram showing an example of an interpolation function in distortion correction. In the conventional linear interpolation, a function having a triangular shape as shown by a dotted line is used as an interpolation function. However, the use of such an interpolation function has the following problems. That is, at the center of the photocathode 10, X
Since the position i of the detection channel where the line is incident coincides with the position n of the channel to be originally detected, P (i) becomes P ′ (n) as it is, but outside the photocathode 10,
i and n do not match and the data at i (n) is i (n)
Is obtained by linear interpolation from the projection data of two points adjacent to each other, noise in these two points may be canceled. In other words, P ′ (n) has a non-homogeneous noise because a channel in which noise remains as it is and a channel in which noise has been canceled are mixed, and the same applies to each position of rotation. A ring-shaped false image may occur in the constituent image.

Thus, the interpolation function f _{I according} to the present embodiment is described.
(Δi) is

F _I (Δi) = (1 + COS (2πΔi / 3)) / 3 [if | Δi | <1.5] = 0 [if || Δi |> = 1.5] where Δi = Ii (n). … (3) It has a period corresponding to three times the discrete data interval, an amplitude of 2/3, and -π to π
Is a function obtained by moving the function of the cosine in the range of by １ ／ in the positive direction of the vertical axis.

The projection data P (i) is given by the following equation:

'The _{(n) = Σ i f I} (i-i (n)) · P (i) ... (4), the projection data P' [Equation 4] P is converted to (n).

By using such an interpolation function, i (n)
Can be converted using the values of the projection data of the three points closest to the position, and the increase / decrease of the noise can be homogenized as a whole.

This is proved as follows. Assuming that the interpolation phase is x (-1/2 to 1/2) and the noise of the data before interpolation is σ, the noise σ ′ after interpolation is a three-point average with weight,

Equation 5] _{σ '= σ · (f I} 2 (x-1) + f I 2 (x) + f I 2 (x + 1)) 1/2 ... (5) in it represented, specifically the f _I Substituting equation (3) and transforming, the right side is constant at σ / 2 ^1/2 irrespective of x,
The increase / decrease factor of the noise is constant at 1/2 ^1/2 . End of proof.

Therefore, according to the distortion correction of the computer tomography apparatus according to the present embodiment, the interpolation value at a predetermined position between the data is determined based on the data values of the three points closest to this position. Thus, the increase and decrease of noise can be made uniform, and a tomographic image with few ring-shaped false images can be obtained.

The same effect can be obtained even if the interpolation function is a sine function or a function similar to the sine function without using a strictly cosine function. Also,
Not only for distortion correction, but also for interpolation when converting data sequences obtained at equal intervals into data sequences at equal angular intervals, and when changing data intervals to change the magnification of tomographic images. It can also be applied to interpolation.

Next, the processing in the slice extracting section 11 will be described. FIG. 4 is a diagram showing a screen of the display 9 when the detected transmission data is displayed as a transmission image as it is. A pixel No. (i, i) where i is the horizontal axis and j is the vertical axis
j) corresponds to the position of the detection channel matrix, and the transmission data is represented by d (i, j).

Assuming that the slice line at which the imaging tomographic plane 73 intersects the detection channel matrix is jc, in order to obtain a high-quality image, the slice line jc is located at the position where the line connecting the detection channel in the vertical axis direction intersects. Transmission data d (i, j
c (i)) (hereinafter “slice data d ′ (i)” as appropriate)
You need to get However, when the slice line jc is inclined with respect to the detection channel matrix as shown in the figure, transmission data on the slice line jc cannot be detected, which causes image quality deterioration. Therefore,
The slice data d '(i) is obtained by interpolating the transmission data d (i, j) as shown below.

The interpolation is performed using the transmission data d (i, j) within a fixed width SWj (hereinafter, appropriately referred to as “slice width”) centered on the slice line jc. SWj determines the number of samples of the transmission data d (i, j) in the vertical axis direction used for interpolation, and is set in advance.

FIG. 5 is a diagram showing an example of the interpolation function.
A dotted line interpolation function 1 and a solid line interpolation function 2 are shown. The interpolation function 1 is a function having a trapezoidal shape, and the width of the inclined portion is 1,
The height is 1 / SWj and the half width is SWj. The interpolation function 1, g _I (Δj) is

[6] _{w 1 = (SWj -1) /} 2 w 2 = (SWj +1) / 2 and when _{put, g I (Δj) = 1} / SWj [| Δj | <= case of w _1] = (Δj + w _{2) / SWj [-w 2 <} Δj < case of _{-w 1] = (w 2 -Δj} ) / SWj [ the case of _{w 1 <Δj <w 2]} = 0 [| Δj |> = case of w _2] … (6) Here, SWj is reduced, and w ₁ =
Since SWj = 1 when −w ₁ = 0, the lower limit of SWj is 1.

The interpolation function 2, g _I (Δj) shown by the solid line is
The slope of the interpolation function 1 is replaced by a cosine curve so that the width is 1.5.

And put the following equation 7] _{w 1 '= (SWj -1.5)} / 2 w 2' = (SWj +1.5) / 2, g I (Δj) = 1 / SWj [| Δj | <= w 1 In the case of ') = {1 + COS (2π · (Δj + w ₁ ') / 3)} / (2 · SWj) [in the case of −w ₂ '<Δj <−w ₁ '] == 1 + COS (2π · (Δj−w _{1 ') / 3)} /} (2 · SWj) [w 1'<Δj<w 2 ' for] = 0 [| Δj |> = w 2' represented by] ... (7). Here, when w ₁ ′ = −w ₁ ′ = 0, SWj =
Since it is 1.5, the lower limit of SWj is set to 1.5. Also,
At this time, it matches the interpolation function f _I (Δi) shown in FIG.

Using the interpolation function g _I (Δj), the transmission data d (i, j) is given by the following equation:

Equation 8] _{d '(i) = Σ j} g I (j-jc (i)) · d (i, j) by (8), slice data d' is converted to (i).

When the interpolation function 1 is used, there is an advantage that the calculation speed is fast. On the other hand, when SWj is set to the lower limit value 1, the function has a triangular shape, as described with reference to the conventional interpolation function shown in FIG. There is a problem that the increase and decrease of the noise are not uniform.

On the other hand, when the interpolation function 2 is used, the slice width SWj is set to 1.5 + n (n is 0, 1, 2, 3,...)
If it chooses, the increase and decrease of noise will be homogenized.

This is proved as follows. Assuming that the interpolation phase is x (− ／ − １ ／) and the noise of the data before interpolation is σ, the noise σ ′ after interpolation is

Equation 9] _{σ '= σ · (Σ j} g I 2 (x + j)) is represented by ^1/2 (9) Specifically deformed by substituting equation (7) to g _I, SWj Is 1.5 + n, the coefficient on the right-hand side, that is, the noise increase / decrease factor, is constant regardless of x.

The proof is over.

Therefore, according to the slice extraction of the computed tomography apparatus according to the present embodiment, the interpolation value on the slice line where the imaging tomographic plane intersects the detection channel matrix is extracted, so that the detection channel matrix becomes A high-quality tomographic image can be obtained even when inclined with respect to the imaging tomographic plane.

Although the interpolation function 2 uses a cosine curve for the inclined portion, a similar effect can be obtained by using a sine curve or a curve similar thereto, even if the curve is not exactly a cosine curve. Obtainable.

Further, it is also possible to provide a plurality of lines parallel to the slice line jc and extract a slice for each line. As a result, a plurality of tomographic images can be created by one imaging. However, there is a limit because the image quality decreases as the distance from the slice line jc increases.

Next, the rotary table 5 is rotated by a fixed angle by the rotary mechanism 19, and the rotary table 5 is linearly moved upward by the tomographic plane feed mechanism 20 by a predetermined interval in the direction of the rotary shaft 72 to transmit the transmission data. The processing performed by the slice extracting unit 11 in the spiral scan that repeatedly performs the above-described operations will be described.

FIG. 6 is a diagram showing a screen of the display unit 9 when a transmission image detected in the spiral scanning is displayed. When the slice line jc, at which the imaging tomographic plane 73 intersects the detection channel matrix, is moved up and down in parallel with the rotation angle of ± 180 °, it corresponds to the 360 ° rotation angle surrounded by two lines. The range is defined as an extraction range, and a line parallel to the slice line and following the movement of the target tomographic plane of the subject within the extraction range is defined as a follow-up slice line 24.

FIG. 7 is a diagram showing slice extraction in spiral scanning. A case of obtaining an image of the target tomographic plane will be described below. First, when the target tomographic plane that moves upward as the rotation angle advances enters the above-described extraction range (for convenience, the rotation angle at this time is set to 0 °), it corresponds to the target tomographic plane. A tracking slice line is set, and the transmission data (hereinafter, referred to as the
“Follow-up slice data”) is obtained at the position of each rotation angle by the same processing as the above-described slice data extraction.

Here, the distance between the target tomographic planes (hereinafter referred to as “cross-sectional pitch”) Sp is set to the distance that the subject 4 moves upward by a linear motion during one rotation (hereinafter referred to as “spiral pitch”). If the rotation angle is set to 1/2 of Zp, the target slice enters the extraction range at the rotation angle of 180 °. Therefore, a new tracking slice line is provided, and the tracking slice data of the target slice plane is obtained. Start extraction. When the rotation angle becomes 360 °, slice data for 360 ° for the target tomographic plane has been obtained, and this is sent to the reconstruction unit 15 via the air correction unit 12 and the like to perform image reconstruction. For the target tomographic plane to, the following slice line is similarly provided, and the following slice data is extracted.

