JP2005331527A - Computerized tomography method and device - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To contribute to increase of speed or improvement of image quality in photography. <P>SOLUTION: This computerized tomography device is provided with a radiation source 1 radiating radiation beams, a radiation detector 3 detecting the radiation beams 2 from the radiation source 1 with a two-dimensional spatial resolution, rotation means 5 and 6 relatively rotating a specimen 4 inside the radiation beams 2, a slant part varying two-dimensional transparent data from multiple directions for the specimen acquired by the radiation detector 3 during rotation by the rotation means 5 and 6 so that inclination becomes 0.5 in the rotation axis position and changes 0 to 1 respectively giving inclinations symmetrical in the lateral direction crossing the rotation axis at right angles, and a reconstruction means 9 creating a three-dimensional image of the specimen by performing back projection after multiplication of a window function giving 0 on one side and 1 on the other side on the outside of the slant part. <P>COPYRIGHT: (C)2006,JPO&NCIPI

Description

  The present invention relates to a computer tomography method and apparatus among non-destructive inspection apparatuses, and more particularly to a computer tomography method and apparatus capable of contributing to high-speed imaging or high image quality.

  In recent years, high-resolution industrial computer tomography apparatuses (hereinafter referred to as CT scanners) have been manufactured for the purpose of inspecting small electronic components and the like with high resolution.

  FIG. 13 is a schematic diagram showing a system configuration example of this type of conventional high-resolution CT scanner, in which both a transmission image and a cross-sectional image are obtained.

  In FIG. 13, an X-ray tube 101 and a detector 103 that detects a cone-shaped X-ray beam 102 emitted from the X-ray tube 101 with two-dimensional spatial resolution are arranged to face each other. A transmission image of the subject 104 inside is obtained.

  The transmission image can be displayed as a moving image in real time, or can be added by the data processing unit 109 and displayed with reduced noise.

  When the rotary table 105 and the rotary lifting mechanism 106 are moved closer to or away from the X-ray tube 101 by the shift mechanism 107 to change the shooting distance and reduce the shooting distance, the shooting magnification can be increased.

  When photographing a cross-sectional image, a large number of transmission images are obtained while rotating the subject 104 on the rotary table 105 by the rotary elevating mechanism 106.

  The multiple transmission images are processed by the data processing unit 109, and a cross-sectional image on the imaging surface 111 is obtained from the transmission image passing through the imaging surface 111 orthogonal to the rotation axis 112.

  As the reconstruction method, for example, a filter-corrected back projection method (FBP method) shown in, for example, “CT scanner” (Edited by Yoshinori Iwai: Corona) is used.

  The position of the cross-sectional image of the subject 104 is changed by moving the subject 104 up and down in the direction of the rotation axis 112. However, a helical scan that simultaneously rotates and lifts the rotary table 105 is performed on the imaging surface 111 in one imaging. There is also a method for obtaining a plurality of substantially parallel cross-sectional images (three-dimensional images).

  Further, there is a method of performing a helical scan and obtaining a plurality of cross-sectional images substantially parallel to the imaging tomographic plane 111 using a transmission image passing outside the imaging plane 111.

  The reconstruction method at this time is described in the following document as an application of the FBP method (see Patent Document 1).

  In the first reconstruction method disclosed in Patent Document 1, since the transmission image plane that passes outside the imaging surface 111 is regarded as a parallel plane when it is inclined with respect to the imaging surface 111, it is projected back slightly. On the other hand, the resolution in the direction of the rotating shaft 112 is reduced, whereas the second reconstruction method disclosed in the same document performs back projection according to the tilt, so that the resolution in the direction of the rotating shaft 112 can be increased.

This second reconstruction method can be said to be a method (Feldkamp method) described in the following document applied to a helical scan (see Non-Patent Document 1).
JP-A-4-224736 Practical cone-beam algorithm ”L.A.Feldkamp, L.C.Davis, and JW.Kress J.0pt.Soc.Am./Vol.1,No.6,pp.612-619/June1984

  By the way, recently, there is a strong demand for examinations using a CT scanner, and the types of examination objects and examination contents are expanding. For this reason, there is a tendency that the demand for higher image quality and the demand for higher imaging speed are increasing.

  However, when a plurality of cross-sectional images are obtained by a single scan using the cone-shaped X-ray beam 102 with the above-described conventional high-resolution CT scanner, it takes a long time to reconstruct the cross-sectional images. There is a problem of becoming a bottleneck in speeding up.

