JPS633842A - X-ray ct apparatus - Google Patents

Abstract

(57) [Summary] This bulletin contains application data before electronic filing, so abstract data is not recorded.

Description

[Detailed Description of the Invention] [Object of the Invention] (Industrial Application Field) The present invention is directed to x-ray CT (computed tomography)
The present invention relates to an apparatus, and particularly relates to a reconstruction means for obtaining a radiation distribution on a tomographic plane of a subject.

(Prior Art) In this type of X-ray CT system, for example, a plurality of detectors are fixedly arranged in a circular body surrounding the subject, and fan beam-shaped X-rays are emitted continuously or in a pulsed manner. The X-ray source is rotated around the subject.

Hereinafter, such a CT device will be referred to as a fourth generation CT device. This fourth generation CT apparatus collects fan beam-shaped X-rays emitted from an X-ray source in the negative direction and transmitted through a subject from a plurality of directions to obtain source fan data. The source fan data from these multiple directions are rearranged into detector fan data spread out from each detector in the form of a fan beam. These detector fan data are then convolved and back-projected to reconstruct an image. In this way, the source
The reason for converting fan data to detector fan data is that converting source fan data to detector fan data and performing reconstruction calculations has better spatial resolution than performing reconstruction calculations on the source fan data as is. This is because a tomographic image can be obtained. That is, since the sampling pitch of the source fan data depends on the pinch of the detector, it is difficult to obtain source fan data with a fine numbering pitch. On the other hand, since the sampling pitch of the detector fan is determined by the interval of X-ray irradiation, it is possible to easily obtain data with a fine sampling pitch. For this reason, in the fourth generation CT apparatus, reconstruction calculations are usually performed after converting source fan data into detector fan data.

Using this fourth generation CT device, the so-called dynamic scan method, in which X-rays are continuously rotated around the subject and the temporal changes in the image of the resting area of the subject are reflected, is shown in Figures 3 and 5. This will be explained with reference to FIG.

The detector ring 1 shown in FIG. 3 is fixed in a donut shape with a hole in the center that is large enough for the subject 3 to fit inside. This detector ring 1 has, for example, 2304 X
The ray detectors 1 #1 to #2304 are arranged in a donut shape to form a groove, and detect the X-rays from the subject 3.

The XIjl source 2 rotates around the detector ring 1 around the central axis of the detector ring 1 in the direction of the arrow while continuously emitting X g. The rotation angle positions of this X-ray source 2 are equally assigned from P1, which is three times the number of detectors, to P6912. At this time, when the rotational position of the X-ray source 2 is P1,
The X-ray beam from the radiation source 2 passes through the object and reaches the detector 1#.
1682 to #2304 and #1 are detected. When the rotation angle position is P4, the detectors 91683 to #23
X-rays are detected at #04 and #1 and #2.

FIG. 5 shows this skimming state using a sinobram in which the vertical axis is the detector number and the horizontal axis is the rotational angular position and time of the X-ray source 2.

Now, when t1=0, if the rotation angle position of X-ray source 2 is Pl, the X-ray beam from X-ray source 2 is transmitted to detector 1 #1682.
It is detected by the detectors #2304 to #1. Further, the rotation angle position 2.3 is also detected by the same detector as the edge at the rotation angle position P1. Rotation angle position P4. P5.
P6 corresponds to detectors 1 #1683 to #2304 and #1 to 2. Thereafter, detection is performed in the same manner, and when the rotation angle positions are P6910 to P6912, detectors 1 #1681 to #2304 correspond.

In other words, source fan data is arranged in the column direction, and detector fan data is arranged in the row direction. Furthermore, since scanning is performed continuously, this data group continuously advances diagonally downward to the right on the sinobram.

In order to further reconstruct one tomographic image, it usually takes 36
Detector fan data from the 0° direction, for example, detector fan data corresponding to detectors 1#1 to #2304 is required. This corresponds to portions B and A' shown in FIG. To obtain one tomographic image at this time, the data detected from X-ray source position 1 (t1 = o) to X-ray source position 1872 (t3 = 1.27) is used. rIfI containing "blur"
It becomes a layered image. The "blur" at this time corresponds to about 1.27 seconds.