Therefore, according to the slice extraction in the helical scan of the computed tomography apparatus according to the present embodiment, the extraction corresponding to the rotation angle of ± 180 ° around the slice line is performed for each of the plurality of target tomographic planes. 360 ° on a line parallel to the slice line within the range
By obtaining the following slice data by interpolation, a high-quality tomographic image can be obtained for each target tomographic plane.

Note that the sectional pitch Sp is not limited to の of the spiral pitch Zp. If a finer sectional pitch is used, a plurality of following slice lines corresponding to the sectional pitch are provided within the extraction range. Thus, a plurality of following slice data can be extracted by the same processing as described above.

The extraction range is not 360 ° but 180 ° + fan angle (fan-shaped angle of the X-ray beam 3).
The range may be determined by: In this case, a known half-scan reconstruction is performed. Since data can be collected in a narrower extraction range, the helical pitch can be relatively increased, and the imaging efficiency can be improved.

A scanogram image can be taken by applying the above-described slice extraction in the spiral scanning. The scanogram image is a kind of transmission image, but is a transmission image with a parallel beam in the rotation axis direction. In this case, the rotation is not performed, and only the tomographic plane feed is performed, and the tracking slice data is extracted within the extraction range for each tracking slice line, and this is averaged for each tracking slice line.
A scanogram image is obtained by arranging in the order of the target tomographic plane.

Since the scanogram image becomes a transmission image in the direction exactly along the target tomographic plane, a plurality of target tomographic plane positions to be photographed on the scanogram image displayed on the display unit 9 are designated and automatically set. The position can be accurately specified when performing a so-called planned scan of the section.
When capturing the scanogram image, the data processing unit 8 stores the tomographic plane feed position when each target tomographic plane matches the imaging tomographic plane, so that the designated target tomographic plane can be accurately captured. . Shooting is fixed at each feed position and 1
Images are taken one by one, but if they are densely specified, spiral scanning can also be used.

In the case of taking a scanogram image,
It is preferable that the distance of the target tomographic plane and Sp be set to a length corresponding to one pixel of the transmission image, since the vertical and horizontal magnifications of the created scanogram image can be matched. Also, the extraction range can be freely selected, and if it is wide, the noise of the image can be reduced. However, if it is too wide, the resolution in the direction of the rotation axis decreases.

Next, setting of a slice line will be described. In order to match the position of the target tomographic plane of the subject to the imaging tomographic plane orthogonal to the rotation axis, it is necessary to know the position of the slice line corresponding to the imaging tomographic plane in advance.

In this regard, Japanese Patent Application Laid-Open No. Hei 5-3329 discloses
No. 53, the transmission image of a flat square bar is taken, the square bar is linearly moved up and down in the direction of the rotation axis, and the upper surface of the square bar positioned parallel to the front and back with respect to the detector. An attempt is made to grasp the position of the imaging tomographic plane by finding a position where the two sides seem to match. However, since the coincidence of the two sides of the rectangular bar is visually observed, there is a problem that it is difficult to accurately determine the imaging tomographic plane.

FIG. 8 is a view showing a slice phantom 26 for setting a slice line, and FIG. 9 is a view showing a transmission image when transmission data transmitted through the slice phantom 26 is displayed on the display 9. is there. Slice phantom 2
In 6, a cylindrical member 26 has a planar gap 26 c parallel to the bottom surface, and an upper portion a and a lower portion b are bonded via a spacer 27 having good X-ray transparency. The slice phantom 26 is set on the turntable 5 so that the spacer 27 is perpendicular to the rotation axis 72.

FIG. 9A shows a state in which the plane of the spacer 27 is not at the X-ray focal point F. When the slice phantom 26 is moved up and down by the tomographic plane feed mechanism 20, when the plane of the spacer 27 is in focus on the X-ray focus F, the gap image becomes clearest as shown in FIG. The line located at the center coincides with the line (slice line) where the ideal imaging tomographic plane perpendicular to the rotation axis intersects the detection channel matrix, and the slice line setting unit determines the coordinate position of this slice line. Find and memorize. As a result, the imaging tomographic plane to be sliced exactly matches the ideal imaging tomographic plane. The stored slice lines are displayed by the slice line display unit when inspecting the object, as shown in FIG.
Displayed on the display 9 together with the scale display, the operator
The position of the subject is adjusted by the tomographic plane feed mechanism 20 so that the target tomographic plane of the subject matches this slice line.

Further, the flat visible light beam 29
6c so as to be parallel to the slice line and coincide with the slice line.
By adjusting 8, the position of the target tomographic plane of the subject can be adjusted so that the visible light 29 coincides with the line intersecting the subject. At this time, if a material such as an acrylic resin that easily transmits visible light is used for the spacer 27, the adjustment can be performed easily and accurately by adjusting the light projector 28 so that the visible light passes through the gap 26c.

Therefore, according to the slice line setting in the computed tomography apparatus according to the present embodiment, the position of the slice line corresponding to the imaging tomographic plane can be accurately grasped, and the imaging tomographic plane to be slice-extracted can be determined. It can be made to pass through the X-ray focal point F exactly perpendicular to the rotation axis 72, and a high-quality tomographic image can be obtained.

The spacer 27 has members 26a,
By using a substance having a lower X-ray transmittance than 6b,
The contrast of the gap image may be reversed.

The spacer 27 does not need to fill the entire gap 26c, but may be only partially sandwiched.

Further, the display of the slice line may be switched between ON and OFF, or may be displayed by blinking or changing the display color.

It is needless to say that the shape of the slice phantom does not need to be a cylinder, but may be, for example, a prism or a cone.

Next, calibration of the calculated distance between the focal point F and the rotation axis 72 will be described.

The distance between the focal point F and the rotating shaft 72 (hereinafter referred to as “FC”)
D) does not correspond to the distance set in the image reconstruction calculation (hereinafter referred to as “FCDC”),
There is a problem that the dimensions of the reconstructed image are shifted. For example,
As shown in FIG. 10, when the FCDC is longer than the FCD, the size of the tomographic image of the subject is enlarged in proportion thereto. In this regard, in the past, several fixed positions of the FCD were provided, objects of known dimensions were photographed in advance at each position, and the size per pixel of the tomographic image at each position was obtained. Then, it is difficult to obtain a tomographic image with good dimensional accuracy.

First, the operator sets an arbitrary FCD with the focal length changing mechanism 21 and measures the approximate FCD measured with a finger or the like.
The value is input to the data processing unit 8 as an FCDC value. Next, a tomographic image of the subject having a known size is captured. The FCD calibration unit in the data processing unit 8 compares the size of the subject with the size on the tomographic image, and calculates the magnification k of the FCDC with respect to the FCD.
Is calculated, and the value of FCDC is changed to 1 / k times. Thereafter, reconstruction is performed using this FCDC value. FCD
If the above calibration is performed each time is changed, an image with high dimensional accuracy can always be obtained.

Further, instead of using a finger, an FCD measuring unit 30 can be provided as shown in FIG. FC
The D measuring unit 30 has a configuration in which a scale plate 31 is attached to a fixed portion of the focal length changing mechanism 21, and an index 32 indicating the position of the rotary shaft 72 is attached to the tomographic plane sending mechanism 20.

In this case, the index 32 or the scale plate 31 is shifted so that the 1 / k times FCDC value obtained by photographing a subject having a known size becomes the designated value.
Calibrate. Thereafter, even if the FCD is changed, the FCD measurement unit 3
Since the correct FCD value can be read by 0, this value is
An image with good dimensional accuracy can be obtained only by inputting the data as a CDC value to the data processing unit 8.

Therefore, according to the calibration of the FCD in the computed tomography apparatus according to this embodiment, the FC
Even if D is arbitrarily changed, a tomographic image with high dimensional accuracy can be easily obtained.

Note that the FCD measurement may be directly sent to the data processing unit 8 instead of being read by the operator.

Needless to say, the present invention can be applied to a case where a detector having a one-dimensional detection channel is used.

Therefore, according to the computer tomography apparatus according to the first embodiment of the present invention, distortion correction for interpolating data to correct distortion due to a shift in the detection position, Slice extraction to interpolate data on the slice line when intersecting at an angle to the matrix, slice extraction to interpolate data in spiral scanning and scano image shooting, setting of slice line position,
By performing FCDC calibration comprehensively, a high-quality tomographic image can be obtained.

In this embodiment, the X-ray I.D.
I decided to perform detection using a TV camera.
For slice extraction, other detectors having a two-dimensional channel matrix may be used.

In the above embodiment, the processing is performed in the order of slice extraction, air correction, LOG conversion, and distortion correction. However, the order can be changed as necessary.

Next, a second embodiment of the present invention will be described.

In a conventional CT scanner, a phantom for calibrating distortion caused by a shift in detection position, calibration of a position where an imaging tomographic plane intersects a detection plane of a detector, change of tomographic plane feed and change of focal length Calibration of the angle of rotation axis installation and calibration of the distance between the focal point and the rotation axis (FCD) were performed by manually adjusting each mechanism. This is the case.

FIG. 12 is a diagram showing the configuration of a computer tomography apparatus according to the present embodiment. The feature of the computer tomography apparatus is that a slice phantom 105 for determining the position of a slice line and the position of a rotation axis is connected to a rotary table 5. And a distance measuring unit 101 that measures the amount by which the focal length changing mechanism 21 has moved the subject and transmits the measured amount to the data processing unit 78.
And a calibrating unit 100 in the data processing unit 78 for performing various types of calibration processing, and a distortion calibrating phantom 113 as shown in FIG.
4, the X-ray I.I.2 can be installed / removed in front of the X-ray, and a distortion calibration phantom is installed and distortion is calibrated. Setting of slice lines for finding the position of the intersecting line, finding the center of rotation for finding the position on the detection channel row where the path passing from the radiation source and passing through the rotation axis 72 is incident, and determining the position of the radiation source 1 and the rotation axis 72 The point is that the calibration of the calculated distance is automatically performed with high accuracy.

