WO2013191001A1 - X線ct装置 - Google Patents
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- WO2013191001A1 WO2013191001A1 PCT/JP2013/065639 JP2013065639W WO2013191001A1 WO 2013191001 A1 WO2013191001 A1 WO 2013191001A1 JP 2013065639 W JP2013065639 W JP 2013065639W WO 2013191001 A1 WO2013191001 A1 WO 2013191001A1
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Definitions
- the present invention relates to an X-ray CT apparatus, and more particularly to a technique for improving spatial resolution and improving subject measurement accuracy.
- An X-ray CT (Computed Tomography) apparatus photographs X-ray transmission data of a subject while rotating a pair of an X-ray source and an X-ray detector (hereinafter referred to as a scanner) opposed to each other across the subject. It is an apparatus for computationally reconstructing images (hereinafter referred to as CT images), and is widely used in the fields of industrial and security inspection apparatuses, medical image diagnostic apparatuses, and the like. In the field of medical X-ray CT apparatuses, in recent years, the area of X-ray detectors has increased and the speed of scanner rotation has increased, and a wide range of imaging regions can be measured in a short time.
- the detection element of the X-ray detector In order to improve the spatial resolution in the measurement by the X-ray CT apparatus, it is usually necessary to make the detection element of the X-ray detector finer, that is, to reduce the size. However, when the X-ray dose incident on the X-ray detector is the same, if the detection element is miniaturized, the number of X-ray photons incident on one detection element is reduced, and the S / N of the detection signal is reduced. In order to improve the S / N, it is necessary to increase the X-ray dose, but in the case of medical measurement, an increase in the X-ray dose is accompanied by an increase in the exposure of the subject.
- the size of the detection element of the X-ray detector is determined by the trade-off between the spatial resolution and the exposure dose.
- an X-ray input surface with a size of about 1 mm square is usually used.
- An X-ray device is used.
- Patent Document 1 As a method for improving the spatial resolution without reducing the size of the detection element of the X-ray detector, the method of Patent Document 1 has been proposed.
- the method disclosed in Patent Document 1 is a method called Flying Focal Spot (FFS) method.
- FFS Flying Focal Spot
- Patent Document 2 in order to achieve high resolution and high sensitivity of a medical X-ray CT apparatus, one of the linear projection images of an X-ray shielding member of an X-ray grid on the detection surface of the X-ray detector is disclosed.
- a configuration example is disclosed in which the period in the direction is substantially an integer multiple of twice or more the period of the array of detection elements in this one direction.
- FIG. 30 is a diagram for explaining the relationship between the X-ray generation point S and the data sample point R in a general X-ray CT apparatus that does not employ the FFS method.
- FIG. 31 is a diagram for explaining the relationship between the X-ray generation points S1 and S2 and the respective data sample points R1 and R2 in the X-ray CT apparatus adopting the FFS method.
- the X axis is a straight line that passes through the rotation axis of the scanner and is parallel to the input surface of the X-ray detector 2, and the individual detection elements P constituting the X-ray generation point S and the X-ray detector 2.
- the data sample point R is the intersection of the straight line connecting X and the X axis.
- the X-ray beam emitted from the X-ray generation point S and incident on the detection element P is measured after passing through the subject at the data sample point R near the rotation axis of the scanner.
- the relative position between the X-ray generation point S and the X-ray detector 2 rotates the scanner. However, it is always fixed during a plurality of shootings.
- the position of the X-ray generation point is alternately changed to S1, S2, S1, S2,. .
- the positions of S1 and S2 are set so that the data sample points R1 and R2 formed for the respective X-ray generation points are equally spaced in the X-axis direction.
- the sampling interval in the X-axis direction of the data sampling points R1 and R2 obtained in the FFS method is ⁇ X / 2 that is half of the interval ⁇ X in the X-axis direction of the data sampling point R obtained in the general method.
- FIGS. 32A and 32B are diagrams for explaining frequency characteristics in the X-axis direction of data acquired at the data sample points R shown in FIGS. 30 and 31, respectively. That is, FIG. 32A shows the frequency characteristic of the general method, and FIG. 32B shows the frequency characteristic of the FFS method.
- a frequency characteristic 3200 shown in FIG. 32 is a frequency characteristic of the aperture function of the X-ray detection element P. The frequency information included in the subject is measured by being subjected to band limitation by the frequency characteristic 3200 of the aperture function.
- a so-called aliasing region 3201 in which a frequency characteristic 3200 repeatedly appears at every frequency twice as high as the Nyquist frequency fn by a known sampling theorem is compared. It becomes big.
- the aliasing causes high frequency information contained in the subject to be lost and causes a reduction in spatial resolution.
- the Nyquist frequency is 2fn, which is twice that of the general method, and therefore the frequency characteristic 3200 appears repeatedly every 4fn. Accordingly, the aliasing area 3202 formed at this time is significantly reduced compared to the aliasing area 3201 of the general method.
- loss of high frequency information due to aliasing can be reduced, so that the spatial resolution can be improved without reducing the size of the detection element.
- An object of the present invention is to provide an X-ray CT apparatus that improves spatial resolution without reducing the S / N of a CT image to be measured and enables high-precision measurement of a fine structure of a subject.
- an X-ray CT apparatus for generating a CT image of a subject, the X-ray generation unit and an X-ray for detecting an X-ray image irradiated from the X-ray generation unit
- a detection unit a rotation mechanism unit that rotates the X-ray generation unit and the X-ray detection unit around a rotation axis, a signal reading unit that reads a signal detected by the X-ray detection unit at a predetermined imaging timing, and a signal reading unit
- a signal processing unit that generates a CT image of the subject, and a signal reading unit is provided in a plurality of detection pixels configured in a matrix on the X-ray input surface of the X-ray detection unit.
- NX detection pixels adjacent in the X direction and / or Nz detection pixels adjacent in the z direction (where NX and Nz are integers of 1 or more, and one of them is 2 or more) And add the output
- An X-ray CT apparatus configured to change an addition position of a detection pixel to be added in synchronization with imaging timing is provided.
- the spatial resolution can be improved without reducing the S / N of the measured X-ray CT image.
- FIG. 1 is a schematic front view of an X-ray CT apparatus according to a first embodiment. It is a perspective view for demonstrating the positional relationship of the X-ray focus S of an X-ray tube and an X-ray detector based on a 1st Example. It is sectional drawing for demonstrating the cross-section of the X direction of the X-ray detector based on a 1st Example. It is a figure for demonstrating arrangement
- FIG. 32 is a diagram for explaining frequency characteristics in the X-axis direction of data acquired at the data sample point R shown in FIGS. 30 and 31. It is a figure for demonstrating the relationship between the X-ray generation point S and the data sample points R1 and R2 in one structure of the X-ray CT apparatus of this invention.
- the X direction refers to the X-ray generation unit and the X-rays of a large number of detection pixel detection pixels arranged in a matrix on the X-ray input surface of the X-ray detection unit constituting the X-ray CT apparatus. It is a direction perpendicular to the rotation axis for rotating the line detection unit, and the z direction is a direction parallel to the rotation axis of the detection pixel.
- FIG. 33 is a diagram for explaining the relationship between the X-ray generation point S and the data sample points R1 and R2 in the principle configuration of the X-ray CT apparatus of the present invention.
- the detection elements P1, P2, P3,... Constituting the X-ray detector 2 have respective X-direction sizes that are used in the general method shown in FIG. It is designed to be half the size of P in the same direction.
- a signal addition circuit (not shown) is provided, which adds and outputs signals detected by two detection elements adjacent in the X direction. Further, the addition position of the signal addition is changed for each photographing timing.
- signal addition is performed in the form of P1 + P2, P3 + P4,... (Hereinafter referred to as addition pattern 1) at a certain shooting timing, and P2 + P3, P4 + P5,.
- addition pattern 1 is performed in the form of addition pattern 2), and shooting is performed while repeating these two addition patterns alternately.
- the data sample points for the addition patterns 1 and 2 are R1 and R2, respectively, and data having the same sample interval ⁇ X / 2 as in the FFS method shown in FIG. 31 can be acquired.
- the Nyquist frequency of the data sample can be expanded to 2fn, which is twice that of the general method, and the reduction in spatial resolution due to aliasing can be reduced.
