JP2002541896A - Apparatus and method for use in a computed tomography system utilizing a reduced size detector that covers only the half-field - Google Patents

Abstract

PROBLEM TO BE SOLVED: To acquire projection data of a subject by a computed tomography (CT) system including a detector shifted by half of the width with respect to an isocenter. SOLUTION: For each projection view, a value Va of one detector element is selected from the detector elements closest to the isocenter of the CT system, and for the selected detector element, from the opposite direction or the same. From the forward projection of the direction, the value Vb of the detector element is estimated. Next, a smoothing function that can remove the difference between Va and Vb is selected. The smoothing function is then applied to remove the difference between Va and Vb. Next, a smooth transition area is generated by applying a weighting function to smooth the steps when the true projection data and the estimated projection data are combined.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

TECHNICAL FIELD OF THE INVENTION

The present invention covers only half of the field of view, thereby increasing (or substantially without increasing) artifacts in an area detector.
(volumetric computed tomography) using a reduced-size area detector that can reduce the size and cost of (area detector)
The present invention relates to a method and apparatus used in a mography (VCT) system.

[0002]

BACKGROUND OF THE INVENTION

In computed tomography (CT), generally, x-raying a patient, collecting digital x-ray projection data of a portion of the patient's body, processing and backprojecting the digital x-ray projection data, and then a CT system It is necessary to display an image on a display monitor. CT systems typically include a gantry, table, X-ray tube, X-ray detector array, computer, and display monitor. The computer sends a command to the gantry control device, and the gantry sends X
The tube and / or the detector array are rotated at a certain rotational speed.

In the third generation of CT systems, there is a relative rotational movement between the gantry and the patient's body, part of which is constituted by a detector array and an X-ray tube. As this relative rotational movement occurs, the computer controls the data acquisition process performed by the x-ray tube and detector array to acquire a digital radiograph. The computer then processes and backprojects the digital radiographic data by executing a reconstruction algorithm, and displays the reconstructed CT image on a display monitor.

[0004] Many CT systems currently in use utilize a single row of detectors in the gantry, and this single row detector is commonly referred to as a linear array of detector elements. A further improved CT system uses a linear array of two to four rows of detectors to create a multi-row detector. Both detector arrangements can be used in a helical scan protocol. However, with a multi-row detector, the patient can be scanned more easily by increasing the helical pitch of the detector array so that the specified axial coverage of the patient can be scanned more quickly. The helical pitch is typically the gantry 1 of the table supporting the patient.
It is defined as the ratio of displacement during rotation to detector pitch. For example, a helical pitch of 1 means that during one revolution of the CT gantry of the CT system, the patient table translates by an amount equal to the detector pitch.

[0005] Typically, the X-ray source emits X-
The entire imaging area for the line fan beam is covered. In other words, X-rays that have passed through the subject being scanned (which may or may not be a patient) or illuminated that region of the subject are absorbed by the detector array.

[0006] In some CT imaging systems, it is desirable to reduce the size of the detector array, and in some cases, the size of the detector array. For example, recent developments in CT technology have used area detector arrays composed of multiple rows of linear detector arrays for CT data collection. At present, detector panels that cover the entire field of view, ie the entire area of the patient being imaged, are not yet available. In addition, some systems that use linear detector arrays provide an extremely large field of view compared to the patient being scanned. Even in this situation, it is desirable to reduce the size and cost of the detector array.

One technique that has been used to overcome these limitations is to translate a detector array that is smaller in size and half its width. For example,
Assume that the original size of the detector array required to cover the desired imaging area of the patient is 80 cm. A width equal to half of the original detector width (40 in this case)
cm) can be used. The detector covers approximately half the field of view of the CT imaging system,
It is displaced by half the detector width (20 cm in this example). According to this example, the same field of view of the patient can be acquired by a detector having a width equal to half the original value.

[0008] Furthermore, it is possible to enlarge the imaging area with a system having a fixed width detector. Typically, the projection of the center of rotation of the CT imaging system coincides with the center of the detector panel. The rotation center of CT imaging is X around it.
The physical location of the point at which the source and detector array are rotating. But,
By displacing the detector by half the detector width with respect to the original position, the field of view (FOV) of the system can be expanded. The projection of the center of rotation of the imaging system is near the edges of the shifted linear or multi-row detector. However, the detector still measures projection data from X-rays passing through the physical center of rotation of the imaging system (ie, the isocenter location). On the other hand, this arrangement effectively doubles the imaging area of the imaging system from its original configuration, thereby greatly increasing the imaging area of the imaging system. A system configuration that shifts a detector by half its width is typically referred to as a semi-detector shift.