Here, the slice phantom 105 constitutes the turntable according to the eleventh aspect, and the calibrating unit 100 comprises the distortion calibration means, the intersection line setting means, the FCD calibration means, and the rotation center according to the twelfth aspect. The grid driving section 114 forms a phantom installation / removal section, and the distance measuring section 101 forms a distance measuring section. The same components as those in FIG. 1 are denoted by the same reference numerals.

The distortion correction phantom 113 is a plate-like substance having a low X-ray absorption and a columnar wire having a high absorption embedded at regular intervals (grid).

FIG. 18 is a diagram showing the slice phantom 105. FIG. 19 is a cross-sectional view of the slice phantom 105.
8 shows a state where 8 protrudes upward. Pin drive unit 1
09 is capable of completely storing the pin 108 therein, and the slice phantom 105 is installed such that the spacer 104 between the upper portion 105a and the lower portion 105b is perpendicular to the rotation axis 72.

In such a CT scanner, each calibration is performed by an operator selecting a menu of the data processing section 78.

First, the distortion calibration will be described. FIG.
FIG. 15 is a diagram showing a transmission image of a grid when transmission data is collected using the distortion calibration phantom 113, and FIG. 15 is a diagram showing a distortion correction curve of the grid. The calibration unit 100 calculates the i _k that vertical line-shaped transmission image crosses the slice line 116, the detection channel No. i, to be detected vertical axis position on the horizontal axis as shown in FIG. 15 Channel No.
The measurement points ( _ik , _nk ) are plotted on the coordinates of n, a distortion correction curve is approximately obtained and stored using the least square method or the like, and this distortion correction curve is appropriately stored in the distortion correction unit 14. To provide interpolation processing.

Next, setting of a slice line will be described. Slice phantom 10 by tomographic plane feed mechanism 20
5 is moved in the vertical direction. When the linear image corresponding to the spacer 104 becomes the clearest in the density distribution of the transmitted image, the calibration unit 100 sets the line located at the center of the image as a slice line. Remember.

Next, how to find the center of rotation will be described.
The pin 108 of the slice phantom 105 is protruded from the upper surface, and transmission data is collected at two rotational positions different by 180 ° with respect to the pin 108. X passing
The position of the detection channel (hereinafter referred to as “center channel”) on which the line is incident is obtained and stored. When examining the subject, the reconstruction data is calculated by shifting the transmission data so that the stored center channel position becomes the center of rotation in calculation. The collection of transmission data is 180
° The detection may be performed for a plurality of sets each including a pair of detection positions facing each other.

Next, the calibration of the FCD will be described. The calibration unit 100 photographs the slice phantom 105 at an arbitrary FCD position, and measures the FCD measured by the distance measurement unit 101.
Is used as an FCDC value to reconstruct a tomographic image. Next, the dimension of the slice phantom 105 is compared with the dimension on the tomographic image to determine the magnification k of FCDC with respect to the true FCD, and F
The value of CDC is multiplied by 1 / k to determine the true FCD value, and this FCD value is determined.
The difference between the CD value and FCDC is obtained and stored. Since then, FC
When D is changed, the FCD is measured by the distance measuring unit 101, the difference stored in the FCD is added by the calibrating unit 100, and the value of FCDC is obtained and used for reconstruction.

Therefore, according to the present embodiment, the installation / removal of the strain calibration phantom and the calibration of the strain, the setting of the slice line, the determination of the rotation center, and the calibration of the FCD are automatically performed. Manually adjust each mechanism to set up a phantom for distortion correction due to the shift of the detection position, calibrate the position where the imaging tomographic plane intersects the detection plane, calibrate the angle of rotation axis installation, and FCD calibration As a result, the same effect can be obtained, and a high-quality tomographic image can be easily obtained.

The phantom for distortion correction is shown in FIG.
As shown in (a), a concentric phantom 118 in which a substance having high X-ray absorptivity is concentric, and as shown in (b), a grid phantom having a grid 119 in which small cylinders are arranged at equal intervals on a straight line. 17 using one of the side pin phantoms 122 having one pin 121 at a position away from the center as shown in FIG.
As shown in (a), a fixed image may be taken on the rotary table 5 to take a transmission image. However, in the side pin phantom 122, a correction curve is obtained using a plurality of transmission images at different rotation positions. FIG.
As shown in FIG.
May be installed underneath.

The slice phantom may be configured such that the pin 111 is embedded in the phantom as shown in FIG. However, if the width of the gap is too small, it becomes difficult to obtain transmission data of the pin 111 and the accuracy of center calibration is reduced. Therefore, a wide gap may be provided separately from the one for tomographic plane calibration. The installation position may be below the rotation mechanism 19 as shown in FIG. 21, or a pin may be provided below the rotation mechanism 19 as shown in FIG.

In the above-described embodiment, the detection is performed using the X-ray II.2 and the television camera 7. However, other detectors having a two-dimensional detection channel matrix may be used. For calibration of distortion, determination of the center of rotation, and calibration of FCD, similar effects can be obtained even if a detector having a one-dimensional detection channel array is used.

The slice phantom does not need to be cylindrical, and may be, for example, prismatic or conical.

Next, a third embodiment of the present invention will be described.

The above-described calculation of the center of rotation for finding the position on the detection channel array at which the path starting from the focal point of the radiation source and passing through the rotation axis is incident is performed using a predetermined phantom before imaging the subject. However, it is more convenient if the rotation center can be obtained without using a predetermined phantom. In this regard, in Japanese Patent Application Laid-Open No. 8-114558, a contour of a transmission image is extracted during imaging of a subject, and a rotation center is obtained. However, this method has a problem that an error is large in a subject whose contour is not clear.

FIG. 23 is a diagram showing a configuration of a computer tomography apparatus according to the present embodiment, wherein an X-ray tube 51,
A collimator 52, an X-ray beam 53, a subject 54,
Rotation table 55, mechanism control unit 56, data collection unit 57, detector 60, display 59, data processing unit 5
8 has an air correction unit 62, a LOG conversion unit 63, a rotation center finding unit 66, and a reconstruction unit 65. The feature is that the data processing unit 58 is provided with a rotation center finding unit 66 so that the 360
The position of the center channel corresponding to the rotation center is obtained using the projection data of °. Here, the rotation center finding section 66 constitutes a rotation center finding means.

The processing in the rotation center finding section 66 will be described below. FIG. 24 is a diagram illustrating the calculation of the rotation center. The average projection data Pm is obtained by averaging the data for 360 ° for each channel n on a sinogram in which the projection data P for 360 ° are arranged with the horizontal axis as the detection channel No. n and the vertical axis as the rotation angle φ. Pm is bilaterally symmetric as described later, and the position nc of the center is determined to be the position of the center channel. The center nc can be calculated by the following equation by using the centroid finding method.

Nc = (Σ _n Pm (n) · n) / (Σ _n Pm (n)) (10) Here, if the calculation including the large value and the small value of Pm is performed, the accuracy of finding the center of gravity is reduced. Therefore, the weight may be reduced for these portions. For example, a value larger than the upper limit L2 is replaced with a constant value, and a value smaller than the lower limit L1 is replaced with 0, and then the center of gravity is calculated.

The center nc is not necessarily the average data P
m can be obtained without finding m. For example,

M (φ) = Σ _n P (φ, n) · n (11) W (φ) = Σ _n P (φ, n) (12) nc = (ΣφM (φ)) / ( ΣφW (φ)) (13) Even if nc is obtained, the result is the same as that of Expression (10).

Here, the principle that the center nc coincides with the center channel will be described. FIG. 25 is a diagram illustrating the principle of finding the rotation center. The detection channels n _L and n _R symmetrically positioned with respect to the center channel on which the path from the focal point F passing through the center C is incident as shown in FIG. Since we collect data for the same set of paths over 360 ° through a certain distance from
The projection data added by 60 ° has the same value for both. That is, the projection data Pm obtained by averaging the projection data becomes symmetrical as shown in FIG.
c corresponds to the center channel.

The method of finding the center nc using this principle is not limited to the above method of finding the center of gravity. FIG.
FIG. 9 is a diagram showing another example of finding a rotation center. FIG.
And a plurality of levels Li are selected, and for each level, two points di and u are located at symmetric positions where Li crosses Pm.
i is set as a set, and ni obtained by averaging these positions is obtained.
nc can be obtained by further averaging i for all levels Li. If there are a plurality of sets where the level Li crosses Pm, any of the sets may be selected or all the sets may be used. Further, by setting the level Li to an intermediate level with a small error with high frequency, the accuracy of the calculation of nc can be improved.

As shown in FIG. 13B, the temporary center n0
(Design values and the like) are set, Pm ′ obtained by folding Pm around this n0 is obtained, and the result of adding the absolute value of the difference between the value of Pm and the value of Pm ′ for all n is obtained. n
By changing 0, the addition result is obtained in the same manner, and n0 when the addition result becomes minimum can be set to nc. Here, since the data point of P′m is shifted from the integer position of n in order to change n0 as a real value, a difference is obtained using primary interpolation. Further, the addition may be performed only on the left or right of n0.