- the X-ray CT apparatus of the present invention can perform imaging while fixing the position of the X-ray generation point S in the same manner as in the general method. Occurrence can be prevented. Further, unlike the FFS method, it is not necessary to use a special X-ray tube having a mechanism for moving the position of the X-ray generation point, so that the running cost of the X-ray CT apparatus can be reduced.
- the first embodiment is an embodiment of an X-ray CT apparatus, which is an X-ray CT apparatus that generates a CT image of a subject, and an X-ray generation unit and an X-ray image irradiated from the X-ray generation unit.
- An X-ray detection unit to detect, a rotation mechanism unit that rotates the X-ray generation unit and the X-ray detection unit around a rotation axis, and a signal reading unit that reads out a signal detected by the X-ray detection unit at a predetermined imaging timing;
- a signal processing unit that processes a signal output from the signal reading unit and generates a CT image of the subject, and the signal reading unit is configured in a matrix on the X-ray input surface of the X-ray detection unit.
- NX and Nz are integers equal to or greater than 1 and either 1 And add it to 2 or more
- FIG. 1 is a schematic front view of the X-ray CT apparatus according to the first embodiment.
- the horizontal direction, the vertical direction, and the vertical direction on the paper surface are the X, y, and z directions, respectively.
- An X-ray CT apparatus according to the present embodiment includes an X-ray tube 1 that is an X-ray generation unit, an X-ray detector 2 that is an X-ray detection unit, a rotation plate 3 that is a rotation mechanism unit, a bed top plate 4, and a gantry 5. , Line quality filter 7, bow tie filter 8, collimator 9, console CSL, imaging controller CTL, computer CPU, memory MEM1, memory MEM2, and monitor MNT.
- the X-ray tube 1, the quality filter 7, the bow tie filter 8, the collimator 9, and the X-ray detector 2 are disposed on the rotating plate 3, and these are hereinafter collectively referred to as a rotational imaging system.
- the entire rotational imaging system is stored inside the gantry 5.
- An opening 6 is provided at the center of the gantry 5, and a subject 10 is disposed near the center of the opening 6.
- the X-ray detector 2 in this embodiment is shown in FIG. 1 in addition to the function as the X-ray detector described above. Has the function of a signal reading unit in which is omitted.
- a human body is assumed as the subject 10, and the normal subject 10 is arranged on the bed top 4 in a normal state.
- the rotating plate 3 that is the rotating mechanism unit described above is rotated by a drive motor (not shown) to capture an X-ray transmission image of the subject 10 from the entire circumference.
- the rotating plate 3 rotates around a rotation axis that passes through the center of the opening 6 and is parallel to the z-axis.
- the position of the couchtop 4 can be moved in the z direction by a driving device (not shown).
- a known spiral scan can be performed by simultaneously rotating the rotating plate 3 and moving the bed top 4.
- a representative example of the distance between the X-ray generation point of the X-ray tube 1 and the X-ray input surface of the X-ray detector 2 is 1040 [mm].
- a typical example of the diameter of the opening 6 is 650 [mm].
- a typical example of the rotational speed of the rotating plate 3 is 3 [rotation / second], and the rotational imaging system captures X-ray transmission images of the subject 10 from various rotational angles.
- a typical example of the number of times of photographing in one rotation of the rotating photographing system is 1600 times, and every time the rotating plate 3 rotates 0.225 degrees, one photographing is performed while synchronizing.
- the wire quality filter 7 is a publicly known filter configured by superimposing a single material or a plurality of metal plates.
- the X-ray filter 7 is arranged in the path of the X-ray beam irradiated from the X-ray tube 1 toward the X-ray detector 2, and the X-ray quality (energy spectrum) after passing through the X-ray filter 7 is measured. It has a function to change. In particular, it is used for the purpose of reducing the exposure of the subject 10 by blocking low-energy X-rays or reducing the influence of the BH effect.
- Typical examples of the metal plate used for the wire quality filter 7 include a copper plate having a thickness of about 0.05 to 0.2 mm, an aluminum plate having a thickness of about several mm, or a laminate of these.
- a plurality of types of the quality filter 7 are prepared, and the user can change the above types according to the imaging application. At this time, the designated quality filter 7 is placed in the path of the X-ray beam prior to imaging by a moving mechanism (not shown).
- the bow tie filter 8 is a known one formed of a material such as aluminum.
- the bow tie filter 8 is disposed in the path of the X-ray beam irradiated from the X-ray tube 1 toward the X-ray detector 2.
- the bow tie filter 8 has a shape in which the thickness thereof changes so that the transmission path length of the X-ray beam in the bow tie filter 8 is the shortest at the center position of the opening 6 and becomes longer as it approaches the peripheral position. Yes.
- the intensity of the X-rays that enter the X-ray detector 2 after passing through the subject 10 is made uniform in the Xy plane direction.
- the bow tie filter 8 is prepared in a plurality of shapes according to the size of the subject 10 and the imaging region, and the user can change the type. At this time, the designated bow tie filter 8 is placed in the path of the X-ray beam prior to imaging by a moving mechanism (not shown).
- the collimator 9 is a known X-ray shielding plate formed of a material such as lead, and limits the irradiation range of X-rays emitted from the X-ray tube 1 in the Xy plane direction and the z direction.
- the irradiation range of the X-ray in the Xy plane direction is limited to coincide with the input range of the X-ray detector 2 in the Xy plane direction.
- the irradiation range in the z direction (hereinafter referred to as slice width) can be variously changed by the user in accordance with the photographing purpose.
- a moving mechanism (not shown) moves the position of the collimator 9 to limit the slice width to a specified size.
- the X-ray detector 2 is a well-known one composed of a scattered radiation removing collimator 300, a scintillator array 303, a photodiode array 304, and the like which will be described later.
- the X-ray detector 2 has a two-dimensional input surface in which a large number of X-ray detection elements are arranged in a matrix, and is arranged so that the input surface faces the X-ray tube 1.
- a typical example of the number of arrays is 2000 elements (Xy plane direction) ⁇ 128 elements (Z direction).
- the X-ray detection elements are arranged on an arc that is substantially equidistant in the Xy plane direction with respect to the X-ray tube 1.
- a typical example of the size in the Xy plane direction and the z direction of each X-ray detection element is 0.5 [mm].
- the user instructs shooting start after instructing shooting conditions through the console CSL.
- the imaging conditions include known conditions such as the tube voltage and tube current of the X-ray tube 1, the rotational speed of the rotating plate 3, the imaging slice width, the type of the quality filter 7, the type of the bow tie filter 8, and the imaging range of the subject 10.
- the photographing controller CTL starts to rotate the rotating plate 3.
- the imaging controller CTL instructs the X-ray irradiation of the X-ray tube 1 and the operation of the X-ray detector 2 to start imaging.
- the X-ray detector 2 generates and outputs imaging data while performing a pixel addition operation peculiar to the X-ray CT apparatus according to the designated value of the imaging mode condition described above. The details of the pixel addition operation of the X-ray detector 2 will be described later.
- the imaging data output from the X-ray detector 2 is temporarily transferred to the memory MEM1 after being transferred from a rotating imaging system to a stationary system, that is, a non-rotating system in the gantry 5 via a known optical slip ring (not shown).
- the computer CPU sequentially performs a preprocessing operation described later on the photographic data, and stores the result in the memory MEM2.
- the reference data stored in the MEM 2 in advance is referred to by the computer CPU. The details of the preprocessing calculation and the details of the reference data will be described later.
- the computer CPU reads out the pre-processed imaging data from the memory MEM2, reconstructs the CT image of the subject 10 using a known reconstruction algorithm, and stores the obtained CT image data in the memory MEM2. save.
- the computer CPU also reads out the CT image data MEM2 and uses a known image processing technique such as a volume rendering method, MIP (MaXimum Intensity Projection) method, MPR (Multi Planar Reconstruction) method, or the like.
- a display image is created, and the created display image is displayed on the screen of the monitor MNT.
- a dedicated arithmetic unit or a known general-purpose arithmetic unit is used as the computer CPU.
- the memory MEM1 uses a RAM (Random Access Memory), and the memory MEM2 uses a known recording means such as a hard disk, SSD (Solid State Drive), or a combination thereof.
- FIG. 2 is a perspective view for explaining the positional relationship between the X-ray focal point S of the X-ray tube 1 and the X-ray detector 2 of the present embodiment.
- a large number of X-ray detection elements are arranged in a matrix on the input surface of the X-ray detector 2.