  In addition, the above-described Literature 1, Literature 2, etc. do not describe a specific method for speeding up reconstruction when the cone-shaped X-ray beam 102 is used (and simultaneously helically scanned).

  On the other hand, the above-described conventional high-resolution CT scanner has a feature that a high-resolution image can be obtained by bringing the subject 104 close to the X-ray focus. Since the subject 104 is irradiated with the ray beam 102, the scattered X-rays scattered by the subject 104 and incident on the detector 103 increase, and the cross-sectional image becomes unclear depending on the subject 104, so that sufficient inspection can be performed. There is a problem that can not be done.

  In addition, the subject 104 may be soft and easily deformed, and there is a problem that a cross-sectional image becomes unclear due to movement during imaging.

  Furthermore, depending on the arbitrarily selected shooting conditions, the reconstruction filter function may be inappropriate, resulting in a noisy image. In such a case, the filter function is selected again and reconstructed.

  An object of the present invention is to provide a computer tomography method and apparatus capable of contributing to high-speed imaging or high image quality.

  In order to achieve the above object, a computed tomography apparatus according to a first aspect of the present invention provides a radiation source that emits a radiation beam and a radiation detection that detects the radiation beam from the radiation source with two-dimensional spatial resolution. A rotating means for relatively rotating the subject within the radiation beam, and two-dimensional transmission data of the subject obtained by the radiation detector during rotation by the rotating means from a plurality of directions. 0.5, the right and left sides orthogonal to the axis of rotation have a slope that changes from 0 to 1 with a symmetrical slope, and the outside of the slope is multiplied by a window function with 0 on one side and 1 on the opposite side. Reconstructing means for creating a three-dimensional image of the subject by back projection.

  Therefore, in the computer tomography apparatus of the invention corresponding to claim 1, for the purpose of imaging a large subject, a rotation center is set at the end of the radiation beam detected by the radiation detector, When transmission data is obtained by projecting an object outside the beam (offset scan), a sharp change in projection data can be avoided by applying a window function to obtain a three-dimensional image with little false image (ring shape). Can do. Further, it is not necessary to reorganize the data string (conversion to parallel data, etc.), and reconfiguration can be performed at high speed.

  The computer tomography apparatus of the invention corresponding to claim 2 is the computer tomography apparatus of the invention of claim 1, wherein the reconstruction means converts the two-dimensional transmission data into the rotation axis of the volume element of the three-dimensional image. Back projection (two-dimensional centering) onto a centering surface that is approximately parallel to one surface that is approximately parallel to the surface, and backprojects the values on the centering surface to each volume element for each set of volume elements that are approximately parallel to the centering surface. Like to do.

  Therefore, in the computed tomography apparatus of the invention corresponding to claim 2, since the second backprojection is a backprojection between parallel planes, the calculation time of the backprojection coefficient can be shortened and reconstruction can be performed at high speed. Can do.

  Furthermore, the computer tomography apparatus of the invention corresponding to claim 3 is the computer tomography apparatus of the invention corresponding to claim 1 or claim 2, wherein the subject is relatively moved in the axial direction of rotation within the radiation beam ( The translation data is collected while performing the rotation by the rotation means and the translation by the translation means almost simultaneously.

  Therefore, in the computed tomography apparatus of the invention corresponding to claim 3, it is possible to obtain a three-dimensional image over a wide region of the subject in the rotation axis direction at a time.

  On the other hand, a computed tomography apparatus according to a fourth aspect of the present invention is the computed tomography apparatus according to any one of the first to third aspects, wherein the radiation beam is changed to a fan beam along a rotation surface. A collimator to be limited is provided, and scattered radiation correction means for correcting the scattered radiation of the transmission data not blocked by using the transmission data of the path blocked by the collimator.

  Therefore, in the computed tomography apparatus of the invention corresponding to claim 4, the intensity of only the scattered radiation is obtained from the transmission data of the blocked passage, and the transmission data that is not blocked is close to the two-dimensional position of this transmission data. Scattering ray correction can be performed by subtracting the transmission data of the blocked passage, and a high-quality cross-sectional image can be obtained.

  On the other hand, the computed tomography apparatus of the invention corresponding to Claim 5 is the computed tomography apparatus according to any one of Claims 1 to 4, wherein the reconstruction means is a signal noise of the transmission data. A filter function is selected on the basis of the ratio, and filter-corrected back projection is performed to create a three-dimensional image.

  Therefore, in the computer tomography method of the invention corresponding to claim 5, the filter function is selected by the signal-to-noise ratio of the transmission data even if the imaging conditions such as the imaging magnification, tube voltage, tube current, scan time are arbitrarily changed. Therefore, it is possible to obtain a cross-sectional image in which the relationship between image noise and spatial resolution is optimal.