Further, FIG. 13C shows an example in which the rotation center calculation shown in FIG. 13B is applied when half of the subject is outside the imaging region as in a CT scanner described later. In this case, since the portion of Pm that was not in the photographing region is missing, the range where Pm overlaps Pm 'when Pm is folded back at the temporary center n0 to be Pm' is limited. Therefore, the absolute value of the difference between the value of Pm and the value of Pm 'is added only for n within this overlapping range, and n0 that minimizes the addition result is determined as nc. Here, similarly, the addition may be performed only on either the left or right of n0.

Therefore, according to the present embodiment, the position of the center of rotation can be obtained for any subject during imaging, and a predetermined phantom can be used and the mechanism can be adjusted so that the rotation axis does not shift. Can be eliminated, and a high-quality tomographic image can be easily obtained.

In the present embodiment, the calculation of the rotation center is performed after the LOG conversion. However, it can be performed before the LOG conversion.

Further, only at the time of the first photographing, the rotation center is obtained to obtain nc, and the processing time can be reduced by using this nc in the second and subsequent times.

Further, the calculation of the rotation center is performed by spiral scanning CT.
When applied to a scanner, the rotation center is obtained only for the first target tomographic plane, and this is used for the subsequent tomographic plane, so that the processing time can be reduced. However,
This can be applied only when the rotation axis does not shift for each tomographic plane due to the tomographic plane feeding.

Next, a fourth embodiment of the present invention will be described.

Conventionally, a microfocus X-ray tube having an electron beam focusing coil or electrode capable of finely adjusting the focus, or an X-ray fluorescence intensifier having an electron beam focusing coil or electrode and an optical lens are provided. In a CT scanner using a TV camera, the focus is adjusted by visually observing a transmission chart with a resolution chart or a pinhole as a subject, and the adjustment depends on the skill of the adjuster or it takes time and effort. In some cases, the image quality deteriorates due to reduced frequency and shooting in an inappropriate state.

FIG. 27 is a diagram showing the configuration of a computed tomography apparatus according to the present embodiment,
A microfocus X-ray tube 14 that converges an electron beam 137 output from the electron beam 137 with a focusing coil 138 and collides the electron beam 137 with an anode target 139 to generate X-rays
6, an X-ray focus adjustment unit 133 that adjusts the current of the focusing coil 138 to adjust the size of the focal point F on the anode target 139, and the detected X-rays are converted by the X-ray / electron beam conversion film 140. Electron beam 141 to the focusing electrode 1
X-ray I.I. converged on the I.I. output surface 143 by the I.I.
147, an I.I. focus adjuster 134 for adjusting the voltage of the focusing electrode 142 to adjust the focus of the transmitted image projected on the I.I. output surface 143, and contacting the X-ray I.I.147. A television camera 148 that is installed in a state where the transmission image on the II output surface 143 is projected onto the CCD sensor 145 via the optical lens 144, an optical lens 144 from the II output surface 143, and the CCD sensor 145. And a MTF (Modulation Transfer Function) which determines a signal value and a contrast value from a transmitted image signal transmitted from the TV camera 148, and a camera focus adjustment unit 135 for adjusting the distance between
The configuration includes a data processing unit 128 having a focus evaluation unit 130 for calculating a curve, a display 9 for displaying an operation result such as an MTF curve, and a phantom 131 for focus evaluation. The focus evaluation phantom 131 is shown in FIG.
As shown in the figure, several types of rectangular patterns in which lead or the like that is difficult to transmit X-rays are repeatedly arranged with a constant width and spacing are provided on a plate-like substance.

Here, the focus evaluation section 130 constitutes a focus evaluation section, the X-ray focus adjustment section 133 constitutes a radiation source focus adjustment section, and the I.I.
4 and the camera focus adjustment unit 135 constitute a detector focus adjustment unit. In addition, the same components as those in FIG. 1 are denoted by the same reference numerals.

The operator sets the focus evaluation phantom 131 at the position A in FIG. 27 and collects transmission data from one direction. The transmission data is sent to the data processing unit 128, and the focus evaluation unit 130 determines the ratio of the contrast value C to the signal value S (hereinafter, referred to as “C / S value”) from the repeated pattern of the intensity distribution as shown in FIG. ) Is calculated, an MTF curve is calculated with the C / S value on the vertical axis and the spatial frequency f (1 / mm) on the horizontal axis, and the display 9 displays the MTF curve and the C / S value. The operator determines the C /
The focus of the X-ray II.147 is adjusted so that the S value is maximized.
I. Adjust using the focus adjustment unit 134. Since the MTF curve changes at any time in real time, the focus can be easily adjusted. Next, the focus of the television camera 148 is similarly adjusted using the camera focus adjustment unit 135. By repeating this alternately, X-ray II.
And the television camera 148 can be precisely focused. Next, the focus evaluation phantom is shown in Fig. 2.
7, the focal point of the X-ray tube 146 is adjusted so as to maximize the C / S value of the MTF curve by using a focal length changing mechanism (not shown).

Therefore, according to the present embodiment, each focus of the X-ray tube, the X-ray fluorescence intensifier, and the television camera can be accurately adjusted without depending on the skill of the operator, and high quality can be achieved. A simple tomographic image can be easily obtained. Also, MTF
By recording the curve, it is possible to know the aging of the CT scanner.

[0139] In this embodiment, the TV camera may use an image pickup tube instead of the CCD sensor. Further, the X-ray II is transmitted through the fluorescent plate and the light II (with the lens).
Can be replaced by The entire detector may be a fluorescent screen and a camera, or an X-ray image pickup tube having an X-ray / electron beam conversion film attached to the input surface of the image pickup tube. In any case, the focus can be accurately adjusted in the same manner.

Next, a fifth embodiment of the present invention will be described.

Conventionally, along a path passing through a portion of the subject which has a high X-ray absorptivity, there is a problem that the detected X-ray amount is reduced, so that the noise is relatively increased and the image quality is deteriorated.

FIG. 29 is a diagram showing a configuration of a computer tomography apparatus according to the present embodiment. The feature thereof is that a constant addition section 91 is provided after the air correction section 62 and is substantially constant with respect to transmission data. That is, the values are added. Here, the scanning mechanism 64 is a mechanism for performing TR or RR scanning of the subject 61, and the constant adding unit 91 constitutes a noise reduction unit according to claim 16. The same components as those in FIG. 23 are denoted by the same reference numerals.

FIG. 30 is a diagram showing the flow of the process for adding a constant. This is an example of a case where a cylindrical subject 61 is imaged. The transmission data d (n) collected by the data collecting unit 57 is based on the central part of the subject that absorbs a large amount of X-rays. In this case, the detected amount decreases and noise relatively increases. For this reason, as shown in FIG.
The noise at the center of the projection data after the G conversion is amplified.

Therefore, in the present embodiment, the transmission data h (n), which has been gain-corrected by the equation (1) in the air correction section 62, is added to the following equation by the constant addition section 86:

The constant h0 is added by the following equation: h ′ (n) = h (n) + h0 (14) For example, if h0 = 0.05, as shown in the figure, noise at the center of the projection data P ′ (n) obtained by logarithmically converting h ′ (n) by the LOG converter 63 is compressed. , H0, the compression is further applied. h0 is, 4 · h _NEP when the signal level of the signal noise ratio is 1 and the h _NEP
When set in the range of ~ 20 · h _NEP , the effect of noise reduction is great.

Therefore, according to the present embodiment, by adding a constant value to transmission data before logarithmic conversion, the smaller the detected X-ray dose, the greater the noise can be compressed, and the lower the spatial resolution. A tomographic image in which noise is reduced can be obtained without causing noise.

In the present embodiment, the constant value h0 is added by the constant adder 91. However, it is not strictly required that the constant value h0 be added. If not, the same effect can be obtained.

Further, the addition of the constant may be performed before the air correction, and the detector 60 or the data collection unit 57 is configured to have the noise reduction means according to claim 17 so as to increase the transmission data by a substantially constant value. Even so, a similar effect can be obtained.

Next, a sixth embodiment of the present invention will be described.

FIG. 31 is a diagram showing a configuration of a computer tomography apparatus according to the present embodiment. The feature thereof is that a noise compression section 92 is provided after the LOG conversion section 63,
That is, a function of compressing noise is applied. Here, the noise compression section 92 forms a noise compression unit. The same components as those in FIG. 29 are denoted by the same reference numerals.

FIG. 32 is a diagram showing a flow of processing for performing noise compression. The projection data P (n) obtained by logarithmic conversion in the LOG conversion unit 63 is subjected to a noise compression function in the noise compression unit 92 to obtain the following equation:

P ′ (n) = LOG (1 / (EXP (−P (n)) + h0)) (15) This noise compression function is a monotonically increasing function in which the slope becomes gentler with an increase in the input value. The larger the value of the projection data, that is, the smaller the detected X-ray dose, the greater the compression. ing. Also, the inclination angle can be adjusted by the value of h0,
Increasing h0 reduces the slope.

Therefore, according to the present embodiment, by applying the above-mentioned noise compression function to the projection data obtained by logarithmic transformation, a tomographic image in which noise has been reduced without lowering the spatial resolution can be obtained. Obtainable.

Note that the noise compression function is not limited to the equation (15), and the same effect can be obtained as long as the function has a shape similar to this.

For example, the following equation:

P ′ = P / [｛(1 / (m−1)) ＾ (1 / m) * P / P0｝ ＾ m + 1] [if P <P0] = P0 * (m−1) / m [if P> = P0] (16) Here, the parameter m determines the curvature of compression with a real number exceeding 1, and the parameter P0 determines the saturation point with a real number.