- a unit detection area of the X-ray detector 2 formed by each X-ray detection element is referred to as a detection pixel.
- X and z represent the position of the detection pixel, where X is a direction perpendicular to the Z axis and z is a direction parallel to the Z axis.
- FIG. 3 is a cross-sectional view for explaining a cross-sectional structure in the X direction of the X-ray detector 2 of the present embodiment.
- the X-ray detector 2 includes a scattered radiation removing collimator 300, a scintillator array 303, a photodiode array 304, and the like.
- the scintillator array 303 includes a scintillator block 301 and a light reflecting material 302.
- the scintillator block 301 is made of a known scintillator material that converts light into X-rays, and is arranged in a matrix in the X direction and the z direction.
- the gaps and upper surfaces between the scintillator blocks 301 are filled with a light reflecting material 302 formed of a known material.
- the light reflecting material 302 has a function of reflecting the light generated inside the scintillator block 301 at the interface between the scintillator block 301 and the light reflecting material 302, and the light enters the adjacent scintillator block 301 and is caused by crosstalk. This is used to prevent a reduction in spatial resolution.
- the light reflecting material 302 has a function of preventing the attenuation of the optical signal due to the crosstalk or the emission of light from the upper surface of the scintillator block 301.
- the light reflecting material 302 has a function of an adhesive that bonds and fixes the scintillator block 301.
- a scattered radiation eliminating collimator 300 made of a known material is disposed on the upper surface of the scintillator array 303.
- the scattered radiation elimination collimator 300 has a lattice-like shape when observed from above in the figure perpendicular to the Xz plane. For this reason, the X-direction and z-direction cross-sectional views are shown in FIG. It has a shape that is arranged in large numbers.
- the slit 307 is arranged so as to jump one above the gap in the X direction of the scintillator blocks 301a and 301b.
- the present invention is not limited to this. It may be arranged above, two jumps, three jumps, or the like.
- the crosstalk generated between the gaps is set by setting the gap distance g2 as large as possible in the blocking area. Can be reduced.
- the use efficiency of X-rays can be prevented from being lowered by setting the distance g1 of the gap to be relatively small.
- the slits 307 are arranged so as to be directed to the X-ray focal point S shown in FIG.
- the scattered radiation eliminating collimator 300 is scattered in the subject 10 and the like, and then the scattered X-rays input to the X-ray detector 2 are in the X direction and the z direction with respect to the input surface of the X-ray detector 2.
- the photodiode array 304 is a known back-illuminated photodiode array formed on a silicon substrate, and photodiode elements 305 are formed in an array in the Xz plane direction in the substrate.
- the photodiode elements 305 are designed so that the scintillator blocks 301a and 301b and the positions in the Xz plane direction coincide with each other.
- Light generated inside each of the scintillator blocks 301a and 301b is incident on the photodiode element 305 from the lower surface thereof, converted into an electrical signal, and then output through the output signal line 306.
- the cross-sectional structure in the X direction of the X-ray detector 2 has been described above, the cross-sectional structure in the z direction is the same as that in the X direction, and the description thereof is omitted.
- FIG. 4 is a diagram for explaining the arrangement of the detection pixels P (X, z) viewed from the input surface side of the X-ray detector 2 in the X-ray CT apparatus of the present embodiment.
- the symbol P (X, z) is displayed only on a representative one of a plurality of detection pixels for simplification of the drawing.
- the detection pixels P (X, z) are arranged in a matrix in the X direction and the z direction.
- Each detection pixel P (X, z) is a unit detection region formed individually by a single element pair of the scintillator block 301 and the photodiode element 305 shown in FIG.
- FIG. 5 is a diagram for explaining a pixel addition operation unique to the X-ray CT apparatus of the present embodiment.
- the plurality of detection pixels P (X, z) of the X-ray detector 2 are used.
- the addition position Q of the pixel addition is characterized in that it changes synchronously for each of the plurality of shooting timings.
- FIG. 5 shows an example in which the pixel addition is performed by adding two detection pixels in the X and z directions, for a total of four detection pixels, but it is needless to say that the present invention is not limited to this.
- FIG. 5 shows an example in which the addition position Q for pixel addition is shifted by one detection pixel in the X and z directions at each photographing timing. At this time, there are a total of four different variations of the addition position Q, and each addition position Q (T 1 ) to Q (T 4 ) (where T 1 to T 4 represent four different shooting timings). Are shown in FIGS. 5A to 5D, respectively. All the squares of the thick line shown in FIG.
- FIGS. 5A to 5D represent the addition position Q, and it is assumed that the detection pixel P (X, z) inside the area surrounded by the thick line is added. Also, in FIGS. 5A to 5D, the symbols Q (T 1 ) to Q (T 4 ) are added to only one representative of a plurality of addition positions (thick lines) for simplification of the drawing. It is displayed.
- (A) to (D) in FIG. 5 are merely examples of changes in the addition position corresponding to four different shooting timings T 1 to T 4, and are not limited thereto.
- the addition position Q may be changed to another mode corresponding to the imaging timings T 1 to T 4 .
- the number of pixel additions in the X and z directions is generalized as NX and Nz, respectively.
- the shift amounts in the X and z directions of the addition position are represented as MX and Mz, respectively.
- the number G of variations of the addition position of the pixel addition is expressed by the following equation 1.
- G (UX / MX) * (Uz / Mz) (where 1 ⁇ MX, 1 ⁇ Mz) (Formula 1) However, UX is the least common multiple of NX and MX, and Uz is the least common multiple of Nz and Mz.
- G 4 from Equation 1.
- the pixel addition condition defined by the values of NX, Nz, MX, and Mz as described above is referred to as a shooting mode condition.
- NX or Nz 1
- pixel addition in each direction is not performed (addition number 1).
- MX ⁇ NX and Mz ⁇ Nz are set, oversampling for the added pixels can be realized, and this has the advantage of reducing the reduction in space due to aliasing.
- FIG. 6 is a diagram for explaining the relationship between the imaging timing and the pixel addition position in the X-ray CT apparatus of the present embodiment.
- S (T k ) represents the position of the X-ray generation point at the imaging timing T k .
- the position of the X-ray generation point S sequentially changes on the circle 600 as S (T k ), S (T k + 1 ), S (T k + 2 ),.
- the X-ray detector 2 sets the pixel addition position to Q (T 1 ), Q (T 2 ),... Q (T G ), Q (T 1 ), Q ( T 2 ),...
- a periodic position change is performed for each photographing frame. For example, in the pixel addition example shown in FIG. 5, Q (T 1 ), Q (T 2 ), Q (T 3 ), Q (T 4 ), Q (T 1 ), Q (T 2 ),. ⁇ Repeat four cycles of position changes.
- FIG. 7 is a diagram for explaining another example of pixel addition.
- the symbols Q (T 1 ) to Q (T 3 ) are added to only one representative of a plurality of addition positions (thick lines) for simplification of the drawing. It is displayed.
- the addition position changes with a shift amount (1 pixel) that is 1/3 of the addition number (3 pixels).
- the Nyquist frequency in the X direction increases three times compared to the case where the addition position is not changed, the spatial resolution is reduced by aliasing compared to the example shown in FIG. 5 (Nyquist frequency is increased two times). Can be further reduced.
- FIG. 8 is a diagram for explaining still another example of pixel addition.
- the symbols Q (T 1 ) to Q (T 4 ) are added to only one representative of a plurality of addition positions (thick lines) for simplification of the figure. It is displayed.
- temporal resolution may be prioritized over spatial resolution.
- the number of additions may be set larger as shown in the figure. At this time, the total number of added pixels is reduced and the amount of signal output from the X-ray detector 2 is reduced. Therefore, it is possible to increase the frame rate of imaging and cope with improvement in time resolution.
- FIG. 9 is a diagram for explaining the detector center area 900 and the detector peripheral area 901 set in the X-ray detector 2 of the present embodiment.
- the shooting mode condition may be set individually for each area set as described above, for example.
- the setting of the area in the detector is not limited to this example, and the number and position may be variously changed.
- FIG. 10 is a diagram showing an example of a setting screen that functions as a designation unit for setting the imaging mode condition in the X-ray CT apparatus of the present embodiment.
- This setting screen is displayed, for example, on the monitor MNT in FIG. 1, and various setting values are input from the console CSL using a known information input means such as a keyboard or a mouse.