  On the other hand, the computed tomography method of the invention corresponding to claim 6 is the computed tomography method using the computed tomography apparatus according to any one of claims 1 to 5, wherein A step of causing the subject to perform the rotation at a constant speed, and a step of collecting the two-dimensional transmission data while maintaining the rotation at the constant speed to take a three-dimensional image or a cross-sectional image. And

  Therefore, in the computer tomography method of the invention corresponding to claim 6, since imaging is performed without changing the rotation speed after continuing to rotate at a constant speed, it is possible to perform imaging even with a subject that is easily deformed, A high-quality three-dimensional image can be obtained.

  ADVANTAGE OF THE INVENTION According to this invention, the computer tomography method and apparatus which can contribute to the improvement in the imaging speed or the quality improvement of an image can be provided.

Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings.

(First embodiment)
FIG. 1 is a schematic diagram showing a system configuration example of a CT scanner according to the present embodiment.

  In FIG. 1, as an X-ray tube 1 as a radiation source, a microfocus X-ray tube having a focal point F of several to several tens μm of a radiating X-ray beam 2 is used. As a radiation detector 3, a photodiode array is used as a scintillator. X-ray planar solid state detector (or X-ray ll (image intensifier tube) and TV camera) is used.

  The X-ray tube 1 and the radiation detector 3 are arranged to face each other and supported by a support member (not shown) on the floor.

  The subject 4 is placed on the rotary table 5, rotated along the imaging surface 11 in the X-ray beam 2 by the rotation / elevating mechanism 6, and raised and lowered at right angles to the imaging surface 11.

  The subject 4 is moved between the X-ray tube 1 and the radiation detector 3 along the imaging surface 11 by the shift mechanism 7 supported on the floor together with the rotary table 5 and the rotation / lifting mechanism 6. The shooting magnification is changed.

  Although not shown, the mechanism portion is surrounded by X-ray shielding.

  On the other hand, as other components, a data processing unit 9 that processes a transmission image from the radiation detector 3, a display unit 10 that displays processing results, and a mechanism unit are controlled by commands from the data processing unit 9. A mechanism control unit 8 and an X-ray control unit 13 for controlling the tube voltage and tube current of the X-ray tube 1 are provided.

  The data processing unit 9 and the display unit 10 are ordinary computers, and include a CPU, a memory, a disk, a keyboard, an interface, and the like, and store software for reconstructing a three-dimensional image from tomographic sequences and data.

  The operator uses the data processing unit 9 and the display unit 10 to perform menu selection, condition setting, manual operation of the mechanism unit, start of tomography, apparatus status reading, 3D image display, 3D image analysis, etc. Do.

  The data processing unit 9 includes an air correction unit 14 that processes digital data from the radiation detector 3, a LOG conversion unit 15, and a reconstruction unit 16, and the reconstruction unit 16 further includes a filtering unit 16a. And a volume BP portion 16b.

  Next, the operation of the CT scanner configured as described above according to the present embodiment will be described.

  In FIG. 1, when obtaining a transmission image, the operator places the subject 4 on the table 5, sets the tube voltage and tube current, turns on the X-ray, and displays the transmission image on the display unit 10.

  In the case of a three-dimensional image, the subject 4 is moved up and down by the rotation / lifting mechanism 6 so that the center of the examination position is aligned with the imaging surface 11.

  This can also be done while looking at the transmission image.

  When tomography is started, the rotary table 5 is rotated, and transmission images are collected by the data processing unit 9 during this time, and transmission data in the vicinity of the imaging surface 11 obtained at intervals of Δφ in the 360 ° direction are used. A three-dimensional image is reconstructed and displayed on the display unit 10.

  First, the transmission image is corrected by the air correction unit 14 for each channel by taking a ratio with the data da when there is no subject 4 collected in advance.

h = (d-doff) / (da-doff) (1)
Here, doff is data when the X-ray is OFF.

  Next, each channel is logarithmically converted by the LOG converter 15 and converted into projection data p corresponding to the line integral of the absorption coefficient.

p = LOG (1 / h) (2)
The projection data of the n and m channels at the rotation angle φ is described as pφ (n, m), and then a high frequency enhancement filter is applied in the n direction along the imaging surface 11 by the filter application unit 16a.

  This is done by performing a Fourier transform in the n direction, applying a filter function that is approximately proportional to the frequency in the frequency space, and returning by an inverse Fourier transform.