The noise reduction or noise compression according to the fifth and sixth embodiments can be performed simultaneously with the logarithmic conversion. In this case,

P ′ (n) = LOG (1 / (h (n) + h0)) (17) When h (n) is input, the function on the right side becomes h
When (n) is large, it matches the LOG function, and as it becomes smaller, it is compressed to a smaller value than the LOG function. If the function on the right side is created in a table including input values and output values, noise compression can be performed simultaneously with LOG conversion. The LOG conversion including such noise compression is not limited to the expression (17), and a similar effect can be obtained as long as the function has a shape similar to this.

Next, a seventh embodiment of the present invention will be described.

The transmission data changes when the detection time changes due to the influence of the intensity variation of the radiation. In this regard, conventionally, a dedicated comparison detector has been installed at a position where X-rays that do not pass through the subject can be detected, the intensity fluctuation of the X-rays is measured, and the transmission data is corrected.

FIG. 33 is a diagram showing the configuration of a computer tomography apparatus according to the present embodiment. The feature of the apparatus is that after processing by the air correction section 62, the X-rays outside the subject are detected by the air area identification section 93. An area on the transmission data (hereinafter, referred to as an “air area”) corresponding to the area in which is detected is determined, an average value of the X-ray dose in the air area is obtained,
In step 4, the reciprocal of the average value is multiplied by the entire transmission data to correct the intensity variation of the radiation source.
Here, the air region identification unit 93 constitutes a data identification unit, and the REF correction unit 94 constitutes an intensity fluctuation correction unit. The same components as those in FIG. 29 are denoted by the same reference numerals.

FIG. 34 is a diagram showing identification of an air region and correction of intensity fluctuation. First, in the air region identification unit 93,
As shown in FIG. 11A, the transmission data h obtained at the same time after the processing in the air correction unit 62 has a value of 1 ± ε.
(Ε is a very small value)
(The range of n represented by n <n1 and n2 <n) is obtained, and this region is slightly reduced to obtain an air region A ′ (n <n1 ′, n2 ′).
<Range of n represented by n) is obtained, and the average value k of the transmission data h in this region is obtained, and 1 / k is set as a correction multiple m.

Next, when the transmission data h is multiplied by the correction multiple m in the REF correction unit 94, the average value of the transmission data in the air region becomes 1, and the fluctuation due to the X-ray intensity fluctuation is corrected. Subsequently, projection data P is obtained through logarithmic conversion in the LOG conversion unit 63, and image reconstruction is performed using 360 ° projection data P in the reconstruction unit 65.

Therefore, according to the present embodiment, it is possible to correct the fluctuation due to the influence of the radiation intensity fluctuation without using a dedicated comparison detector, and to obtain a high-quality tomographic image.

In the present embodiment, the air region identification and the REF correction are performed before the LOG conversion, but may be performed after the LOG conversion. In this case, as shown in FIG. 34B, the provisional air area A in which the value of the projection data P after the LOG conversion is in the range of 0 ± ε.
Is obtained, the average value Pa in the air region A ′ obtained by reducing this region is obtained, and −Pa is added to the projection data, which is mathematically equivalent.

Next, an eighth embodiment of the present invention will be described.

Conventionally, the detection channel array is arranged so as to be shifted by 1/4 channel with respect to the path of the radiation passing through the rotation axis starting from the focal point of the radiation source. The X-rays on the path shifted by two channels are detected, and the projection data facing 180 ° are combined to form one projection data, and the projection data for 180 ° is apparently obtained every 1/2 channel. There are known CT scanners that improve the resolution in such a manner. Since the CT scanner performs the reconstruction process after combining the facing data, the reconstruction is started after collecting the data for 360 °, so that it takes a long time to form a tomographic image. To solve this, Japanese Patent Laid-Open No. Sho 6
In Japanese Unexamined Patent Application Publication No. 2-231626, half of 360 ° projection data is inserted by inserting 0 values between data without combining projection data facing 180 °.
A data value is obtained for each channel so that the resolution is high and the reconstruction can be performed in parallel with the data acquisition. However, in this method, there is a problem that a ring-shaped artifact (false image) may be generated on a tomographic image because a change between adjacent data becomes severe due to insertion of a 0 value.

FIG. 35A is a diagram showing a configuration of a computer tomography apparatus according to the present embodiment, and FIG. 35B is an enlarged view of a central portion A of the detector 60. Its features include a detection channel 81 of the detector 60.
Are shifted by 1/4 channel with respect to a center line 71 connecting the focal point F and the rotation axis, and a data insertion unit 95 is provided after the LOG conversion unit 63 in the data processing unit 84 to provide projection data. In this case, an interpolated value for interpolating the values of two adjacent points is inserted between these data.
Here, the data insertion unit 95 constitutes a data insertion unit. The same components as those in FIG. 23 are denoted by the same reference numerals.

FIG. 36 is a diagram showing the flow of the data insertion process. With respect to the detected projection data (shown by a black circle in the figure), an interpolation value (shown by a white circle) for interpolating two adjacent points is obtained by linear interpolation, and this interpolation value is inserted between the data. The number of samples of the projection data after inserting the interpolation value is doubled.

Therefore, according to the present embodiment, in a CT scanner in which detectors are arranged shifted by 1/4 channel, an interpolation value for interpolating two adjacent points is inserted between data to reduce the number of samples of projection data. Since the resolution is substantially doubled, the resolution of the reconstruction processing can be doubled, so that the resolution of the tomographic image is improved. Further, since there is no unreasonableness between adjacent data, a tomographic image having less ring-shaped artifacts can be obtained. Obtainable. In addition, since reconstruction processing can be performed in parallel with data collection, a tomographic image can be obtained in a short time.

In this embodiment, the data insertion is performed after the LOG conversion. However, the same effect can be obtained by performing the data insertion before the LOG conversion.

Next, a ninth embodiment of the present invention will be described.

Conventionally, there has been known a CT scanner which is capable of imaging a large object by detecting projection data in a state in which one side of the object is out of an imaging region by shifting the position of a rotation axis. (JP-A-58-116342).
However, since the projection data rapidly changes between the photographed area where data was actually collected and the non-photographed area where data was not collected, if image reconstruction is performed using such projection data, a ring-shaped image will be formed. There is a problem that artifacts tend to occur.

FIG. 37 is a diagram showing a configuration of a computer tomography apparatus according to the present embodiment. The feature of the apparatus is that a rotation axis of a rotary table 150 on which a large subject 54 is mounted is set outside in an imaging area. In addition, a window function multiplying unit 96 is provided in the data processing unit 85 so that a predetermined window function acts on the projection data as a filter to obtain a reconstructed image. Here, the window function multiplication unit 96
Constitutes a window function calculating means. The same components as those in FIG. 23 are denoted by the same reference numerals.

FIG. 38 is a diagram showing the flow of the process of multiplying by the window function. The window function w (n) has a horizontal axis representing the detection channel N
o.n, the path connecting the focal point F and the rotation axis has an inclined portion varying from 0 to 1 around the incident detection channel (nc), and the value at nc is 0.5. In the window function multiplication, the projection data P
By multiplying (n) by the window function w (n), new projection data P ′ (n) is obtained. This P ′ (n) is 360 °
Tomographic images can be obtained by performing reconstruction processing by the FBP method.

Therefore, according to the present embodiment, when detecting projection data in a state where one side of a large subject is out of the imaging region, a window function is applied so that the projection data is calculated from the detected value of the imaging region. By smoothly changing to the zero value in the non-imaging area, a ring-shaped tomographic image with few artifacts can be obtained.

The window function has a value at nc of 0.5.
Having an inclined portion that changes from 0 to 1 in a shape whose inclination is symmetric on the left and right of nc, that is, the following condition:

## EQU16 ## As long as w (nc + n) + w (nc-n) = 1 (18), another function similar to this may be used. For example, the same effect can be obtained by using a window function in which a slope portion is replaced with a curve as shown in FIG.

Next, a tenth embodiment of the present invention will be described.

Conventionally, when image reconstruction is performed using detectors having detection channels arranged at equal intervals, projection data detected at equal intervals are converted into projection data at equal angular intervals, and then image reconstruction is performed. Therefore, there is a problem that image quality is deteriorated due to an error in the conversion process.

FIG. 40 is a diagram showing a configuration of a computer tomography apparatus according to the present embodiment. The feature thereof is that X-ray beams 3 generated from an X-ray tube 1 are arranged at equal intervals on a plane. Detected by the planar solid state detector 151 having the detected detection channel matrix, and the slice extraction unit 11 and the air correction unit 1
2. The projection data P is obtained through the LOG conversion unit 13, and the filter unit 98 applies a filter performed by the normal FBP method. Then, the linear data back projection unit 99 converts the equally spaced projection data to the equally spaced projection data. That is, a reconstructed image is calculated without converting the image into. Here, the straight line data back projection unit 99 constitutes back projection means. In addition, the same components as those in FIG. 1 are denoted by the same reference numerals.