Therefore, by paying attention to the fact that part A and part 8' are the same in position, and moving the data of A to A' and re-creating the bridge, the "blur" time is reduced to 1 second. I will make it happen. This "blur" occurs because the X-rays rotate around the subject while being irradiated to create two tomographic images. It refers to the error in the time change until the end. That is, since the subject moves during X-ray exposure due to breathing, body movement, heart, etc., fluctuations (errors) occur in the values of the source fan data.

The error due to this "blur" corresponds to the rotation time of xjlA required to create one tomographic image, and the shorter the rotation time, the smaller it is.

Next, a dynamic scan in which the same area is continuously scanned using such a fourth generation CT apparatus to obtain a plurality of temporally different images will be described with reference to the sinobram in FIG. 6. In the dynamic scan in the same area, first the xi source is 1 from rotation angle position P1 to P6912.
One image is obtained from the data obtained by rotation. Next, for example, data obtained by rotating the X-ray source from rotational angular position P6912 to P6913 is replaced with previous data corresponding to the rotational angular position, and an image shifted by the rotational time is obtained. Thereafter, similarly, the X-ray source is rotated little by little, and the data obtained thereby is replaced with previous data corresponding to the rotational angular position to sequentially obtain images temporally shifted by the rotational valve.

When performing a reconstruction operation by replacing this newly obtained data with the previous data whose rotation angle position corresponds to it, the previous data that was not replaced is replaced by the convolution pack projection performed when obtaining the previous image. It takes time because it has to be repeated again. Therefore, in order to shorten the reconstruction calculation, the same inventor as the present inventor has proposed a technique disclosed in Japanese Patent Application Laid-Open No. 57-134142@. This takes advantage of the fact that convolution and backprojection are linear operations, finds the difference between newly obtained data and the corresponding previous data, performs convolution on that difference, and calculates the previous backprojection. This is a back project to the projected data.

In this way, it is not necessary to repeatedly perform convolution and back projection on duplicate data, so the reconstruction time can be shortened.

(Problem to be Solved by the Invention) However, when applying such technology to 4th generation CT, the first image is a source fan obtained by rotating the xjlA source from rotational angular position P1 to P6912. The data is reconstructed from the rearranged detector fan data corresponding to detectors #1 to #2304. The next image shows the source fan data obtained when the X-ray source rotates from rotation angle position P6912 to P6913 and the source fan data from the previous overlapping rotation angle position MP2 to P6912 on detectors #2 to #2304. and #1, and are rearranged and reconstructed. Here, referring to FIG. 6, the source fan data C at the rotational angular position P1 is the source fan data C at the rotational angular position P6.
913 C'. Looking at this from the detector fan data, detectors #1682 to #2304
A change will occur in the detector fan data corresponding to #1. In other words, detector fan data D
In addition to finding the difference between and D' and performing convolution and backprojection on the difference value, it is also necessary to perform convolution and backprojection again for these detectors #1682 to #2304.

Therefore, when such dynamic scanning is performed in a fourth generation CT apparatus that converts source fan data into detector fan data, there is a problem in that the time required to obtain an image is longer than in the third generation CT.

A first object of the present invention is to provide an X-ray CT apparatus that can obtain high-resolution images without converting source data to a detector fan.

The second object of the invention is to continuously rotate the radiation source so that temporally different sources detected at the same rotational angular position
The purpose of the present invention is to provide an X-ray CT apparatus that obtains a difference value between 7 images, performs convolution on this difference value, back-projects the convolved data onto the previous image, and continuously obtains images. shall be.

[Configuration of the Invention (Means for Solving the Problems)] In order to achieve this object, the present invention provides a radiation source that generates fan-beam-shaped radiation toward a cross section of a subject; a radiation detection means comprising a plurality of radiation detection elements arranged so as to surround the specimen, and detecting the fan beam-shaped radiation transmitted through the specimen; and a scanning device that rotates the radiation source around the specimen. and a convolution function that inserts a zero between each data of the source fan data sequentially supplied from the radiation detection means in accordance with the rotation of the radiation source, and corresponds to the source fan data into which the zero is inserted. This convolved source
An X-ray CT apparatus comprising: a reconstruction means for back-projecting fan data; and a display means for displaying an image of the cross section of the subject reconstructed by the reconstruction means.