- a setting screen can be provided on the console CSL.
- a normal shooting mode and a high-resolution shooting mode are prepared as shooting modes, and the user can select a shooting mode through selection of radio buttons 1000 and 1001 according to the shooting purpose.
- input lists 1002 and 1003 for setting values of NX, Nz, MX, and Mz, which are shooting mode conditions, are prepared, and individual values are preset.
- the imaging mode conditions can be preset for both the detector center area 900 and the detector peripheral area 901. Each preset value can be changed by the user.
- the user presses a pull-down button 1004 so that a desired set value can be selected from a selection list 1005 displayed at this time.
- NX, Nz are individually applied to the detector central area and the detector peripheral area, which are a plurality of different small areas set in advance. Etc. can be set.
- the method for setting the shooting mode condition for each area is not limited to this example, and various known methods may be used instead.
- FIG. 11 is a diagram for explaining the arrangement of the detector modules 1100 constituting the X-ray detector 2 in the X-ray CT apparatus of the present embodiment.
- the X-ray detector 2 is formed by arranging a plurality of detector modules 1100 in the X direction.
- this detector module 1100 is called a small detector for convenience. With such a configuration, even when a failure such as a pixel defect occurs in some of the detection pixels of the X-ray detector 2, only the detector module 1100, which is a corresponding small detector, is replaced. Low cost repairs are possible.
- each detector module 1100 is usually used having a planar shape, but these small detectors are placed on a frame on the same arc (not shown) with the X-ray focal point S as the center.
- the X-ray detector 2 having a substantially arc-shaped input surface as shown in FIG. 2 can be formed.
- FIG. 12 is a perspective view for explaining a specific example of the structure of the small detectors constituting the X-ray detector 2, that is, the individual detector modules 1100, in the X-ray CT apparatus of the present embodiment.
- the detector module 1100 includes a scattered radiation elimination collimator 300, a scintillator array 303, a photodiode array 304, a substrate 1200, a substrate 1201, a flexible wiring 1202, a DAS (Data Acquisition System) chip 1203, and the like.
- the DAS chip 1203 functions as the signal reading unit of the above-described embodiment.
- the scintillator array 303 is disposed on the upper surface of the photodiode array 304, and the scattered radiation eliminating collimator 300 is disposed on the upper surface of the scintillator array 303.
- the photodiode array 304 is fixed on the upper surface of the substrate 1200, and the substrate 1201 is fixed on the rear surface of the substrate 1200 so that the substrate surface is vertical.
- the output signal line 306 of the photodiode array 304 shown in FIG. 3 is connected to a flexible wiring 1202 having one end fixed to the back side of the substrate 1200 through a through-through hole (not shown) formed in the substrate 1200. Yes.
- the other end of the flexible wiring 1202 is fixed on the substrate 1201 and connected to the DAS chip 1203 via a terminal (not shown) formed on the substrate 1201.
- the DAS chip 1203 is a circuit that performs AD (Analog-to-Digital) conversion on the electrical signal output from the photodiode array 304 and outputs it as a digital signal, and is formed using a known CMOS technology or the like.
- the outputs of the plurality of DAS chips 1203 corresponding to the plurality of detector modules 1100 are stored in the memory MEM1 as imaging data as described with reference to FIG. An image is generated.
- the structure of the detector module 1100 that is a small detector is not limited to this example.
- the DAS chip 1203 may be directly connected and fixed on a silicon substrate on which the back-illuminated photodiode array 304 is formed using a known technique such as flip chip bonding.
- FIGS. 13 to 15 are circuit diagrams for explaining the outline of the circuit configuration of the DAS chip 1203 which is a signal reading unit in the X-ray CT apparatus of this embodiment.
- Circuits in the DAS chip 1203 which is a signal reading unit mainly include a pixel addition circuit 1300, a CA (Charge Amplifier) circuit array 1301, an SH (Sample Hold) circuit array 1302, and an ADC (Analog-to-Digital Converter) circuit array 1303.
- the CA circuit array 1301, the SH circuit array 1302, and the ADC circuit array 1303 are in the above order from the signal input direction to the output direction (left to right in the drawing) in all the circuit configurations shown in FIGS.
- the positions of the pixel addition circuits 1300 are different from each other.
- the pixel addition circuit 1300 is arranged at the first stage (the previous stage of the CA circuit array) in the example of FIG. 13, and is arranged between the SH circuit array 1302 and the ADC circuit array 1303 in the example of FIG. In this example, it is arranged at the last stage (after the ADC circuit array 1303).
- the CA circuit array 1300 a known charge amplifier that accumulates a charge signal generated in the photodiode array 304 and converts it into a voltage signal is formed in a parallel array for each input signal line.
- the SH circuit array 1302 is configured by forming a known sample hold circuit that samples the output voltage of the charge amplifier at a predetermined timing in a parallel array for each input signal line.
- sampling timing of the sample hold circuit is the imaging timing of the X-ray CT apparatus.
- the ADC circuit array 1303 is a known ADC circuit that converts a voltage signal sampled by the sample hold circuit into a digital signal, and is formed in a parallel array for each input signal line.
- the pixel addition circuit 1300 is a circuit that performs the pixel addition described above.
- NX and Nz can be set as the addition number in pixel addition, but the minimum values NXo and Nzo that can be set as NX and Nz When either one of them is 2 or more, signal addition is always performed in the pixel addition circuit 1300, so the number of output signal lines 1305 of the pixel addition circuit 1300 is reduced with respect to the number of input signal lines 1304. Can do. For this reason, the number of parallel circuits that follow the subsequent stage can be reduced as the pixel addition circuit 1300 is in the upper stage. Such a reduction in the number of circuits has advantages such as reducing power consumption of the circuits to reduce heat generation from the circuits and improving the manufacturing yield of the DAS chip 1203.
- the pixel addition circuit 1300 is arranged at the uppermost stage, there is an advantage that the number of circuits in the subsequent stage can be reduced most. However, on the other hand, there is a demerit that it is most easily affected by circuit noise generated by pixel addition.
- the pixel addition circuit 1300 is arranged in the middle stage, and the CA circuit array 1301 and the SH circuit array 1302 in the previous stage are formed in parallel with the same number as the output signal lines 306 of the photodiode array 304. There is a need to.
- the signals are added after the signals are amplified by the CA circuit array 1301, there is a merit that it is less susceptible to the influence of circuit noise generated during the addition than in the case of FIG. Further, there is an advantage that the circuit amount of the ADC circuit array 1303 having the largest power consumption and heat generation amount can be reduced.
- a known addition circuit that adds the voltage signals output from the S / H circuit array 1302 is used for the pixel addition circuit 1300 in FIG.
- the pixel addition circuit 1300 is arranged at the last stage, and there is a demerit that the circuit amount cannot be reduced.
- the signals are added after being converted into digital signals by the ADC circuit array 1303, there is an advantage that no circuit noise is generated due to the addition.
- a known digital addition circuit is used for the pixel addition circuit 1300 in FIG.
- FIG. 16 is a schematic diagram for explaining the addition position Q of pixel addition performed in each detector module 1100 in the X-ray CT apparatus of the present embodiment.
- the detection pixels P (X, z) arranged in the X direction and the z direction are finite, and therefore when the pixel addition is performed, the number of additions is NX and Nz at the end in each direction. There may be less than one. Therefore, depending on the shooting timing, the number of signals output after addition may change.
- the DAS chip 1203 as a signal reading unit having the pixel addition circuit 1300 is individually prepared for each detector module, and therefore can be processed in parallel in the pixel addition circuit 1300.
- the pixel addition capability and the number of output signal lines 1304 output after the addition must be set to the maximum value of the number of output signals. Now, let the number of detection pixels P (X, z) that each detector module 1100 has in the X direction and the z direction be KX and Kz, respectively.
- G 3 from Equation 1
- the addition positions Q (T 1 ) to Q (T 3 ) at the respective photographing timings are shown in FIGS. 16 (A) to (C), respectively.
- numbers are added in the form of Q 1 (T), Q 2 (T),...
- the maximum number added by the above numbering is 3, and the maximum number of signals output in the X direction is 3. It becomes a piece.
- LX can be calculated by the following equations 2, 3, and 4.
- FIG. 17 is a diagram for explaining another example of pixel addition performed in each detector module 1100.
- G 3 from Equation 1
- the addition positions Q (T 1 ) to Q (T 3 ) at the respective photographing timings are shown in FIGS. 17 (A) to (C), respectively.