  Next, the volume BP unit 16b performs back projection on each volume element (voxel).

  This back projection processing will be described with reference to the geometric diagram shown in FIG. 2 and the flowchart shown in FIG.

  Note that description of the above-described reference signs and signs that are obvious from the drawings is omitted.

  FIGS. 2A and 2B are the positions of the X-ray focal point F and the detection surface 17 viewed at coordinates x, y, and z fixed to the subject 4. Let the rotation angle of F be φ.

  Each voxel of the volume to be reconstructed is numbered by i, j, k, and the voxel size is Δx, Δy, Δz.

  Here, Δy is preferably set equal to Δx, and Δz is also preferably equal to Δx.

Next, in FIG. 3, the φ loop for 360 ° is obtained in step S1, and the x and y coordinates xF and yF of F are calculated in step S2, and xF = FCD · sinφ (3)
yF = FCD · cosφ (4)
FCD: focus, enter the voxel j, i of the loop rotation center distance step S3, S4, determined nij, △ mij, the L ² in step S5.

nij, mijk: channel number BP to voxel i, j, k Δmij: increment of mijk when 1 increases 1 / L ² : BP weight x = (i−ic) · Δx ( ² ) 5)
y = (jc−j) · Δy (6)
Δx, Δy, Δz: Voxel size φ ′ = arctan ((xF−x) / (yF−y)) (7)
θ = φ−φ ′ (8)
nij = nc + FDD.tan.theta. / d (9)
L ² = ((xF−x) ² + (yF−y) ² ) / FCD ² (10)
Δmij = FDD / (√ (L ² ) · FCD · cos θ) · Δz / d (11)
d: Channel interval FDD: Focus, detection surface distance In step S5, an initial value of mijk is further set.

mijk = mc + (ks−kc) · Δmij (12)
Next, in step S6, the k loop is entered, and in step S7, addition to the voxels i, j, k is performed.

pbp (i, j, k) = pbp (i, j, k) + pφ (nij, mijk) / (L 2 + (kc-k) 2 · △ z 2 / FCD 2)
≒ pbp (i, j, k) + pφ (nij, mijk) / L ² (13)
Here, since nij and mijk are generally not integers, interpolation calculation is performed.

  Next, in step S8, mijk is updated for the next k.

mijk = mijk + Δmij (14)
By repeating k, i, j in steps S9, S10, and S11, the entire volume is added. In step S12, φ is repeatedly applied to one volume, and the backprojection processing is completed. I can make an image.

  In the above description of the backprojection process, the formula is not the fastest for easy understanding, and the formula includes many parts that can be put out of the loop.

  As described above, in the CT scanner according to the present embodiment, since the loop of voxels in the direction of the rotation axis 12, that is, the k loop is the innermost, the calculation of the corresponding channel nij, mijk (or Δmij) is performed within the k loop. Since it can be made common, there is no waste in calculation compared to the case where the k loop is outside the i and j loops as in the conventional case described above.

  In addition, since the data can be reconfigured without waiting for other data in the order in which the data is collected, the reconfiguration speed can be increased.

  Furthermore, since the back projection is performed according to the inclination from the imaging surface 11, the resolution in the direction of the rotation axis 12 can be increased.

(Modification of the first embodiment)
In addition to 1 / L ² , the BP weight may be multiplied by a weight w (φ) depending on the rotation angle as shown in FIG.

  In this case, data collection is performed during 360 ° + 2α rotation. When this weight is applied, before and after 360 ° rotation is averaged, even if the subject 4 slightly moves on the rotary table 5 during imaging, a false image is hardly generated.

  The weight w (φ) may have a curved slope portion. In other words, it is only necessary to be 1 when added with the one shifted by 360 °.

(Second Embodiment)
The system configuration of the CT scanner according to this embodiment is the same as that shown in FIG. 1, and only the method of back projection processing by the volume BP unit 16b is different.

  Therefore, here, the back projection processing (action) of the CT scanner of this embodiment will be described with reference to the geometric diagrams shown in FIGS. 5 and 6 and the flowchart shown in FIG.

  Note that description of the above-described reference signs and signs that are obvious from the drawings is omitted.

 FIG. 5 shows the positions of the X-ray focal point F and the detection surface 17 viewed at coordinates x, y, z fixed to the subject 4. Let the rotation angle of F be φ.

  Similarly in FIG. 6, each voxel of the volume to be reconstructed is numbered by i, j, k, and the voxel sizes are set as Δx, Δy, Δz (Δy = Δx).