The radiation source is continuously rotated, a difference value between temporally different source and fan data detected at the same rotation angle position is determined, a convolution is performed on this difference value, and the convolved data is calculated. is characterized by back-projecting the image onto the previous image.

(effect) First, the principle of the present invention will be explained.

No. 732,280, filed in the United States on September 5, 1985 by the same applicant as the present applicant, describes an expanded diverging end method. This extended divergent method can be used when certain data and its opposite data exist, and if the sum of their respective weights is constant, even if both data are multiplied by an arbitrary weight and then reconstructed, the reconstructed image There is no difference.

Here, as shown in FIG. 10, the X-ray beam emitted from the X-ray tube 2 at the rotational position θ1 and detected by the detector group is R1-R.
4, the X-ray beams R1' to R3' existing between these beams R1 to R4 are obtained when the X-ray tube 2 reaches the rotational position between each of these beams R1 to R4. . And each of these beams R1 to R4, R1' to R3'
The beam as "opposing data" is shown in the broken line, and the weighting of the data given by 1!7 by the X-ray beams R1 to R4, R1' to R3' and these "opposing data" is determined as shown in the figure. This then satisfies the conditions of the extended divergent end method described above.

Therefore, focusing on the XFA fan beam at the rotational position S, the data obtained from the X-ray beams R1 to R4 are weighted with "1", and the data between each channel is weighted with "0" for reconstruction. However, when looking at the entire data system, it satisfies the conditions of the extended divergent end method.

FIG. 7 shows detection elements DI arranged circumferentially at equal intervals,
D3 , ......, Dll and the detection elements [)1 , [)3 , [)] while rotating around this circumference, respectively.
5..., sampling pitch 31, 32 with twice the density of Dll.

. . . shows the relationship with the X-ray source that generates fan-beam radiation in S12. In the figure, the X-ray source rotates on the circumference of the detection element to simplify the explanation, but usually one of the radii is small.

Now, when the X-ray source is set at position 81 and a fan beam-shaped X-ray is emitted in the direction of the central axis, the fan-beam-shaped X-rays are detected by the detection elements D3, D5, D7, and D9.

[)11 to detect.

Next, when the X-ray source rotates to 82 and irradiates, the X-ray beam spread out in a fan beam shape is transmitted to the detection elements DS, D7. D9
．． Detected at D11. Furthermore, the X-ray source rotates and irradiates at positions 83 to S12 in FIG. FIG. 8 shows the numbers of the detection elements that detect each X-ray beam at each irradiation position of the X-ray source.

FIG. 8 shows the detection elements DI, D3, D5 in the vertical direction.
.

D7, D9, and Dll are lined up, X-ray source positions S1 to 812 are lined up in the horizontal direction, and the X-ray bus from the X-ray source position to the detection element is shown in the form of Xs, d. The meaning of Xs, d is an abbreviation, for example, the X-ray beam from 81 to [ ) 11 in FIG. 7 is represented by Xl, 11. Therefore, this X 1
．． 11 is shown in the square at the bottom left of FIG. 8 where 81 and [)11 intersect. According to this figure, the locations corresponding to the respective detection elements, for example, between Dl and D3, between D3 and D5,
There are no X-ray buses in the horizontal direction.

This corresponds to 32, 34, 36 in the layout diagram in Figure 7.
, 88, 810, and 812 because there are no detection elements at the positions. In FIG. 8, convolution with source fan data is performed along the vertical grid. The detector fan performs convolution on the detector fan data rearranged along horizontal grids. In FIG. 8, the detection element DI,
D3, D5, D7.

For Q9 and [)11, there are more paths aligned in the horizontal direction, so when reconstructing with the 4th generation) ICT device, it is better to convolve the detector fan data than the source fan data. improves spatial resolution.