- the addition positions are numbered sequentially from the left in the form of Q 1 (T), Q 2 (T),.
- the maximum numbering is four.
- FIG. 18 is a diagram for explaining still another example of pixel addition performed in each detector module 1100.
- G 2 from Equation 1
- the respective addition positions Q (T 1 ) and Q (T 2 ) at each photographing timing are shown in FIGS. 18A and 18B, respectively.
- Q 1 (T) in order from the left side with respect to each addition position, Q 2 (T) are numbered in such a way that the,.
- FIG. 19 is a diagram for explaining an example of the arrangement of the detection pixels P (X, z) in the detector module 1100 in the X-ray CT apparatus of the present embodiment.
- FIG. 20 is a diagram for explaining an example of pixel addition performed in the detector module 1100 shown in FIG.
- FIG. 21 is a circuit diagram for explaining a specific example of the pixel addition circuit 1300 for realizing the pixel addition shown in FIG.
- X and z directions of the detection pixels P (X, z) are represented by numbers 1 to 4, respectively.
- G 2 from Equation 1
- each addition position Q (X, z) at each photographing timing is shown in FIGS. 20 (A) and 20 (B), respectively.
- the pixel addition circuit 1300 shown in FIG. 21 is assumed to be arranged at the first stage of the DAS chip 1203 shown in FIG. 13, and a signal input from the input signal line 1304 is generated by the photodiode array 304.
- switching between two imaging timings T1 (corresponding to the pixel addition pattern in FIG. 20A) and T2 (corresponding to the pixel addition pattern in FIG. 20B) is performed on the signal lines 2101 and 2100, respectively.
- This is realized by alternately switching ON / OFF of the input switch voltages T1 and T2.
- the switch voltage T1 is switched to the switches 2115, 2116, 2117, 2118 via the signal lines 2101 and 2113, And 2119 are added to the gate voltage to turn on all the switches. Therefore, in the X-ray CT apparatus of the present embodiment, the pixel addition position is switched using the imaging timings T1 and T2, which are the sampling timing of the sample hold circuit, and the imaging timing and the pixel addition position switching are synchronized. .
- the signal charges generated in the detection pixels P (1,1), P (2,1), P (1,2), and P (2,2) are the signal lines 2103, 2104, 2105, 2106 and the switch 2115, respectively. 2116, 2117, and 2118 to the signal line 2114 to be added. Further, the addition signal is output as an addition pixel signal Q (1, 1) via the switch 2119.
- the switch voltage T2 is OFF, the switches 2123, 2124, and 2125 are all OFF. Therefore, other signal charges do not flow into the added pixel signal Q (1,1).
- all six added pixel signals from Q (1,1) to Q (3,2) are generated.
- a GND signal is output for them. .
- the GND signal is supplied via signal lines 2102, 2107, and 2109 and output via switches 2121 and 2122.
- the switch voltage T1 input to the signal line 2101 is OFF and the switch voltage T2 input to the signal line 2100 is ON
- the switch voltage T2 is applied to the switches 2123, 2124, and 2125 via the signal lines 2100 and 2112. It is added to the gate voltage to turn on all the switches.
- the signal charges generated in the detection pixels P (1,1) and P (2,1) flow into the signal line 2111 through the signal lines 2103 and 2105 and the switches 2123 and 2124, respectively, and are added.
- the addition signal is output as an addition pixel signal Q (1,1) via the switch 2125.
- FIG. 22 is a flowchart for explaining the procedure of the preprocessing operation in the X-ray CT apparatus of the present embodiment described with reference to FIG. Note that the steps of S01 to S07 shown in FIG. 22 are performed by the calculation of the computer CPU functioning as the signal processing unit shown in FIG. Imaging data that is output from the X-ray detector 2 and then transferred from the rotational imaging system to the stationary system is stored in the memory MEM1. At the same time that the photographic data is stored in the memory MEM1, the computer CPU first reads the photographic data from the MEM1 (step S01) and performs an end pixel addition operation (step S02). Here, the contents of the end pixel addition will be described with reference to FIGS.
- FIG. 23 is a schematic diagram for explaining an example of calculation contents of end pixel addition in the X-ray CT apparatus of the present embodiment.
- FIG. 23 shows some of the detector modules 1100a and 1100b when the X-ray detector 2 is configured by arranging the detector modules 1100 shown in FIG. 19 in the X direction.
- the added pixel signal is converted into a digital signal for each detector module 1000 and output.
- an addition pixel output less than the set addition number NX is generated in the addition pixels near the module boundary 2300.
- FIG. 23 shows some of the detector modules 1100a and 1100b when the X-ray detector 2 is configured by arranging the detector modules 1100 shown in FIG. 19 in the X direction.
- FIG. 23 also shows the addition pixel position Q (X, z) (shooting mode condition: N
- FIG. 24 is a diagram for explaining another example of calculation contents of edge pixel addition.
- FIG. 24 shows some detector modules 1100a, 1100b, and 1100c when the X-ray detector 2 is configured by arranging the detector modules 1100 shown in FIG. 19 in the X direction.
- G 3 from Equation 1, and there are three shooting timings.
- end pixel addition is performed on the addition pixel positions Qa (2,1) and Qb (1,1) near the module boundary 2300.
- step S03 a reference correction calculation is performed (step S03).
- the reference correction is a known calculation performed to normalize the output of the shooting data, and is calculated by the following equation (5).
- g is the photographing timing
- VX and Vz are the numbers of added pixels in the X and z directions generated after the end pixel addition
- Q g (X, z) is the signal value of the added pixel.
- the crosstalk correction is an operation for recovering the spatial resolution reduced by the crosstalk by removing the crosstalk signal flowing from the adjacent detection pixel by calculation, and various calculation methods have been proposed so far. For example, as one of the simplest known calculation methods, there is a method shown in Equation 6.
- crosstalk ratio values derived in advance based on known experiments, known simulations, and the like are tabulated and recorded in the memory MEM2.
- the table value is read from the memory MEM2 at the time of crosstalk correction, and is used for the calculation by the above equation (6).
- FIG. 25 is a diagram for explaining a configuration of a crosstalk ratio table stored in the memory MEM2 in the X-ray CT apparatus according to the present embodiment.
- the value of the crosstalk ratio needs to be obtained for all the added pixels generated in various shooting modes.
- the table of the crosstalk ratio is in the form of a table 2500 for the imaging mode 1, a table 2501 for the imaging mode 2, a table 2502 for the imaging mode 3,.
- step S04 air correction
- step S05 air correction
- BH Beam Hardening
- a g (X, z) is a known air data measured in advance to the addition pixel Q g (X, z).
- the edge pixel addition (step S02), the reference correction (step S03), and the crosstalk correction already described with reference to FIG. 22 are performed with respect to the photographing data photographed without the subject 10 and the couch top 4.
- Step S04 After executing (Step S04), a plurality of these are averaged.
- ⁇ g (X, z) and ⁇ g (X, z) are well-known BH correction coefficients measured in advance for the added pixel Q g (X, z).
- a publicly known method as disclosed in Japanese Patent Publication No. 61-54412 Patent Document
- Patent Document Patent Document
- the air data A g (X, z) and the BH correction coefficients ⁇ g (X, z) and ⁇ g (X, z) are tabulated in advance and recorded in the memory MEM2.
- the table values are read from the memory MEM2 at the time of air correction and BH correction, respectively, and are used for calculations according to the above formulas 7 and 8.
- FIG. 26 is a diagram for explaining a specific configuration of a table of air data and BH correction coefficients stored in the memory MEM2 in the X-ray CT apparatus according to the present embodiment.
- the values of the air data and the BH correction coefficient are all generated under various imaging conditions such as the tube voltage of the X-ray tube 1, the type of the quality filter 7 and the combination condition that can be set for the type of the bow tie filter 8. It is necessary to obtain the sum pixel.
- the air data and BH correction coefficient tables are all settable shooting conditions in the form of a table 2600 for shooting conditions 1, a table 2601 for shooting conditions 2, a table 2602 for shooting conditions 3, and so on. Prepared for.
- tables are prepared for all settable shooting modes in the form of a table 2603 for shooting mode 1, a table 2604 for shooting mode 2, and so on. Further, in each shooting mode, a table is prepared in the form of a table 2605 for shooting timing T1, a table 2606 for shooting timing T2,... For all G shooting timings in the shooting mode.