A fan beam CT system, which is a CT system in which the X-ray source is a point and irradiates only the detector panel and emits X-rays having a shape similar to a fan at a certain angle aperture, is a CT gantry. Need to collect projection data for a fraction of the full rotation of. Specifically, projection data needs to be collected while the gantry rotates around the patient by an angle region equal to 180 degrees plus the fan angle. Again, the fan angle is a measure of the angular aperture of the x-rays such that only the detector array is illuminated in the axial plane of the imaging system. Obviously, some of the projection data is redundant because the measurement of the projection does not necessarily have to be made during the full 360 degree rotation of the gantry around the patient.

In a semi-detector shift configuration of a CT system, data is collected for a full 360 degree rotation of the gantry. At each view angle of the gantry, only half of the projection data is measured. Data from other views of the gantry is used to complete projection data for a given view angle. The manner in which this process is performed is well known in the art. However, if the measured projection data covering half the field of view of the imaging system is combined with data from other views of the gantry, the resulting projection data may not match well near the center of the projection data. is there. These mismatches, if not reduced or eliminated, can cause undesirable artifacts in the reconstructed image.

One technique currently used to mitigate artifacts caused by discontinuities in projection data within the field of view is to use a weighting function to smooth the discontinuities in the transition region. ing. This technique requires that the detector have some additional detector elements that extend beyond the projection of the center of rotation of the imaging system onto the detector. As the gantry rotates 360 degrees around the patient, the area of the shifted detector panel that extends slightly beyond the projection of the center of rotation on the detector in both directions is called the transition area. The actual data is measured by the detector in half of the transition area. Also within the second half of the transition area
Data can be created from another view of the gantry. The data in the transition region is multiplied by a weighting factor to smooth discontinuities. In general, the image quality is improved when the transition area is large. However, the imaging area of this system configuration is slightly smaller than the area provided by the half-detector shift configuration, which leads to an increase in system cost.

There is a need to improve the integration of the measured data with the data created in the transition area so that the entire imaging area can be realized with this detector array.

In a stereoscopic CT system utilizing a half-detector shift configuration, the projection data is measured within half of the imaging area of the imaging system, while the other half of the projection radiograph is taken from an opposite ray. Need to be created. However, the projection data measured at another projection angle of the CT gantry is twice the original width and the deviation (
(Offset) does not have the same azimuth as the ray when measured by a detector having no offset. Therefore, there is a need for a VCT system that utilizes an area detector in a semi-detector shift configuration, realizes its benefits, and overcomes the above problems.

[0014]

Summary of the Invention

1 is a computed tomography (CT) system including an X-ray source and a detector for acquiring projection data of a subject. The detector has been shifted by half its width relative to a center position corresponding to the projection of the center of rotation of the CT system onto the detector.
According to the method of the invention, for each projection view, a value Va of one detector element is selected from the detector elements closest to the isocenter of the CT system. Then
For the selected detector element, the value Vb of the detector element is estimated from the opposite direction or from the forward projection in the same direction. Next, a smoothing function that can remove the difference between Va and Vb is selected. The smoothing function is then applied to remove the difference between Va and Vb. Next, a weight function is applied to smooth the transition when combining the true projection data and the estimated projection data, thereby generating a smooth transition area.

[0015]

BEST MODE FOR CARRYING OUT THE INVENTION

Prior to describing the method and apparatus of the present invention, referring to FIG.
Here is a general discussion of the T system. FIG. 1 is a block diagram of a stereoscopic CT scanning system suitable for implementing the method and apparatus of the present invention. This stereoscopic CT scanning system will be discussed for use in reconstructing images for patient anatomical features, but the invention is limited to imaging any particular object Please understand that it is not. Further, as will be appreciated by those skilled in the art, the present invention may be used for industrial processes. Furthermore, the invention is not limited to medical CT devices, but rather an industrial system in which the geometry of the X-ray source and detector is fixed and the subject rotates during the scan time. Is also included.

In a stereoscopic CT scanning system, the gantry rotates around a subject, such as a patient, and projection data is collected. The computer 1 controls the operation of the stereoscopic CT scanning system. In this specification, the term gantry rotation is intended to mean the rotation of the X-ray tube 2 and / or the rotation of the detector 3 (preferably a high-resolution area detector). The X-ray tube 2 and the area type detector 3 are included in a gantry. The control devices 4A and 4B are three-dimensional CT
It is controlled by a computer 1 of the scanning system and is coupled to an X-ray tube 2 and a detector 3, respectively. The X-ray tubes 2 and / or the control devices 4A and 4B
Alternatively, the detector 3 is given an appropriate relative rotational movement. The control devices are not necessarily individually required. The gantry can also be rotated using a single controller component. It should also be noted that the computer 1 controls image scan time variation, image resolution and / or axial coverage to implement the method of the present invention.