Next, in FIG. 7, the projection data p for 360 ° in step S1 is divided into quarters for 90 °, and in step S2, the first 90 ° φc (calculation rotation angle) loop is formed. Calculate the collection rotation angle φ. Φ = φc + nq · π / 2 (15)
nq: Quarter number In step S4, it is determined whether or not φ is the φ being collected, and the head is started or terminated.

  In step S5, x and y coordinates xF and yF of F are calculated.

xF = FCD · sinφc (16)
yF = FCD · cosφc (17)
Next, centering is performed in steps S6 to S11.

  In FIG. 5, centering is a process of back-projecting data on the detection surface 17 onto a pq matrix on the xz plane.

  In step S6, a p-loop is entered, and np is obtained in step S7.

xo = (p−pc) · cp (18)
x1 = xo · cosφc (19)
mag1 = FDD / √ ((xF−xo) ² + yF ² −x1 ² )) (20)
np = nc + x1 · mag1 / d (21)
np, mpq: channel number cp, cq: centering pitch BP to the matrix p, q Next, the q loop is entered in step S8, mpq is obtained in step S9, and centering data pc (p, q) is obtained in step S10.

mpq = mc + (q−qc) · cq · mag1 / d (22)
pc (p, q) = pφ (np, mpq) (23)
Here, since np and mpq are generally not integers, interpolation calculation is performed.

  Next, the following q and p are repeated in step S11.

  Next, BP is performed in steps S12 to S22.

  In FIG. 6, the j loop is entered in step S12, and poj, Δpj, and Δqj are obtained in step S13.

mag2 = yF / (yF + (j−jc) · Δy) (24)
poj = pc + (xF−mag2 · (xF + ic · Δx)) / cp (25)
Δpj = mag2 · Δx / cp (26)
Δqj = mag2 · Δy / cp (27)
pij, qjk: matrix number BP assigned to voxel i, j, k Δpj: increment of pij when i increases by 1 Δqj: increment of qjk when k increases by 1 Set the value.

pij = poj (28)
Then, enter the i loop S step 14, it calculates the ^{L 2} in step S15.

x = (i−ic) · Δx (29)
y = (jc−j) · Δy (30)
L ² = ((xF−x) ² + (yF−y) ² ) / FCD ² (31)
In step S15, an initial value of qjk is further set.

qjk = qc + (ks−kc) · Δqj (32)
Next, in step S16, the k loop is entered, and in step S17, addition to the voxels i, j, k is performed.

pbp (i, j, k) = pbp (i, j, k) + pc (pij, qjk) / (L 2 + (kc-k) 2 · △ z 2 / FCD 2) ≒ pbp (i, j, k ) + Pc (pij, qjk) / L ² (33)
Here, since pij and qjk are not generally integers, interpolation calculation is performed or the nearest neighbor pc is selected.

  Next, in step S18, qjk is updated for the next k.

qjk = qjk + Δqj (34)
The k loop is repeated in step S19, and pij is updated for the next i in step S20.

pij = pij + ΔPj (35)
In steps S21 and S22, the entire volume is added by repeating i, j, and φc is repeatedly added to one volume in step S23.

  In step S24, the image is rotated by 90 degrees, and in step S25, the back projection process is repeated by repeating the quota, and a three-dimensional image of the volume is formed.

  In the above description of the backprojection process, the formula is not the fastest for easy understanding, and the formula includes many parts that can be put out of the loop.

  As described above, in the CT scanner according to the present embodiment, first, when the centering data pc is obtained by back projecting the data on the detection surface 17 onto the xz plane at each rotational position φ, pc is one surface ik of the voxel. Since it is obtained at equal intervals on a plane parallel to the plane, the backprojection calculation to the ik plane (specifically, calculation of pij and qjk) can be remarkably simplified, and the reconstruction speed can be increased. It becomes.

  Similarly to the first embodiment described above, since the loop of voxels in the direction of the rotation axis 12, that is, the k loop, is the innermost, the calculation of the corresponding channels pij, qjk (or Δqj) is performed within the k loop. Since it can be made common, there is no waste in calculation compared to the case where the k loop is outside the i and j loops as in the conventional case, and the reconstruction speed can be increased.

  Furthermore, since the data can be reconfigured without waiting for other data in the order in which the data is collected, the reconfiguration speed can be increased.

  Furthermore, since the back projection is performed according to the inclination from the imaging surface 11, the resolution in the direction of the rotation axis 12 can be increased.