- step S06 the data after the BH correction (step S06) is finally stored in MEM2 (step S07).
- the pre-processing calculation process from step S01 to S07 is repeatedly performed every time the imaging data output from the X-ray detector 2 is stored in the memory MEM1 during imaging, and the calculation is performed on all imaging data. When the processing is finished, the preprocessing calculation is finished.
- FIG. 27A is a flowchart for explaining a procedure of an example of the CT image reconstruction operation described with reference to FIG. 1 in the X-ray CT apparatus of the present embodiment. Note that the steps S11 to S14 shown in FIG. 27A are performed by the calculation of the computer CPU.
- Various methods for reconstructing algorithms have been proposed and are publicly known, but in the following, the most general Feldkamp algorithm (known document 1: J.Opt. Soc. Am. A, vol.1). , pp. 612-619, June) 1984), the procedure of reconstruction calculation will be described. However, as will be described later, the present invention is not limited to this reconstruction calculation.
- the computer CPU first reads the photographic data from MEM2 (step S11) and performs a filter operation (step S12). Since the filter calculation method is described in the above-mentioned publicly known document 1, a detailed description thereof is omitted here, but simply speaking, it is an operation for convolving a digital filter called a reconstruction filter in the X direction with respect to image data. is there. In this X-ray CT apparatus, the filter operation is performed on the data after the pixel addition.
- the filter calculation is performed in the same manner as the conventional method regardless of the position change.
- the back projection calculation is performed using the photographing data after the completion of the filter calculation (step S13). Similar to the filter calculation, the back projection calculation basically uses the method described in the above-mentioned publicly known document 1.
- a method of creating back projection data used for the back projection calculation Details of the backprojection calculation in the X-ray CT apparatus of the present embodiment will be described below with reference to FIGS. 28 and 29.
- FIG. 28 is a diagram for explaining a calculation method of back projection calculation.
- the CT image is generated by calculating reconstruction data for all reconstruction points R on the XYZ coordinates constituting the CT image.
- a position where the straight line 2800 connecting the X-ray generation point S (T k ) and the reconstruction point R at the imaging timing T k intersects the X-ray detector 2 ( X, z) (hereinafter referred to as a projection position) is calculated, and later-described shooting data interpolation calculation (however, the shooting data here is the shooting data after the filter calculation shown in step S12 of FIG. 27A).
- the backprojection data includes PJ k (X, z), PJ for all X-ray generation point positions S (T k ), S (T k + 1 ),. k + 1 (X, z),..., and by adding these, the reconstruction data of the reconstruction point R can be calculated.
- the above calculation for adding backprojection data is called backprojection calculation.
- FIG. 29 is a diagram for explaining a backprojection data calculation method based on an interpolation calculation. Since there are various projection positions (X, z) of the back projection data with respect to the arrangement of the addition pixel positions Q on the X-ray detector 2, the value of the back projection data PJ k (X, z) is the projection position. It is calculated by interpolating the shooting data measured at the addition pixels around (X, z). At this time, the position of the addition pixel changes to Q (T k ) (see FIG. 29A), Q (T k + 1 ) (see FIG. 29B), etc. at each photographing timing. Interpolation data is created in consideration of the change in the position of.
- step S13 the CT image created by the back projection operation (step S13) is finally stored in MEM2 (step S14), and the reconstruction operation is terminated.
- the detection pixels of the X-ray detector are pixel-added and the pixel addition position is changed at every imaging timing.
- the Nyquist frequency can be improved by reducing the data sampling interval.
- the amount of high spatial resolution information included in the subject lost due to aliasing can be reduced, so that the spatial resolution of the CT image can be improved.
- the size of the detection pixel in the X direction and the z direction is set to 0.5 mm [mm]. This is the same size as that used in a general medical X-ray CT apparatus. About half the size. In this case, the amount of X-rays input to one detection pixel is 1/4 of the normal, and the S / N of the measured CT image is reduced.
- the size of the detection pixel is halved in the X direction and the y direction as described above, the total number of detection pixels increases four times. Therefore, if a circuit for reading signals for all the detection pixels is provided in parallel, There arises a problem that the number of parallel circuits and the number of output data increase four times. Such an increase in the number of parallel circuits not only increases the cost and power consumption of the X-ray detector 2, but also increases the amount of heat generated from the circuit and causes a malfunction. Further, when the number of output data is increased by a factor of 4, the amount of information transferred from the rotary imaging system to the stationary system via the optical slip ring is increased by a factor of 4, and malfunctions such as transfer errors are likely to occur. However, in the X-ray CT apparatus shown in the present embodiment, by providing the pixel addition circuit 1300, the above high spatial resolution measurement can be realized with almost no increase in the number of parallel circuits and the number of output data.
- a signal readout circuit can be configured with a slight increase in the number of parallel circuits of about 8% with respect to the total number of 1024 parallel circuits included in the general detector module. Therefore, the high resolution of the X-ray detector 2 can be increased without causing an increase in the cost of the X-ray detector 2, an increase in power consumption, and a malfunction due to an increase in the amount of heat generated.
- An X-ray CT apparatus capable of measurement can be realized.
- the Feldkamp algorithm is used as the reconstruction algorithm
- other known algorithms may be used.
- imaging data measured in a fan beam type coordinate system is converted into a parallel beam type coordinate system, and then reconstruction is performed by filter calculation and back projection.
- fan beam data interpolation is performed in order to obtain desired parallel beam data.
- PJ k X
- the addition pixel Q T k
- parallel beam data corresponding to the change in the addition pixel position is obtained in the same manner as the interpolation calculation method shown in FIG. Can be calculated.
- the reconstruction flow in this case is a configuration in which after shooting data is read (step S11), fan parallel conversion (step S15) is performed, and then a filter operation is performed.
- the X-ray CT apparatus of the present invention can improve the spatial resolution without reducing the S / N of the measured X-ray CT image.
- industrial CT it is possible to measure the fine structure of a subject without increasing the X-ray output from the X-ray generator to improve the S / N, thereby extending the life of the X-ray generator and reducing the life.
- the measurement accuracy of the subject can be improved.
- medical CT it is possible to measure a finer structure such as a blood vessel without increasing the exposure of the subject to improve the S / N, and the diagnostic ability can be improved, which is extremely useful.
- detector central area 901 ... detector peripheral area, 1100... Detector module, 1200... Substrate 1201... Substrate 1202 ... Flexible wiring, 1203 ... DAS (Data Acquisition System) chip, 1300 ... Pixel addition circuit, 1301... CA (Charge Amplifier) circuit array, 1302... SH (Sample Hold) circuit array, 1303 ... ADC (Analog-to-Digital Converter) circuit array, 1304 ... Input signal line, 1305: Output signal line.