The computer 1 controls the data acquisition process by instructing the data acquisition system 6 when to sample the detector 3 and controlling the speed of the gantry. In addition, the computer 1 instructs the data acquisition system 6 to set the resolution of the radiograph obtained by the area type detector 3, thereby making it possible to change the resolution of the system. The data collection system 6 includes readout electronics as shown.

The area type detector 3 is configured by an array (not shown) including detector elements. Each detector element measures an intensity value corresponding to the respective element, which is related to the amount of X-ray energy incident on the detector element. By incorporating the apparatus and method of the present invention into a stereoscopic CT scanning system, a new stereoscopic CT scanning system is created. Therefore, the present invention can provide a novel stereoscopic CT scanning system.

It should be noted that the present invention is not limited to any particular computer in order to perform data collection and perform processing according to the present invention. As used herein, the term "computer" is intended to refer to any device capable of performing the calculations and computations necessary to perform the processes according to the present invention. Intention to do so. Thus, the computer utilized to execute the control algorithm 10 according to the present invention can be any device capable of performing the required processing.

With respect to the present invention, it has been found that according to alternative data smoothing schemes, it is not necessary to use additional detector elements to cover the transition region. further,
One alternative uses an iterative algorithm to estimate the projection data that would have been measured if the data had been collected using a large detector array that covered the entire field of view. . Since the errors in the transition region are handled in a similar manner, the two schemes will be discussed below within the same context.

In this technique, one set of data is obtained by either forward-projecting the reconstructed data obtained from the previous iteration step or by interpolating redundant detector data obtained from a set of oppositely directed rays. X-ray projection {P _a } is formed. Here, another set of projection data in the direction opposite to that of {P _a } is formed for rays in the opposite direction. The technique of forward projection is the process of emitting rays from virtual x-ray sources, which traverse the reconstructed volume towards individual detector elements. The linear attenuation values reconstructed along the ray are summed along the ray, and this is referred to as the linear integral of the linear attenuation coefficient.

The technique of forward projection of reconstructed data (referred to as FPT) is generally suitable for generating projection data corresponding to a larger cone angle (ie, as used in VCT systems), while Technique for interpolating redundant projection data (referred to as PDT)
Is more suitable for projection data at a position closer to the intermediate plane (that is, a position closer to the isocenter). Like the fan angle, the cone angle refers to the angular range of X-rays emitted from the X-ray source in a direction perpendicular to the fan angle direction. Distortion in the image near the isocenter due to the difference between the detector estimate obtained using either FPT or PDT and the original value that would have been obtained if the data were actually measured. May occur.

In order to reduce such distortion, a technique using a smoothing function has been developed. The smoothing function can be described as follows with reference to FIG.

1. For each projection view, select the known detector element that is closest to the isocenter (21). Hereinafter, this is represented by Va. 2. For the same detector element, an estimate is obtained (ie, via interpolated projection data (PDT) from an alternate view or by forward projection (FPT) of the reconstructed data) (22). Hereinafter, this is represented by Vb. 3. Generate an appropriate smoothing function (23). 4. The difference between Va and Vb is reduced by a smoothing function, and this is stepwise smoothed in a region near the center of the imaging area of the imaging system (24).

This can be understood from the following considerations. That is, the amount of discontinuity of the projection data at the center of the shooting area is set to d = Va−Vb. As an example, one of the smoothing functions that can be used to smooth this difference stepwise can be an exponential function defined by:

V = 0.5 ^de -ax ₀ (Equation 1) In the above equation, x ₀ is the absolute value of the distance from the detector position corresponding to the detector element value Va, and a is related to this smoothing function. This is a coefficient for controlling the slope of the curve. This exponential function is added (subtracted) to the projection value located on one side of the central ray position (the detector position corresponding to the projection of the center of rotation of the imaging system onto the detector) and is lower ( The exponential function is added (subtracted) to the projection value on the other side of the central ray position to decrease (increase) the higher (lower) original value, while increasing (decreasing) the estimated value. I have. In other words, by combining the true projection data and the estimated projection data, it is possible to provide a method of reducing the mismatch of the projection data at the position of the center ray. Therefore, this process can provide a smooth data transition region with reduced or eliminated artifacts (2
5).