(Modification of the second embodiment)
As in the modification of the first embodiment described above, the BP weight may be multiplied by a weight w (φ) depending on the rotation angle as shown in FIG. 4 in addition to 1 / L ² . Also in this case, the same effect can be obtained.

  In the second embodiment, when the volume k direction is the same size as the ij direction, the i loop may be located inside the k loop. In this case, equation (33) is not an approximate calculation, but the first row is calculated.

(Third embodiment)
The system configuration of the CT scanner according to this embodiment is the same as that of the second embodiment, and only the mechanism operation at the time of transmission image collection and the method of back projection processing by the volume BP unit 16b are different.

  Therefore, here, the operation at the time of transmission image collection and the back projection processing (operation) of the CT scanner of the present embodiment will be described with reference to the flowchart shown in FIG.

  In the case of the present embodiment, when tomography is started, the rotary table 5 is moved up and down simultaneously with rotation (helical scan), and transmission images are collected during this time, and the three-dimensional volume is obtained from transmission data in a direction of 360 degrees or more. The image is reconstructed and displayed on the display unit 10.

  The volume BP of the present embodiment is substantially the same as that of the second embodiment described above, and will be described using the flowchart shown in FIG. 7 and equations (15) to (35).

  That is, the difference from the case of the second embodiment described above is that kc is not a constant but a function of φ, and equation (33) is different.

kc = kc (φ) = zp / Δz · (φ−φ0) / 2π (36)
pbp (i, j, k) = pbp (i, j, k) + wφ (k) · pc (pij, qjk) / (L 2 + (kc-k) 2 · △ z 2 / FCD 2) ≒ pbp ( i, j, k) + wφ (k) · pc (pij, qjk) / L 2 ··· (33 ')
Here, each equation will be described with reference to FIG.

  First, in equation (36), kc (φ) is the k position of the matrix traversed by the imaging surface 11 at the rotation angle φ.

  In this equation, zp is the height (helical bitch) that moves up and down during one rotation, Δz is the voxel size in the height direction, and φO is the rotation angle when crossing k = 0.

  In FIG. 8A, wφ (z) is a weight of the BP of the voxel in the z direction at the rotation angle φ and has a trapezoidal shape with a half width zp.

  On this trapezoidal slope, BP from φ different by 360 ° is synthesized (dotted line).

  The weight obtained by changing wφ (z) to k scale is wφ (k), and the center shown in FIG. 8B is a trapezoid with kc (φ) and half width zp / Δz.

  zp / Δz at k is a rotation angle and corresponds to 360 °.

Here, the half length kα of the slope is an angle α and kα = Zp / Δz · α / 2π (37)
Therefore, kα is determined by setting the angle half width α to be interpolated.

  In Expression (33 ′), wφ (k) is used as the BP weight, and therefore, when helical scanning is performed, BP corresponding to 360 ° is made to each voxel even if BP is continuously performed for φ.

  As described above, in the CT scanner according to the present embodiment, it is possible to increase the reconstruction speed and the resolution in the direction of the rotation axis 12 as in the second embodiment described above.

  In addition, since the image can be captured efficiently by helical scanning without increasing the inclination of the beam from the imaging surface 11, a wide and high-quality three-dimensional image in the direction of the rotation axis 12 can be obtained.

  Further, since a wide range can be smoothly scanned at a constant speed of rotation and elevation by helical scanning, the subject 4 is not accelerated and a high-quality image can be obtained even on the subject 4 that is easy to move.

  Further, since the BP is multiplied by the trapezoidal weight wφ (k), before and after the rotation of 360 ° is averaged, even if the subject 4 moves during imaging, a false image is hardly generated.

(Modification of the third embodiment)
As shown in FIG. 8C, the weight wφ (k) may have a curved slope. That is, it is only necessary to be 1 when added with the one shifted by zp / Δz (dotted line).

(Fourth embodiment)
The system configuration of the CT scanner according to the present embodiment is substantially the same as that of the third embodiment, and includes a window function multiplying unit between the LOG conversion unit 15 and the reconstruction unit 16, and the rotation center. Is different from the center of the X-ray beam 2 when viewed from above.

  FIG. 9 is a schematic diagram showing the positional relationship between the rotation center C and the X-ray beam 2.

  This is a CT scanner in which the rotary table 5 is set to be shifted and the subject 4 is placed on one side so as to protrude from the X-ray beam 2 and the large subject 4 can be photographed. Called scan.

  The operation of the CT scanner according to the present embodiment is the same as that of the third embodiment, and a window function w (n as shown in FIG. 10 is applied to the projection data pφ (n, m) after LOG conversion. ) Is only different.