- CA Charge Amplifier
- SH Sample Hold
- ADC Analog-to-Digital Converter
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Abstract
Description
ただしUXをNXとMXの最小公倍数、UzをNzとMzの最小公倍数とする。
LX=JX+1 (IX=0, またはIX=1、またはIX=MXかつNXがMXの倍数の場合)
LX=JX+2 (上記以外の場合) ・・・(式2)
(2)NXが1の場合
LX=KX ・・・(式3)
(3)MXが0の場合
LX=JX (IX=0の場合), LX=JX+1(IX≠0の場合) ・・・(式4)
ただし式2および式4において、KX=NX×JX+IX(IXは0≦IX<NXを満たす整数)とする。
1,1)に他の信号電荷が流入することはない。同様にしてQ(1,1)~Q(3,2)迄の合計6個全ての加算画素信号が生成される。以上のようにしてスイッチ電圧T1およびT2のON・OFFを交互に切り替えることで、加算画素の位置を切り替えることができる。
ただしgを撮影タイミング、VX, Vzをそれぞれ端部画素加算後に生成されるXおよびz方向の加算画素の個数、Qg(X,z)を加算画素の信号値とする。
(g=1,2,…G, X=1,2,…VX, z=1,2,…Vz) ・・・(式6)
ただしag(X,z)、bg(X,z)、cg(X,z)、およびdg(X,z)は、加算画素Qg(X,z)に対してその隣接する加算画素Qg(X-1,z), Qg(X+1,z), Qg(X,z+1),およびQg(X,z-1)からQg(X,z)に混入する信号の比率であり、以下ではクロストーク比率と呼ぶ。
(g=1,2,…G, X=1,2,…VX, z=1,2,…Vz) ・・・(式7)
R’g(X,z) =αg(X,z)×Rg(X,z)2 +βg(X,z)×Rg(X,z)
(g=1,2,…G, X=1,2,…VX, z=1,2,…Vz) ・・・(式8)
ただしAg(X,z)は加算画素Qg(X,z)に対して予め計測しておいた公知のエアデータである。
2・・・X線検出器、
3・・・回転板、
4・・・寝台天板、
5・・・ガントリー、
6・・・開口部、
7・・・線質フィルタ、
8・・・ボウタイフィルタ、
9・・・コリメータ、
10・・・被写体、
CPU・・・計算機、
MEM1・・・メモリ、
MEM2・・・メモリ、
CTL・・・撮影コントローラー、
CSL・・・操作卓、
MNT・・・モニタ、
300・・・散乱線除去コリメータ、
301・・・シンチレータブロック、
302・・・光反射材、
303・・・シンチレータアレイ、
304・・・フォトダイオードアレイ、
305・・・フォトダイオード素子、
306・・・出力信号線、
307・・・スリット、
900・・・検出器中心エリア、
901・・・検出器周辺エリア、
1100・・・検出器モジュール、
1200・・・基板、
1201・・・基板、
1202・・・フレキシブル配線、
1203・・・DAS(Data Acquisition System)チップ、
1300・・・画素加算回路、
1301・・・CA(Charge Amplifier)回路アレイ、
1302・・・SH(Sample Hold)回路アレイ、
1303・・・ADC(Analog-to-Digital Converter)回路アレイ、
1304・・・入力信号線、
1305・・・出力信号線。
Claims (15)
- 被写体のCT画像を生成するX線CT装置であって、
X線発生部と、
前記X線発生部から照射されたX線イメージを検出するX線検出部と、
前記X線発生部と前記X線検出部を回転軸を中心に回転する回転機構部と、
前記X線検出部で検出された信号を所定の撮影タイミングで読み出す信号読み出し部と、前記信号読み出し部より出力された信号を処理し、前記被写体のCT画像を生成する信号処理部とを備え、
前記信号読み出し部は、
前記X線検出部のX線入力面上にマトリクス状に構成された複数の検出画素においてx方向に隣接するNx個の検出画素、および/またはz方向に隣接するNz個の検出画素で検出された信号(ここでNxおよびNzは、1以上の整数でありどちらか1方を2以上とする)を加算して出力するとともに、前記撮影タイミングと同期して、前記加算を行う検出画素の加算位置の変化を行う
ことを特徴とするX線CT装置。 - 請求項1に記載のX線CT装置であって、
前記信号読み出し部は、
前記加算を行う検出画素の位置を周期的に変化させる
ことを特徴とするX線CT装置。 - 請求項1に記載のX線CT装置であって、
前記NxおよびNzを指定する指定部を更に有する
ことを特徴とするX線CT装置。 - 請求項3に記載のX線CT装置であって、
前記信号処理部に接続される表示部を更に備え、
前記指定部は、前記表示部の表示画面で構成される
ことを特徴とするX線CT装置。 - 請求項1に記載のX線CT装置において、
前記X線検出部は、検出領域中に予め設定された複数の小領域を備え、
前記NxおよびNzの値は、前記複数の小領域に対して個別に設定可能である
ことを特徴とするX線CT装置。 - 請求項1に記載のX線CT装置であって、
前記信号読み出し部は、
前記X線検出部からの出力信号を、前記検出画素で検出された信号を加算する画素加算回路、前記検出画素で検出された信号を所定のタイミングでサンプリングするサンプルホールド回路、及び前記検出画素で検出された信号をデジタル信号に変換するADC回路を含む
ことを特徴とするX線CT装置。 - 請求項6に記載のX線CT装置であって、
前記撮影タイミングは、前記サンプルホールド回路の前記所定のタイミングである
ことを特徴とするX線CT装置。 - 請求項7に記載のX線CT装置であって、
前記画素加算回路は、前記サンプルホールド回路の前段、前記サンプルホールド回路の後段、あるいは前記ADC回路の後段に設置される
ことを特徴とするX線CT装置。 - 請求項1に記載のX線CT装置であって、
前記信号読み出し部は、
前記加算位置の変化のx方向およびz方向の位置変化量の最小単位をそれぞれMxおよびMzとし(ただしMxは0≦Mx<Nxの整数、Mzは0≦Mz<Nzの整数とする)、
前記MxおよびMzのどちらか1方を2以上とする、あるいは
前記Mxおよび/またはMzの値は、2種類以上の異なる選択肢の中から指定可能である
ことを特徴とするX線CT装置。 - 請求項9に記載のX線CT装置であって、
前記X線検出部は、
x方向にKx個の前記検出画素を有する小X線検出部を複数配列して構成されており、
前記信号読み出し部は、
Nxが2以上かつMxが1以上の場合にLx個(ただしKx=Nx×Jx+IxでIxは0≦Ix<Nxを満たす整数とし、Ix=0、またはIx=1、またはIx=MxかつNxがMxの倍数の場合はLx=Jx+1とし、上記以外の場合はLx=Jx+2とする)の並列処理能力を有し、前記小X線検出部各々のx方向に割り当てられる
ことを特徴とするX線CT装置。 - 請求項9に記載のX線CT装置であって、
前記X線検出部は、
z方向にKz個の前記検出画素を有する小X線検出部を複数配列して構成されており、
前記信号読み出し部は、
Nzが2以上かつMzが1以上の場合にLz個(ただしKz=Nz×Jz+IzでIzは0≦Iz<Nzを満たす整数とし、Iz=0、またはIz=1、またはIz=MzかつNzがMzの倍数の場合はLz=Jz+1とし、上記以外の場合はLz=Jz+2とする)の並列処理能力を有し、前記小X線検出部各々のz方向に割り当てられる
ことを特徴とするX線CT装置。 - 請求項1に記載のX線CT装置であって、
前記信号処理部は、
前記加算位置の変化のフレーム周期内に形成される全ての加算画素(ただし前記検出画素の加算により形成される大きな画素単位を加算画素とする)に対して該加算画素と該加算画素に隣接する加算画素との間のクロストーク量を用いて、前記クロストークによる空間分解能の低下を低減するための補正演算を行う
ことを特徴とするX線CT装置。 - 請求項1に記載のX線CT装置であって、
前記信号処理部は、
前記加算位置の変化のフレーム周期内に形成される全ての加算画素(ただし前記検出画素の加算により形成される大きな画素単位を加算画素とする)に対して前記CT画像を生成する際に必要なキャリブレーションデータを用いて、前記CT画像を生成する
ことを特徴とするX線CT装置。 - 請求項12に記載のX線CT装置であって、
予め算出した前記クロストーク量を記憶する記憶部を更に備える
ことを特徴とするX線CT装置。 - 請求項13に記載のX線CT装置であって、
予め算出した前記キャリブレーションデータを記憶する記憶部を更に備える
ことを特徴とするX線CT装置。
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Cited By (5)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
JP2016154847A (ja) * | 2015-02-20 | 2016-09-01 | 東芝メディカルシステムズ株式会社 | X線コンピュータ断層撮影装置及びx線検出器 |
JP2019503215A (ja) * | 2016-01-14 | 2019-02-07 | プリズマティック、センサーズ、アクチボラグPrismatic Sensors Ab | X線検出器用の測定回路、ならびに対応する方法およびx線撮像システム |
WO2020153272A1 (ja) * | 2019-01-24 | 2020-07-30 | ソニーセミコンダクタソリューションズ株式会社 | 測定装置、測距装置および測定方法 |
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Families Citing this family (14)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
JP2014226376A (ja) * | 2013-05-23 | 2014-12-08 | 株式会社東芝 | X線ct装置 |
JP2014230600A (ja) * | 2013-05-28 | 2014-12-11 | 株式会社東芝 | X線ct装置およびx線ct装置用x線検出器 |
EP3065642B1 (en) * | 2013-11-08 | 2020-01-08 | Koninklijke Philips N.V. | Empirical beam hardening correction for differential phase contrast ct |
US10092257B2 (en) * | 2014-10-27 | 2018-10-09 | Rensselaer Polytechnic Institute | Dynamic bowtie filter for cone-beam/multi-slice CT |
WO2017052443A1 (en) * | 2015-09-24 | 2017-03-30 | Prismatic Sensors Ab | Modular x-ray detector |
WO2017112623A1 (en) * | 2015-12-22 | 2017-06-29 | Carestream Health, Inc. | Tomographic image acquisition using asymmetric pixel binning |
US10192646B2 (en) * | 2016-04-25 | 2019-01-29 | General Electric Company | Radiation shielding system |
KR20190033621A (ko) * | 2016-08-11 | 2019-03-29 | 프리스매틱 센서즈 에이비 | 컴퓨터 단층촬영을 위한 데이터 획득 |
US11350892B2 (en) * | 2016-12-16 | 2022-06-07 | General Electric Company | Collimator structure for an imaging system |