Next, since the X-ray bus XII,n in FIG. 7 is equal to Xn,l, the paths in FIG. 8 are rearranged as shown in FIG. 9. For example, in Fig. 9, the x4.1 x-ray path is S
This is the path when the X-rays from 81 are detected by the detection element D1, or in other words, the X-rays from 81 are detected at the position 84. Using this fact, move the path of Xn,1 to the 81 vertical square. Similarly, other X-ray passes were transferred, 81, 83, S5, S7, S9,
Move to the vertical square of S11. The source fan data has the same density as the deidecta fan, so if the image is reconstructed using the path transferred in this way, the reconstructed 'ili
The resolution of the image is the same as that of the detector fan data. Furthermore, even if zero data is inserted and convolution is performed for the vertical squares of 82, 34, 86, 38, and 310 degrees S12, the image will not change because only zero will be added. The image spatial resolution of the reconstruction result by the source fan at this time is the same as the spatial resolution by the detector fan in FIG. This is because the same number of data (number of passes) is used in both FIG. 8 and FIG. 9, and the same number of data is used for the -degree convolution calculation.

In contrast to FIG. 9, consider bus data not shown in FIG. 11. Figure 11 shows that X using the relationship Xn,m =Xm,n
Of the data of s and d, s = 2.4.6.8°10.1
Instead of the data in 2, I bought 6 columns and 8 rows of data. In other words, the data that was moved when creating FIG. 8 to FIG. 9 is returned to its original state, but the data in column rIAd and row S remains as is.

For example, ×2.5 is in the 5th column and 2nd row in Figure 9, but
Although it is located in the same position in FIG. 11, it is also placed in the second column and fifth row. Now consider FIG. 12a.

FIG. 12a is a diagram in which all the coefficients of the data section added to create FIGS. 9 to 11 are set to zero. Therefore, when the weighting coefficient of FIG. 12a is multiplied by FIG. 11, FIG. 9 is obtained. Therefore, the image obtained by applying the weighting relationship shown in FIG. 12a to FIG. 11 and performing convolution back projection using the source fan is the same as the image obtained by reconstructing FIG. 9 using the source fan. In other words, the image is the same as the image shown in FIG. 8 reconstructed using a detector fan.

Here we apply the extended divergent end method. According to this, when there is opposing data, if the sum of the 1-values is constant, the reconstructed image will be the same even if an arbitrary weight is applied to both data and then reconstructed. .

In FIG. 11, there are cases where the same data exists twice. This weight is 1" on one side and 1" on the other in the m12a diagram.
It is 101P. For example, the data in the 1st column, 4th row, and the 4th column, 1st row are both ×4.1, and the weighting coefficient is 11$1 for each.
゜It is “0”. According to the extended diverging end method, even if this is changed to °JQIM "1", the image remains the same. FIG. 12b shows the same data with the weighting coefficients reversed in this way. Therefore, what is reconstructed by applying the weights shown in FIG. 12b to FIG. 11 and convoluted with the source fan is the same as what is reconstructed by convolving FIG. 8 with the detector fan.

Since the path data is weighted by a function and then convolved, if the weighting function is zero, the weighted result will be zero regardless of the data value. If this is summarized with respect to zero using FIG. 11 and FIG. 12b, the result will be as shown in FIG. 13. It can be seen that FIG. 13 is equivalent to the bus data in FIG. 8 before being reflected with zeros inserted between them.

Therefore, if zero is inserted between the data obtained by the detector and convolution is performed on the fan data, the spatial resolution of the resulting tomographic image will be 1 compared to the image obtained when convolution is performed on the conventional detector fan. is equal to the space valve WI capacity.

In this principle, X-rays are irradiated once from the source between the detection elements, for example Dl and D3, but this is done by irradiating X-rays multiple times between the detection elements and inserting multiple zeros between the detection data. You can also reconfigure it by doing so.

(Example) An embodiment of the present invention will be described below with reference to FIG.