  The window function w (n) is center ch, nc, and is inclined about 0.5 at the center. One outside the inclined area is 0, and the other is 1.

  The reconfiguration unit 16 performs reconfiguration as in the third embodiment.

  As described above, in the CT scanner according to the present embodiment, it is not necessary to reorganize the data string (conversion to parallel data, etc.) even in the case of offset scanning, and can be reconstructed at high speed. Since a smooth weight is applied across the center ch, a sudden change in projection data can be avoided, and a three-dimensional image with few false images (ring-shaped) can be obtained.

  Furthermore, an offset scan can be performed while performing a helical scan, and in addition to obtaining the same effects as those of the third embodiment described above, a CT scanner capable of handling a large subject 4 is realized. It becomes possible.

(Modification of the fourth embodiment)
The offset scan according to the fourth embodiment can also be applied to the first or second embodiment described above.

  The window function w (n) may have a curved slope. That is, it is only necessary that the center ch has an inclined portion that changes from 0 to 1 that is symmetric with respect to the left and right sides that are orthogonal to the axis, and one side of the outside of the inclined portion is 0 and the opposite side is 1.

(Fifth embodiment)
FIG. 11 is a schematic diagram showing a system configuration example of the CT scanner according to the present embodiment, which is substantially the same as that of the first embodiment, and the scattered radiation correction is performed between the air correction unit 14 and the LOG conversion unit 15. The difference is that a portion 19 is provided, and a movable collimator 18 that restricts the X-ray beam 2 to a thin cone shape sandwiching the imaging surface 11 is provided.

  The operation of the CT scanner according to this embodiment is the same as that of the first embodiment. The X-ray beam 2 is limited to a thin cone at the time of tomography, and the data h (n, m after air correction) ) Is different from that in FIG.

  FIG. 12 is a diagram for explaining the content of the scattered radiation correction.

  FIG. 12 shows data h (n, m) on the detection surface 17.

  Except for the irradiation unit 20 of the X-ray beam 2, only the scattered radiation from the subject 4 or the like is detected at a portion blocked by the collimator 18.

Scattered ray correction h ′ (n, m) = h (n, m) − ((m2−m) · h (n, m1) + (m) for the data h (n, m) in the irradiation unit 20 −n1) · h (n, m2)) / (m2−m1) (38)
Add

  Here, the scattered light is obtained by subtracting the upper and lower channels close to h (n, m) by linear interpolation, and subtracted.

  As described above, in the CT scanner according to the present embodiment, the scattered radiation correction can be performed using the data of the portion blocked by the collimator 18 of the two-dimensional detector without using a dedicated detector for scattered radiation measurement. it can.

  Moreover, since the two-dimensional detector is manufactured for all the channels at the same time with respect to the dedicated detector, the characteristics are uniform.

  Furthermore, since the scattered radiation can be measured at a position close to the measurement point at the same time and under a uniform condition, it is possible to correct accurately.

(Modification of the fifth embodiment)
In the fifth embodiment, when performing the correction, not a 2-channel data but a larger number of channel data may be used, and the primary interpolation may not be performed.

  By using this large number of channel data, the statistical accuracy can be improved.

  The scattered radiation correction in the fifth embodiment can be added not only to the first embodiment but also to the second to fourth embodiments. It can also be added to a CT scanner using a conventional two-dimensional detector as shown in FIG.

(Sixth embodiment)
The system configuration of the CT scanner according to the present embodiment is almost the same as that of the first embodiment, except that it is applied to the imaging of the subject 4 that is easily deformed.

  In the case of the present embodiment, the subject 4 is set on the table 5 and is rotated as a preparation at a smooth constant speed. After the subject 4 is sufficiently stabilized, tomography is started.

  The apparatus turns on X-rays while maintaining the same rotational speed, collects transmission images, and creates a three-dimensional image.

  In the case of the helical scan, the preparatory rotation is performed in the same manner, and when the tomography is started, the elevation is gradually started to start collecting transmission images.

  If the preparation rotation time may be short, preparation lifting may be performed together.

  As described above, in the CT scanner according to the present embodiment, the movement of the subject 4 that occurs at the start of rotation is stopped at a constant speed and then imaged without changing the rotation speed. The specimen 4 can be photographed with the movement stopped, and a high-quality three-dimensional image can be obtained.

(Seventh embodiment)
The system configuration of the CT scanner according to this embodiment is almost the same as that of the first embodiment, except that the filter function is automatically selected.