US10191162B2 (en) * | 2017-05-05 | 2019-01-29 | Prismatic Sensors Ab | Radiation hard silicon detectors for x-ray imaging |
US10610191B2 (en) * | 2017-07-06 | 2020-04-07 | Prismatic Sensors Ab | Managing geometric misalignment in x-ray imaging systems |
CN111242913B (zh) * | 2020-01-08 | 2023-04-11 | 浙江大学 | 一种获取肋骨展开图像的方法、系统、装置及存储介质 |
CN111134710B (zh) * | 2020-01-17 | 2021-05-07 | 清华大学 | 一种多能量ct成像系统 |
CN116807502B (zh) * | 2023-06-28 | 2024-03-15 | 赛诺威盛医疗科技(扬州)有限公司 | 用于校正计算机断层扫描设备的扫描数据的方法及装置 |
Citations (5)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
JP2004325183A (ja) * | 2003-04-23 | 2004-11-18 | M & C:Kk | 放射線検出方法、放射線検出器、及び、この検出器を搭載した放射線撮像システム |
JP2005349080A (ja) * | 2004-06-14 | 2005-12-22 | Ge Medical Systems Global Technology Co Llc | 放射線断層撮影装置およびその断層撮影方法、補正データ算出方法 |
JP2006503631A (ja) * | 2002-10-25 | 2006-02-02 | コーニンクレッカ フィリップス エレクトロニクス エヌ ヴィ | コンピュータ断層撮影用のダイナミック検出器インターレーシング |
JP2006314425A (ja) * | 2005-05-11 | 2006-11-24 | Ge Medical Systems Global Technology Co Llc | X線ct装置 |
JP2007529258A (ja) * | 2004-03-17 | 2007-10-25 | コーニンクレッカ フィリップス エレクトロニクス エヌ ヴィ | 複数の焦点取得方法及び装置 |
Family Cites Families (27)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
US4504962A (en) * | 1978-12-22 | 1985-03-12 | Emi Limited | Computerized tomography |
JPS5764047A (en) * | 1980-10-08 | 1982-04-17 | Tokyo Shibaura Electric Co | Radiation tomographing diagnostic device |
US4897788A (en) * | 1988-04-18 | 1990-01-30 | General Electric Company | Image correction for computed tomography to remove crosstalk artifacts |
JP3042810B2 (ja) * | 1992-12-04 | 2000-05-22 | 株式会社東芝 | X線コンピュータトモグラフィ装置 |
JPH08606A (ja) * | 1994-06-23 | 1996-01-09 | Hitachi Medical Corp | X線ct装置 |
US5510622A (en) * | 1994-07-21 | 1996-04-23 | General Electric Company | X-ray detector array with reduced effective pitch |
IL119033A0 (en) * | 1996-08-07 | 1996-11-14 | Elscint Ltd | Multi-slice detector array |
JP3828967B2 (ja) * | 1996-10-30 | 2006-10-04 | 株式会社東芝 | X線ctスキャナ |
US5960056A (en) * | 1997-07-01 | 1999-09-28 | Analogic Corporation | Method and apparatus for reconstructing volumetric images in a helical scanning computed tomography system with multiple rows of detectors |
US6437338B1 (en) * | 1999-09-29 | 2002-08-20 | General Electric Company | Method and apparatus for scanning a detector array in an x-ray imaging system |
DE19956585A1 (de) * | 1999-11-25 | 2001-05-31 | Philips Corp Intellectual Pty | Computertomographie-Verfahren |
JP2001187046A (ja) * | 1999-12-27 | 2001-07-10 | Ge Medical Systems Global Technology Co Llc | マルチスライスx線ct装置及びその制御方法 |
JP3987676B2 (ja) | 2000-07-10 | 2007-10-10 | 株式会社日立メディコ | X線計測装置 |
EP1389956B1 (en) * | 2001-04-25 | 2012-10-31 | Amnis Corporation | Method and apparatus for correcting crosstalk and spatial resolution for multichannel imaging |
US6535572B2 (en) * | 2001-06-15 | 2003-03-18 | Ge Medical Systems Global Technology Company, Llc | Methods and apparatus for compensating computed tomographic channel ganging artifacts |
WO2004064385A1 (en) * | 2003-01-16 | 2004-07-29 | Philips Intellectual Property & Standards Gmbh | Array of sensor elements |
JP4041040B2 (ja) * | 2003-09-08 | 2008-01-30 | ジーイー・メディカル・システムズ・グローバル・テクノロジー・カンパニー・エルエルシー | 放射線断層撮影装置 |
US7539284B2 (en) * | 2005-02-11 | 2009-05-26 | Besson Guy M | Method and system for dynamic low dose X-ray imaging |
JP4881071B2 (ja) * | 2006-05-30 | 2012-02-22 | 株式会社日立製作所 | 放射線検出器、及びこれを搭載した放射線撮像装置 |
JPWO2008010512A1 (ja) * | 2006-07-19 | 2009-12-17 | 株式会社日立メディコ | X線ct装置及び画像ノイズ低減方法 |
US7916836B2 (en) * | 2007-09-26 | 2011-03-29 | General Electric Company | Method and apparatus for flexibly binning energy discriminating data |
US8437526B2 (en) * | 2007-12-05 | 2013-05-07 | Siemens Medical Solutions Usa, Inc. | System for adaptively processing medical image data |
JP5675808B2 (ja) | 2009-08-13 | 2015-02-25 | コーニンクレッカ フィリップス エヌ ヴェ | 独立したx及びz方向の動的フォーカルスポット偏向を持つX線管 |
JP5579505B2 (ja) * | 2010-06-03 | 2014-08-27 | 株式会社日立メディコ | X線ct装置 |
US20120166128A1 (en) * | 2010-12-28 | 2012-06-28 | Ikhlef Abdelaziz | Method and apparatus for detector calibration |
DE102011082878A1 (de) * | 2011-09-16 | 2013-03-21 | Siemens Aktiengesellschaft | Röntgendetektor einer gitterbasierten Phasenkontrast-Röntgenvorrichtung und Verfahren zum Betreiben einer gitterbasierten Phasenkontrast-Röntgenvorrichtung |
DE102012224258A1 (de) * | 2012-12-21 | 2014-06-26 | Siemens Aktiengesellschaft | Röntgenaufnahmesystem zur differentiellen Phasenkontrast-Bildgebung eines Untersuchungsobjekts mit Phase-Stepping sowie angiographisches Untersuchungsverfahren |
-
2013
- 2013-06-06 US US14/409,100 patent/US9818182B2/en active Active
- 2013-06-06 WO PCT/JP2013/065639 patent/WO2013191001A1/ja active Application Filing
- 2013-06-06 JP JP2014521285A patent/JP5963217B2/ja active Active
Patent Citations (5)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
JP2006503631A (ja) * | 2002-10-25 | 2006-02-02 | コーニンクレッカ フィリップス エレクトロニクス エヌ ヴィ | コンピュータ断層撮影用のダイナミック検出器インターレーシング |
JP2004325183A (ja) * | 2003-04-23 | 2004-11-18 | M & C:Kk | 放射線検出方法、放射線検出器、及び、この検出器を搭載した放射線撮像システム |
JP2007529258A (ja) * | 2004-03-17 | 2007-10-25 | コーニンクレッカ フィリップス エレクトロニクス エヌ ヴィ | 複数の焦点取得方法及び装置 |
JP2005349080A (ja) * | 2004-06-14 | 2005-12-22 | Ge Medical Systems Global Technology Co Llc | 放射線断層撮影装置およびその断層撮影方法、補正データ算出方法 |
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US9818182B2 (en) | 2017-11-14 |
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