The X-ray detector 1 includes a plurality of X-ray detection elements, for example, a combination of a scintillator and a photodiode, arranged in a circumferential manner so as to surround the subject 3. An X-ray source 2 that generates fan-beam X-rays rotates around the outside of this detector 1. The nutate section 12 is arranged to prevent the X-ray source 2 from being blocked by the detection element closer to the X-ray source 2.
gives a twisting motion to the X-ray detector 1 in synchronization with the rotation of the The scanning unit 4 rotates the X-ray source 2 and sends a signal synchronized with the rotation to the nutate unit 12.
and supplies it to DAS5. The DAS 5 is connected to each detection element of the detector 1, collects analog source/fan data supplied according to the rotation of the X-ray source 2, and converts it into digital source/fan data.

Since the detection elements are arranged at equal intervals with respect to the rotation center, the angular intervals of the source fan data are unequal. The interpolation unit 6 corrects the digital source fan data at unequal angular intervals supplied from the DAS 5 so that the angular intervals of the source fan data are equal.

For weighting, the circuit 7 inserts O between each data of the interpolated source fan data, apparently doubling the data density. The convolver 8 performs convolution with the source fan data using a convolution function corresponding to this seemingly double density source fan data. The back projector 9 back projects the convolved data in correspondence with the rotation angle of the X-ray source 1. The image memory 10 has a memory space corresponding to the subject 3 and sequentially stores back-projected data. The display section 11 is an image memory 10
Display the reconstructed images stored in . Insert zero between each data of source fan data and then convolution? ! The convolution function 9, gs, or gt to be used when specializing in lli is a (n ・△ψ) = B
n =0-1/2π2 m-2n ・Δψ n+0 However, B=2-Σ(1/2 π2m-2n −△ψ)n
= 1 Here, N is 1/2 of the detected temperature, or gs(n△ψ
)=1/π5in(nΔψ/2)1-4N/n ・・・・・・・・・・・・ (4) Or, gt(n △ψ) =1/40S((n-1)Δψ4
/2) +1/2os(n△4) +1/4g5((n +1)Δψ)・・・・・・・・・
...(5) etc. can be used. Here, Δψ is the angular interval between adjacent source fan data, and n is the bus number from the bus passing through the center of rotation.

Therefore, if the convolution function as shown by these equations (6), (7), and (8) is (1(nΔψ)), then in Fig.
φ) ・Rcos(n△ψ)△ψΔθ (6)

However, Lllll (7, ψ) = [R-rsin (
I Δθ+φ) ]λ + [rCO3(IR△θ+φ)J2'...
... (7) φta-tan(-γcos(m△θ+φ)/R-rs
in (mΔθ+φ) ・・・・・・・・・ (8) M-3606/80N-(fan angle/2)/Δ4 Δθ is the rotation angle al of the X-ray source! The distance R represents the radius of rotation of the XtQ source.

It shows the Kushikon operation. Note that Lffl (r, φ) and Rcos (nΔψ) are Jacobians for FA standard transformation.

Further, in this weighting, the coefficient W is set to 1 or O. Here kv instead of W (k is any constant other than O)
It is also possible to use, for example, =1/2. At this time, W takes a value of 1 or O. However, in the case of k≠1, the grayscale value of the reconstructed image appears to be twice the original value, so it is necessary to multiply it by 1/ after the image reconstruction. For example, when the value is -1/2, the gradation value of the created image may be shifted by 24a.

Next, FIG. 2 shows a data arrangement when there are two corresponding beams between the arrangement pitches of the detectors. From the left of the drawing, the data of ray 1 in FIG. 2 is set to X1111, the data for virtual rays mla and m1b are set to zero, and the rest is set in the same manner as m2->X12. m2a, n+2b->zero.

It is expressed as m3→Xm3. Using these data, a calculation is performed using a convolution function when reconstructing the detector fan, and then a back projection calculation is performed.

When such a zero is inserted and a convolution operation is performed, the result of the convolution operation is added to the position where the zero was inserted later. Therefore, by inserting zeros according to the ray density of the detector fan,
Almost the same spatial resolution as fan reconstruction is obtained.

Next, an example of the function coefficients used in convolution calculation using the source fan beam will be described with reference to FIGS. 14 and 15. FIG. 15 is an example of a conventional convolution function, and shows seven function values when performing - times of convolution operations. Each function value is shown as a value multiplied by -i ooo. and -1
The emitted X-rays are converted into electrical signals by seven detectors.
The two electrical signals, ie, the detection data, and these function values are respectively convolved. Then, the seven convolution results are backprojected.