The processing in the straight line data back projection unit 99 will be described below. FIG. 41 is a diagram showing the detection of equal intervals, and FIG. FIG. 41 shows the focus F when the rotation angle is φ in the coordinate axes x and y fixed to the subject whose rotation axis is the origin.
Shows the relationship between the coordinates (x _F , y _F ) and the measurement line 45 indicating the detection position of the planar solid state detector 151. Each position at equal intervals on the x-axis is represented by a centering number k. In FIG. 42, first, in step 300, the projection data P for 360 ° is divided into four quotas of 90 ° each having −45 ° to 45 ° as one unit.
In a calculation loop for the projection data for °, the coordinates (x _F , y _F ) of the focal point F are calculated at step 320 for each fixed angle φ, and the projections obtained at equal intervals on the measurement line 45 at step 330 The data P is converted into a data string at equal intervals on the x-axis (hereinafter referred to as “centering”). In the centering in step 330, the operation loop for each centering number k is entered in step 340, and the centering data Pc (k) is obtained by the following equation.

[0178]

X ₀ = (k−kc) · cp (19) x ₁ = x ₀ · cos φ (20) L = ((x _F −x ₀ ) ² + y _F ² ) ^1/2 (21) ) n '= nc + (1 / d) · FDD · x 1 / (L 2 -x 1 2) 1/2 ... (22) Pc (k) = P (n') ... (23) (n ' is a real number Therefore, P (n ') is obtained by linear interpolation. However, cp: centering pitch, kc: center k, n
It is assumed that c: center channel, d: channel pitch, FDD: distance between focus and measurement line.

That is, the position of n ′ on the measurement line 45 corresponding to the centering number k is geometrically obtained, and the projection data P (n ′) at this position is obtained by linear interpolation to obtain the projection data Pc at the centering number k. (K). In step 370, when the Pc (k) thus obtained is back-projected to each pixel of the reconstructed image,
A method known in JP-A-54-152490 or the like is used. Subsequently, in step 380, the angle φ is incremented, and the process returns to step 310 to repeat the above processing. After the above processing for the data for 90 ° is completed, the image is rotated by 90 ° in step 390, and
At 00, the process proceeds to the next quota, and the same processing from step 300 is repeated. Back projection is performed on the four quotas to obtain a final tomographic image.

Therefore, according to the present embodiment, the projection data detected by the detection channel matrices arranged at equal intervals can be back-projected without being converted to equal angular intervals, and are converted to equal angular intervals. It is possible to obtain a high-quality tomographic image having no error at the time.

In each of the above embodiments, the rotary table on which the subject is placed is rotated.
The X-ray tube and the detector may be rotated.

Further, in each of the above embodiments, the radiation emitted from the radiation source is X-ray, but other transmissive radiation may be used.

Further, in the third to ninth embodiments, the detector is a one-dimensional detector, but may be a two-dimensional detector. Not limited to TR scanning, other scanning methods, for example, an SR (Stationary-Rotat) in which a ring-shaped detector surrounding the subject is fixedly installed and the X-ray tube is rotated on the inner or outer circumference thereof
e) Needless to say, the method may be adopted.

[0184]

As described above, according to the first aspect of the present invention, when a cosine function is applied to the data sequence, the data sequence is interpolated. As in linear interpolation, an interpolated value obtained using two adjacent data values in which noise may be canceled and an interpolated value using one data value in which noise remains as they are are shown in a data string. Non-uniformity of noise due to coexistence can be prevented, and a high-quality tomographic image can be obtained.

According to the second aspect of the present invention, a cosine function is applied to the data matrix to obtain an interpolated value corresponding to a line where a tomographic plane intersects a detector, thereby obtaining an interpolated value. Even when the plane intersects the data matrix of the detection plane at an angle, the data sequence for the imaging tomographic plane originally required for the calculation of the reconstruction can be accurately obtained, and thus high quality A tomographic image can be obtained.

According to the third aspect of the present invention, a straight line parallel to the line where the imaging tomographic plane intersects the detection plane within the extraction range follows the movement of the target tomographic plane of the subject, and By obtaining an interpolation value corresponding to the above, in tomographic imaging by spiral scanning, a data sequence for the target tomographic plane of the subject can be obtained for all angles of the rotation, and thus a high-quality tomographic image can be obtained. .

According to the fourth aspect of the present invention, a straight line parallel to a line where the imaging tomographic plane intersects with the detection plane within the extraction range is made to follow the movement of the target tomographic plane of the subject. By obtaining the interpolation value corresponding to the above, a high-quality scanogram image along the target tomographic plane of the subject can be obtained in scanogram imaging by scanning the displacement.

According to the fifth aspect of the present invention, a cosine function is applied to the data matrix by applying a trapezoidal function or a function in which a trapezoidal slope is replaced by a cosine curve. The same effect as in the case can be obtained, so that a high-quality tomographic image or scanogram image can be obtained.

According to the sixth aspect of the present invention, the position of the gap perpendicular to the rotation axis in the direction of the rotation axis is adjusted, and the position of the center line of the gap when the gap on the transmission image becomes clearest. , The position of the intersection line where the imaging tomographic plane intersects the detection plane can be grasped, and the imaging tomographic plane can be made exactly perpendicular to the rotation axis, thereby obtaining a high-quality tomographic image. Can be.

According to the seventh aspect of the present invention, by projecting a planar visible light beam on the intersection line, a line where the imaging tomographic plane intersects the subject is indicated by the visible light beam. Thus, the target tomographic plane of the subject is easily adjusted to the imaging tomographic plane.

According to the present invention, by displaying the transmission image of the subject and the intersection line, the target tomographic plane of the subject can be easily adjusted to the imaging tomographic plane by referring to the display. are doing.

According to the ninth aspect of the present invention, the dimensions of the subject are compared with the dimensions of the subject on a tomographic image obtained by imaging the subject, and the radiation source used for calculation is determined. By correcting the distance from the rotation axis, a tomographic image with high dimensional accuracy can be obtained at an arbitrary position from the radiation source.

According to the tenth aspect of the present invention, the position on the data sequence corresponding to the center of the symmetrical data obtained by adding or averaging the data sequence for 360 ° for any subject is obtained. As a result, it is possible to calculate the image reconstruction by grasping the position on the data sequence corresponding to the position on the detector where the radiation passing through the rotation axis is incident, and mechanically adjust the displacement of the rotation axis position In addition, it is not necessary to perform the operation, and a high-quality tomographic image can be easily obtained.

According to the present invention, a phantom for obtaining the intersection line, a phantom for obtaining the position of the rotation axis, and the predetermined phantom for distortion correction are previously set or installed.・ Be able to separate and measure the distance from the radiation source to the axis of rotation to determine the position of the intersection line and the axis of rotation, and to calibrate the distortion or calculate the distance between the source and the axis of rotation. By setting automatically, installation of a dedicated phantom, calibration of the position where the tomographic plane intersects the detection surface, calibration of the angle of rotation axis installation, calibration of the distance from the radiation source to the rotation axis Can be avoided by manually adjusting each mechanism, and a high-quality tomographic image can be easily obtained.

According to the thirteenth aspect of the present invention, by fixing one of the phantoms to the turntable,
While avoiding the trouble of manually installing a phantom, a phantom having a pattern of the concentric circles or a phantom having the grid is a transmission image from one direction, and a phantom having the pins is a transmission image from a plurality of directions. By using this, it is possible to calibrate the distortion caused by the displacement of the radiation detection position.

According to the present invention, the focus state is evaluated from the ratio between the contrast value and the signal value of the transmitted image of the focus evaluation phantom,
The focus of the radiation source or the focus of the detector can be adjusted so that this ratio is maximized, so that these focuses can be easily adjusted without depending on the skill of the operator. Therefore, a high-quality tomographic image can be easily obtained.

According to the present invention, the logarithmic conversion is performed after adding or increasing a substantially constant value to the detected data, or the data sequence after the logarithmic conversion is performed. By applying the compression function to, or by applying the compression function simultaneously with the logarithmic conversion, to reduce noise that is relatively increased due to a small amount of radiation detection Therefore, a high-quality tomographic image can be obtained.

According to the twentieth aspect of the present invention, a dedicated comparison detector for measuring the intensity variation of the radiation is used by correcting the influence of the intensity variation of the radiation using the value of the data in the area. A high-quality tomographic image can be obtained without any problem.

According to the twenty-first aspect of the present invention, the interpolation value is interpolated between each data of the data string obtained by detecting the radiation transmitted through the subject with the detection channel string shifted by １ ／ channel. , The spatial resolution can be improved by substantially doubling the number of data, and a high-quality tomographic image can be obtained.

According to the twenty-second aspect of the present invention, by applying the window function to the data sequence, when imaging a large object, one side of the object is imaged by shifting the rotation axis outward. A data string obtained by detecting radiation transmitted through the subject while being out of the region can be smoothly changed from a detection value in the imaging region to a zero value in the non-imaging region, and a ring shape on the tomographic image can be obtained. A high-quality tomographic image in which the false image of is reduced can be obtained.

According to the twenty-third aspect of the present invention, a data sequence obtained at equal intervals in a straight line is defined by using a rotation axis as an origin.
Since the image is reconstructed by converting the image data into dimensional coordinates, it is not necessary to convert the data sequence into data sequences at equal angular intervals, and an error occurs when the data sequence is converted into data sequences at equal angular intervals. Can be prevented, and a high-quality tomographic image can be obtained.

[Brief description of the drawings]

FIG. 1 is a diagram showing a configuration of a computer tomography apparatus according to a first embodiment of the present invention.

FIG. 2 is a diagram illustrating a distortion correction curve.

FIG. 3 is a diagram illustrating an example of an interpolation function in distortion correction.

FIG. 4 is a diagram showing a screen of a display device 9 when transmission data is displayed as it is as a transmission image.

FIG. 5 is a diagram illustrating an example of an interpolation function in slice extraction.

FIG. 6 is a diagram showing a screen of a display device 9 when a transmission image detected in spiral scanning is displayed.