  That is, the data processing unit 9 calculates a signal noise ratio (average) of data (p or h) before filtering and automatically selects a filter function.

  The signal-to-noise ratio is determined by taking the low frequency component as a signal and the high frequency component as noise, and determining the ratio of signal / noise.

  If the signal-to-noise ratio is small, the filter function is chosen to have a low cut-off frequency to reduce image noise instead of slightly reducing spatial resolution.

  As described above, in the CT scanner according to the present embodiment, the filter function is selected based on the signal-to-noise ratio of the transmission data even if the imaging conditions such as imaging magnification, tube voltage, tube current, and scan time are arbitrarily changed. A cross-sectional image in which the relationship between image noise and spatial resolution is optimal can be obtained.

1 is a schematic diagram showing a first embodiment of a CT scanner according to the present invention. The geometric diagram for demonstrating the effect | action in CT scanner of the said 1st Embodiment. The flowchart for demonstrating the effect | action in CT scanner of the said 1st Embodiment. The figure which shows the modification of 1st Embodiment of CT scanner by this invention. FIG. 10 is a geometric diagram for explaining the operation of the CT scanner according to the second embodiment of the present invention. FIG. 10 is a geometric diagram for explaining the operation of the CT scanner according to the second embodiment of the present invention. The flowchart for demonstrating the effect | action in the CT scanner of the 2nd Embodiment and 3rd Embodiment of this invention. The figure for demonstrating the effect | action in the CT scanner of the 3rd Embodiment of this invention and its modification. The fragmentary schematic diagram which shows 4th Embodiment of CT scanner by this invention. The figure for demonstrating the effect | action in CT scanner of the 4th Embodiment. The fragmentary schematic diagram which shows 5th Embodiment of CT scanner by this invention. The figure for demonstrating the effect | action in CT scanner of the same 5th Embodiment. 1 is a schematic diagram showing an example of a system configuration of a conventional high resolution CT scanner.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 ... X-ray tube, 2 ... X-ray beam, 3 ... Radiation detector, 4 ... Subject, 5 ... Rotary table, 6 ... Rotation / lifting mechanism, 7 ... Shift mechanism, 8 ... mechanism control unit, 9 ... data processing unit, 10 ... display unit, 11 ... imaging surface, 12 ... rotation axis, 13 ... X-ray control unit, 17 ... detection surface, 18 ... collimator, 19 ... scattered ray correction unit, 20 ... irradiation unit.

Claims (6)

A radiation source emitting a radiation beam;
A radiation detector for detecting a radiation beam from the radiation source with a two-dimensional spatial resolution;
Rotating means for relatively rotating the subject within the radiation beam;
Two-dimensional transmission data from multiple directions of the subject obtained by the radiation detector during rotation by the rotation means is 0.5 at the rotation axis position, and the inclination is symmetrical on the left and right sides orthogonal to the rotation axis. Reconstructing means for creating a three-dimensional image of the subject by performing back projection by applying a window function in which one side is 0 and the other side is 0 on the opposite side. ,
A computer tomography apparatus comprising: The computed tomography apparatus according to claim 1,
The reconstruction means back-projects (two-dimensional centering) the two-dimensional transmission data onto a centering surface substantially parallel to one surface substantially parallel to the rotation axis of the volume element of the three-dimensional image, A computer tomography apparatus characterized in that a value is projected back to each volume element for each set of volume elements substantially parallel to the centering plane. The computed tomography apparatus according to claim 1 or 2,
Translation means for relatively moving (translating) the subject in the direction of the axis of rotation within the radiation beam is added, and transmission data is collected while performing rotation by the rotation means and translation by the translation means almost simultaneously. A computer tomography apparatus characterized by that. The computed tomography apparatus according to any one of claims 1 to 3,
A collimator for limiting the radiation beam to a fan beam along the plane of rotation;
Using the transmission data of the passage blocked by the collimator, the scattered radiation correction means for correcting the scattered radiation of the transmission data not blocked,
A computer tomography apparatus comprising: The computed tomography apparatus according to any one of claims 1 to 4,
The computer tomography apparatus according to claim 1, wherein the reconstruction unit selects a filter function based on a signal-to-noise ratio of the transmission data, and performs a filter correction back projection to create a three-dimensional image. A computed tomography method using the computed tomography apparatus according to any one of claims 1 to 5,
Causing the subject to perform the rotation at a constant speed before imaging;
Collecting the two-dimensional transmission data while maintaining the rotation at the constant speed and taking a three-dimensional image or a cross-sectional image;
A computed tomography method comprising:

Images