The first example of the present invention corresponds to the function values shown in FIG.
Shown in Figure 4. FIG. 14 shows 19 functions used in one convolution operation. Two function values are used between each function compared to the conventional example. Same as before 7
The 19 functions are composed of data with two zeros inserted between each of the data from the two detectors. The resulting data is 19, and these are back-projected. The number of data used for back projection addition is 19, which is a large number. As mentioned above, if the data is in a valley, the same image can be reconstructed if the opposing beam data is present there, so a single tomographic image can be precisely obtained.

The present invention is not limited to the above-described embodiments, and can be used with many modifications.

For example, when inserting zeros, two bare zeros are arranged alternately with the actual rays, but any number of zeros may be inserted as long as it is one or more.

Next, an embodiment of dynamic scanning using the present invention will be described with reference to FIG.

First, the tomographic image on the first screen is reconstructed from the projection data of rotary brokers 1 to 6912 using the zero insertion force method according to the present invention. Next, the second screen tomographic image after 1/6912 seconds is the 6913th projection (rotation position 1
The first projection data (t1 = 0 seconds even at rotational position 1) from the data of t1 = 1 + 1/6912 seconds)
Convolution and back projection calculations are performed by using the subtracted data as projection data and inserting zero between the subtracted projection data.

Similarly, the second projection data is subtracted from the 6914th projection data and reconstructed to obtain a third tomographic image. Thereafter, calculations are performed for each brodieximin in the same manner.

If you want to go back in time, you can reverse the projection difference.

By performing as described above, both temporally continuous CT images can be obtained with fewer convolution operations and pack projection operations. Therefore, the next tomographic image can be obtained with a very small number of calculations, that is, in a short time, so that, for example, it is possible to dynamically observe the state of temporal movement of the molding until it is injected into the subject.

[Effects of the Invention] As explained above, the CT wave device of the present invention can obtain high-resolution images without converting source fan data to detector fans, and dynamic scan images can be continuously reduced. It can be obtained by simple calculation.

[Brief explanation of drawings]

Fig. 1 shows how data is arranged when convolution calculation is performed for source fan reconstruction according to the present invention, and Fig. 2 shows how virtual rays are inserted into the rays for source fan reconstruction according to the present invention. 3 is a configuration diagram showing the fourth generation CT wave device according to the present invention, FIG. 4 is a diagram showing the case where reconstruction is performed using a source fan from the fourth generation CT wave device according to the present invention, and FIG. The figure is a diagram when reconstructing with a detector fan from a conventional 4th generation CT wave device.
The figure is a diagram for continuously performing the reconstruction of Figure 5, Figures 7 to 13 are diagrams for explaining the principle of the CT wave device according to the present invention, and Figure 14 is a diagram for explaining the principle of the CT wave device according to the present invention. Convolution (lO
FIG. 15, which is a diagram showing an example of the number, is a diagram showing a conventional convolution function. 1...Detector ring 2...X-ray source 3...Object

Claims (2)

[Claims] (1) A radiation source that generates fan-beam-shaped radiation toward a cross-sectional area of the subject; and the fan, which is made up of a plurality of radiation detection elements arranged so as to surround the subject, and that transmits through the subject. radiation detection means for detecting beam-shaped radiation; scanning means for rotating the radiation source around the subject; and source fan data sequentially supplied from the radiation detection means in accordance with rotation of the radiation source. a reconstruction means for inserting a zero between each data, performing convolution using a convolution function corresponding to the source fan data into which the zero has been inserted, and backprojecting the convolved source fan data; An X-ray CT apparatus comprising: display means for displaying an image of the cross-sectional portion of the subject reconstructed by reconstruction means. (2) Continuously rotate the radiation source, find the difference value between temporally different source and fan data detected at the same rotation angle position, perform convolution on this difference value, and 2. The X-ray CT apparatus according to claim 1, wherein the X-ray CT apparatus back-projects the acquired data onto the previous image.