FIG. 7 is a diagram showing slice extraction in spiral scanning.

FIG. 8 is a diagram showing a slice phantom 26 for setting a slice line.

FIG. 9 is a view showing a transmission image when transmission data transmitted through a slice phantom 26 is displayed on a display 9;

FIG. 10 is a diagram showing a positional relationship between FCD and FCDC.

FIG. 11 is a diagram showing an FCD measuring unit 30.

FIG. 12 is a diagram illustrating a configuration of a computer tomography apparatus according to a second embodiment of the present invention.

FIG. 13 shows that the distortion correction phantom 113 is irradiated with an X-ray II.
It is a figure showing the state where it was installed in the front of.

FIG. 14 is a diagram showing a transmission image of a grid.

FIG. 15 is a diagram showing a distortion correction curve of a grid.

FIG. 16 is a diagram showing another example of the distortion calibration phantom.

FIG. 17 is a diagram illustrating an example of an installation position of a distortion calibration phantom.

FIG. 18 is a diagram showing a slice phantom 105.

FIG. 19 is a diagram showing a cross-sectional view of a slice phantom 105.

FIG. 20 is a diagram showing another example of a slice phantom phantom.

FIG. 21 is a diagram showing another example of the installation position of the slice phantom.

FIG. 22 is a diagram illustrating another example of the installation positions of the pins.

FIG. 23 is a diagram illustrating a configuration of a computer tomography apparatus according to a third embodiment of the present invention.

FIG. 24 is a diagram showing the calculation of the center of rotation.

FIG. 25 is a diagram showing the principle of finding the center of rotation.

FIG. 26 is a diagram showing another example of finding a rotation center.

FIG. 27 is a diagram illustrating a configuration of a computer tomography apparatus according to a fourth embodiment of the present invention.

FIG. 28 is a diagram showing a focus evaluation phantom, a C / S value, and an MTF curve.

FIG. 29 is a diagram illustrating a configuration of a computer tomography apparatus according to a fifth embodiment of the present invention.

FIG. 30 is a diagram showing a flow of processing for performing constant addition.

FIG. 31 is a diagram showing a configuration of a computer tomography apparatus according to a sixth embodiment of the present invention.

FIG. 32 is a diagram showing a flow of processing for performing noise compression.

FIG. 33 is a diagram illustrating a configuration of a computer tomography apparatus according to a seventh embodiment of the present invention.

FIG. 34 is a diagram illustrating identification of an air region and correction of intensity fluctuation.

FIG. 35 is a diagram illustrating a configuration of a computed tomography apparatus according to an eighth embodiment of the present invention.

FIG. 36 is a diagram showing a flow of processing of data insertion.

FIG. 37 is a diagram showing a configuration of a computed tomography apparatus according to a ninth embodiment of the present invention.

FIG. 38 is a diagram showing a flow of a process of multiplying by a window function.

FIG. 39 is a diagram showing another example of the window function.

FIG. 40 is a diagram showing a configuration of a computer tomography apparatus according to a tenth embodiment of the present invention.

FIG. 41 is a diagram illustrating detection of equal intervals.

FIG. 42 is a diagram showing the flow of processing of straight line data back projection.

FIG. 43 is a diagram showing a configuration of a conventional computed tomography apparatus.

[Explanation of symbols]

DESCRIPTION OF SYMBOLS 1 ... X-ray tube 2 ... X-ray II 3 ... X-ray beam 4 ... Subject 5 ... Rotary table 6 ... Mechanism control part 7 ... TV camera 8 ... Data processing part 9 ... Display 10 ... Photoelectric surface 11 ... Slice extraction unit 12 ... Air correction unit 13 ... LOG conversion unit 14 ... Distortion correction unit 15 ... Reconstruction unit 19 ... Rotating mechanism 20 ... Tomographic plane feeding mechanism 21 ... Focal length changing mechanism 23 ... Slice line 24 ... Following slice line 26 ... Slice phantom, 26a ... upper part, b ... lower part,
c ... Gap 27 ... Spacer 28 ... Light projector 29 ... Visible light 30 ... FCD measuring part 31 ... Scale plate 32 ... Index 45 ... Measurement line 51 ... X-ray tube 52 ... Collimator 53 ... X-ray beam 54 ... Subject 55 ... 56 ... Mechanism control unit 57 ... Data collection unit 58 ... Data processing unit 59 ... Display unit 60 ... Detector 61 ... Subject 62 ... Air correction unit 63 ... LOG conversion unit 64 ... Scanning mechanism unit 65 ... Reconstruction unit 66 ... Rotation Center detecting section 67 ... Rotating mechanism 68 ... Tomographic plane feeding mechanism 69 ... Focal length changing mechanism 71 ... Center line 72 ... Rotating axis 73 ... Imaging tomographic plane 78 ... Data processing section 80,82,83,84,85,86 ... Data Processing unit 81 Detection channel 90 Reconstruction unit 91 Constant addition unit 92 Noise compression unit 93 Air region identification unit 94 REF correction unit 95 Data insertion unit 96 Function hook 97 ... center line 98 ... filter section 99 ... linear data backprojection unit 100 ... correction unit 101 ... distance measuring section 104 ... spacer 105 ... slice phantom, 105a ... upper 105
b: lower part, 105c: gap 106: spacer 108: pin 109: pin driver 110: slice phantom, 110a: upper part, 110
b ... lower part, 110c ... gap 111 ... pin 112 ... pin 113 ... distortion correction phantom 114 ... grid driver 115 ... transmission image 116 ... slice line 118 ... concentric phantom 119 ... grid 120 ... grid phantom 121 ... pin 122 ... side pin Phantom 128 Data processing unit 130 Focus evaluation unit 131 Focus evaluation phantom 133 X-ray focus adjustment unit 134 II focus adjustment unit 135 Camera focus adjustment unit 136 Filament 137 Electron beam 138 Focusing coil 139 ... Target 140 ... X-ray / electron conversion film 141 ... Electron beam 142 ... Converging electrode 143 ... I.I. Output surface 144 ... Optical lens 145 ... CCD sensor 146 ... Micro focus X-ray tube 147 ... X-ray I.I.148 ... TV turtle La 150 rotary table 151 flat solid detector 201 X-ray tube 202 collimator 203 X-ray beam 204 subject 205 scanning mechanism 206 detector 207 data collection unit 208 data processing unit 209 display 212: air correction unit 213: LOG conversion unit 215: reconstruction unit

Continued on the front page (72) Inventor Shoji Fujii 2-24-24 Harumi-cho, Fuchu-shi, Tokyo Toshiba FA System Engineering Co., Ltd. (72) Inventor Teruo Yamamoto 2- 24-1 Harumi-cho, Fuchu-shi, Tokyo Toshiba FA System Engineering Co., Ltd. F-term (reference) 2G001 AA01 AA07 BA11 CA01 DA09 FA06 FA08 GA05 HA08 HA12 HA14 JA02 JA06 JA08 JA11 LA01 PA11 SA10

Claims (23)

[Claims] 1. A relative rotation is given to a pair of a radiation source and a detector arranged opposite to each other and a subject located therebetween, and the detector detects radiation from the radiation source that has passed through the subject. A computer tomography apparatus for obtaining a tomographic image of a subject using the obtained data sequence, comprising: a data interpolating unit that obtains an interpolated value between data of the data sequence by applying a cosine function to the data sequence. A computed tomography apparatus characterized by the above-mentioned. 2. A method according to claim 1, further comprising: applying a relative rotation to the set of the radiation source and the detector disposed opposite to each other and the subject positioned therebetween, and transmitting the radiation from the radiation source transmitted through the subject to the detection surface of the detector. In a computed tomography apparatus for obtaining a tomographic image of a subject using a data matrix obtained by two-dimensionally detecting, a cosine function is applied to the data matrix to pass through a radiation source and to be orthogonal to the axis of rotation. A computer tomography apparatus comprising: a data interpolating unit that obtains an interpolated value on a line in the data matrix corresponding to an intersection line at which a tomographic plane to be intersected intersects the detection plane. 3. A radiation source which transmits a subject through relative rotation and a displacement in the direction of an axis of rotation to a pair of a radiation source and a detector arranged opposite to each other and a subject located therebetween. A computed tomography apparatus for obtaining a tomographic image of a subject by using a data matrix obtained by two-dimensionally detecting the radiation by a detection surface of a detector, wherein the tomographic image passes through a radiation source and is orthogonal to the rotation axis. Within an extraction range surrounded by two lines that are arranged substantially symmetrically and in parallel with respect to an intersection line where the plane intersects the detection plane, the movement of the target tomographic plane of the subject due to the displacement is Computed tomography having data interpolation means for following a straight line parallel to the intersection line and obtaining an interpolated value on the line in the data matrix corresponding to the straight line by applying a cosine function to the data matrix. Location. 4. A relative rotation is given to a pair of a radiation source and a detector arranged opposite to each other and a subject located therebetween, and radiation from the radiation source transmitted through the subject is detected by a detection surface of the detector. In a computer tomography apparatus for obtaining a tomographic image of a subject using a data matrix obtained by two-dimensionally detecting the tomographic image, the tomographic position is changed by giving a displacement in the axial direction of the rotation relative to the subject. An extraction area surrounded by two lines that are arranged substantially symmetrically and in parallel with respect to an intersection line where the imaging tomographic plane passing through the radiation source and orthogonal to the axis of rotation intersects the detection plane. Within, a straight line parallel to the intersection line is caused to follow the movement of the target tomographic plane of the subject due to the displacement, and an interpolated value on a line in the data matrix corresponding to the straight line is represented by a cosine of the data matrix. Let the function work Computer tomography apparatus characterized by having a Mel data interpolation means. 5. The computer tomographic method according to claim 1, wherein said data interpolation means applies a trapezoidal function or a function obtained by replacing a trapezoidal inclined portion with a cosine curve to said data matrix. Shooting equipment. 6. A relative rotation is given to a pair of a radiation source and a detector arranged opposite to each other and a subject located therebetween, and radiation from the radiation source transmitted through the subject is detected by a detection surface of the detector. A computed tomography apparatus for obtaining a tomographic image of a subject using a data matrix obtained by two-dimensional detection, comprising: a phantom having a planar gap; and setting the gap so as to be orthogonal to the axis of rotation. A relative displacement in the direction of the axis is given to the phantom, and a line in the data matrix corresponding to the center line of the gap when the transmission image obtained with respect to the planar gap of the phantom becomes clearest. And a crossing line setting means for determining the position of the computer tomography. 7. The computed tomography apparatus according to claim 2, further comprising a light projecting unit that emits a planar visible light beam at the position of the intersection line. 8. The computed tomography apparatus according to claim 2, further comprising display means for displaying a transmission image of the subject and the intersection line on a display device. 9. A relative rotation is given to a pair of a radiation source and a detector arranged opposite to each other and an object located therebetween, and the detector detects radiation from the radiation source transmitted through the object. In a computed tomography apparatus for obtaining a tomographic image of a subject using the obtained data sequence, a distance changing means for arbitrarily changing a distance (FCD) between the radiation source and the rotation axis; Determine the magnification of the dimensions of the subject on the tomographic image obtained using the dimensions of the subject,
An FCD correction unit for dividing the FCD value by the magnification and correcting the FCD value. 10. A relative rotation is given to a pair of a radiation source and a detector arranged opposite to each other and an object located therebetween, and the detector detects radiation from the radiation source transmitted through the object. In a computed tomography apparatus for obtaining a tomographic image of a subject using the obtained data sequence, the computed tomographic image corresponds to the center of left-right symmetric data obtained by adding or averaging a data sequence of 360 ° for an arbitrary subject. A computed tomography apparatus, comprising: a rotation center finding means for finding a position on a data string and using the rotation as a center of rotation. 11. A relative rotation is given to a pair of a radiation source and a detector which are arranged to face each other and a turntable on which a subject placed between them is placed, and the radiation from the radiation source which has passed through the subject is emitted. In a computed tomography apparatus for obtaining a tomographic image of a subject using a data matrix obtained by two-dimensionally detecting a detection surface of a detector, a planar shape perpendicular to the axis of rotation provided on the turntable is provided. And a crossing line setting means for determining the position of a line on the data matrix corresponding to the line at the center of the gap when the transmission image obtained for the gap is the clearest. A phantom installation / removal means that can be installed / removed beforehand, and a distortion calibration means for calibrating distortion on a tomographic image due to a shift in a detection position using a data matrix obtained for the predetermined phantom; A distance measuring means for measuring a distance from a radiation source to the axis of rotation, wherein: 12. A relative rotation is given to a pair of a radiation source and a detector which are arranged to face each other and a turntable on which a subject placed between them is placed, and the radiation from the radiation source which has passed through the subject is emitted. In a computed tomography apparatus for obtaining a tomographic image of a subject using a data matrix obtained by two-dimensionally detecting a detection surface of a detector, a planar shape perpendicular to the axis of rotation provided on the turntable is provided. And a crossing line setting means for determining the position of a line on a data matrix corresponding to the line at the center of the gap when the transmission image obtained for the gap is the clearest, provided on the turntable. Using a pin parallel to the axis of rotation, and a data string obtained for the pin, determine a position on the data string corresponding to a position on the detection surface where radiation passing through the axis of rotation is incident. In the computational rotation Computer tomography apparatus characterized by having a center of rotation Motomede means, a distance measuring means for measuring a distance to the axis of the rotation from the radiation source. 13. A relative rotation is given to a pair of a radiation source and a detector which are arranged to face each other and a turntable on which a subject placed between them is placed, and the radiation from the radiation source transmitted through the subject is given. In a computed tomography apparatus for obtaining a tomographic image of a subject using a data matrix obtained two-dimensionally detected by a detection surface of a detector, the turntable has a pattern of concentric circles about the axis of rotation. A phantom having a grid parallel to the detection surface and having a grid arranged at predetermined intervals in a straight line or a phantom having a pin parallel to the axis of rotation at a position away from the axis is fixed, Distortion correction means for correcting a distortion on a tomographic image due to a shift in a detection position using a data matrix obtained for the rotary table. Shooting equipment. 14. A radiation source which is disposed opposite to a radiation source, a detector and an object positioned between them are given relative displacement, and the radiation from the radiation source transmitted through the object is detected by the detector. In a computer tomography apparatus for obtaining a tomographic image of a subject using a data sequence, a phantom for evaluating a focus of a radiation source, and evaluating a focus state of the radiation source using a data sequence obtained for the phantom A computed tomography apparatus, comprising: a focus evaluation unit that performs a focus; and a radiation source focus adjustment unit that can adjust a focus of the radiation source based on a result of the evaluation. 15. A radiation source obtained by applying a relative displacement to a radiation source, a detector, and a subject located between the radiation source and a radiation source, which are transmitted through the subject, are detected by the detector. In a computed tomography apparatus for obtaining a tomographic image of a subject using a data sequence, a phantom for evaluating a focus of a detector, and evaluating a focus state of the detector using a data sequence obtained for the phantom A computer tomography apparatus, comprising: a focus evaluation unit that performs adjustment; and a detector focus adjustment unit that can adjust the focus of the detector based on a result of the evaluation. 16. A radiation source, a detector and a subject located between them, which are opposed to each other, are given relative displacements, and the radiation from the radiation source that has passed through the subject is detected by the detector. In a computer tomography apparatus for obtaining a tomographic image of a subject by logarithmically transforming a data sequence, a substantially constant value is added to each data of the data sequence to reduce noise after logarithmic transformation before the logarithmic transformation. A computed tomography apparatus having noise reduction means. 17. A radiation source obtained by applying a relative displacement to a radiation source, a detector, and a subject located between the radiation source, which are opposed to each other, and detecting the radiation from the radiation source transmitted through the subject by the detector. In a computed tomography apparatus that obtains a tomographic image of a subject by logarithmically converting a data sequence, when detecting radiation transmitted through the subject, the detected value is increased by a substantially constant value to reduce noise after logarithmic conversion. A computer tomography apparatus comprising a noise reduction unit for reducing noise. 18. A radiation source obtained by applying a relative displacement to a radiation source, a detector, and a subject located between the radiation source and the radiation source disposed opposite to each other, and detecting the radiation from the radiation source transmitted through the subject by the detector. In a computer tomography apparatus that obtains a tomographic image of a subject by performing logarithmic conversion and other predetermined processing on a data sequence, an output value of the data sequence obtained by the logarithmic conversion decreases with an increase in an input value. A computer tomography apparatus, comprising: noise compression means for compressing noise by applying a function to cause the noise to be reduced. 19. A radiation source obtained by applying a relative displacement to a radiation source, a detector, and a subject located between the radiation source and a radiation source that have been transmitted through the subject and detecting the radiation from the radiation source through the subject. In a computed tomography apparatus that obtains a tomographic image of a subject by performing predetermined processing on a data sequence, a logarithmic transformation is performed on the data sequence, and a function that compresses an output value as the output value of the logarithmic transformation increases. A computer tomography apparatus comprising a noise compression means for simultaneously operating 20. A radiation source obtained by applying a relative displacement to a radiation source, a detector, and a subject located between the radiation source and a radiation source, which are transmitted through the subject, are detected by the detector. A computer tomography apparatus that obtains a tomographic image of a subject using a data sequence, wherein the data identifying unit identifies data in the data sequence corresponding to radiation that has passed through an area outside the subject, and is identified by the data identifying unit. An intensity variation correction unit that corrects the influence of the intensity variation of the radiation using the obtained data. 21. A pair of a radiation source and a detector arranged opposite to each other and a subject positioned therebetween are rotated relative to each other, and radiation from the radiation source transmitted through the subject is detected by a detection channel array of the detector. A computed tomography apparatus for obtaining a tomographic image of a subject by using a data sequence obtained by the above method, wherein the detection channel sequence is shifted by 1/4 channel with respect to a path of radiation passing through the axis of rotation. A computer tomography apparatus comprising: a data insertion unit that inserts an interpolation value for interpolating between data into an obtained data sequence. 22. A relative rotation is given to a pair of a radiation source and a detector arranged facing each other and a subject located therebetween, and the detector detects radiation from the radiation source that has passed through the subject. In a computed tomography apparatus for obtaining a tomographic image of a subject using the obtained data sequence, a data sequence corresponding to a position on a detector where radiation passing through the axis of rotation is incident on the data sequence A computed tomography apparatus characterized by having a window function calculating means for applying a window function having a slope centered on a position. 23. A relative rotation is given to a pair of a radiation source and a detector arranged opposite to each other and an object located therebetween, and the radiation from the radiation source transmitted through the object is linearized in the detector. In a computed tomography apparatus for obtaining a tomographic image of a subject using a data sequence obtained by detection by a detection channel sequence arranged at equal intervals, the data sequence may be expressed on two-dimensional coordinates with the rotation axis as an origin. Computer tomography apparatus, comprising back projection means for performing back projection of image reconstruction by converting the data into equally spaced data strings.