[0027] At present, the prior art does not know how to use an area detector in a semi-detector shift configuration in a stereoscopic CT (VCT) system. VCT system,
Having described above various known techniques for eliminating projection data discontinuities in the field of view of an imaging system by creating a transition region, another aspect of the present invention will now be described. To

The variable fθ And fθ ′ To represent two functions obtained at the source angle θ and representing the signal intensities of the forward ray and the opposite ray (the same angle and traversal but opposite directions), respectively. To

Fθ (n) = 0 (when N ₂ <n <N) (Equation 2) fθ ′ (n) = 0 (when 1 <n <N ₂ ) (Equation 3) Ideally, fθ (N ₂ )
Is fθ '(N ₂ ). However, this cannot happen for the following reasons.

(A) The actual shape of each ray is a shallow tetrahedron starting from the source and ending at the detector element, except that the subject is entirely homogeneous and rotationally symmetric. No two rays cross exactly the same location in (B) The subject / patient movement during the scan cycle introduces additional errors in each forward / backward ray pair. (C) no efficient complete interpolation scheme has been developed so far; That is, errors are introduced by the interpolation process.

Fθ at source angle position θ (N ₂ ) and fθ Assuming that the difference of '(N ₂ ) is d (θ), if d (θ) is completely random, the error introduced into the reconstructed image is probably related to other random errors in CT. It must be covered by quantum noise. However, if the error is somewhat systematic,
Obvious artifacts will be introduced in the reconstructed image. For this reason, it is necessary to use a smoothing process in the transition region. In other words, this smoothing function is fθ And fθ ′ Have been developed to be smaller around the central ray position represented by detector element N ₂ .

Fθ And fθ ′ Will be represented using W and W ′. In deriving W and W ', certain criteria must be considered, as those skilled in the art will understand. Further, those skilled in the art will appreciate that a variety of smoothing functions other than those specifically listed herein are suitable for this purpose, such as the following examples. .

W (n) + W ′ (n) = 1 (for all n) (Equation 4) δW / δn = δW ′ / δn = 0 (at the position of n = N ₂ ± Δn) (Equation 5) )

In the above equation, Δn sets the degree of smoothness of W and W ′, and δ
Arithmetic. In addition, traditional smoothing functions for this type of application are generally
Note that it is also known as a feathering function.
We use W and W 'to find the transition region between the forward and the opposite ray. flat
For the smoothing function to work properly, Δn must be an integer greater than zero.
Note that there is no. In fact, the larger Δn, the better the smoothing function
Works well. However, if Δn is too large, the detector element at the center ray position will not be used.
N_TwoIt is necessary to use additional detector elements to extend beyond. did
Thus, in the selection of Δn, Δn is sufficiently large, but additional detection
It must not be so large that additional device elements are required. Is this something
Fθ And fθ 'Has the following limit:

Fθ (n) = 0 (when N ₂ + Δn <n <N) (Equation 6)
) Fθ '(N) = 0 (when 1 <n <N ₂ −Δn) (Equation 7)
)

The actual detector signal used in the backprojection process is Wfθ And W'fθ '. Since a wider transition region tends to remove much of the mismatch error between the forward and opposite rays, it is necessary to construct opposite rays for each half field of view (FOV) projection data. Note further that there is no. In other words, zero is embedded in each half-FOV data (and each value of the additional Δn detector) to obtain detector data of length N, followed by performing a conventional filtered projection procedure. No interpolation is needed.

Increasing the number of detector rows in a CT system, ie, in an area detector, motivates the additional detector elements to be more than (Δn times the number of rows), resulting in Δn
There is a need to improve upon conventional approaches by devising methods and apparatus that can minimize.

In the method according to the present invention, Δn is reduced to 1. On the other hand, computer simulations show that Δn needs to be approximately 20 in the conventional smoothing method in order to achieve an equivalent artifact level. This advantage becomes even more important in VCT applications where the number of detector elements required in the transition area of the area detector is three orders of magnitude higher than the number of elements required in a linear array.

There may be systematic errors in the forward ray as well as in the opposite ray obtained by interpolation or forward projection. fθ (N ₂ ) and fθ The stepwise error can be measured by detecting the difference between '(N ₂ ). here,
d (θ) = fθ (N ₂ ) −fθ '(N ₂ ), d (θ) becomes fθ (N ₂
) And fθ '(N ₂ ) as a step error (where θ is the angular position of the X-ray source).

In this method, the step error is smoothed using the exponential function described in (Equation 1) (where a is a control coefficient for controlling the smoothing of the exponential function). A function of the forward and opposite rays, fθ (N) and fθ '(N) are two separate functions, gθ, (N) and gθ '(N).

Gθ (n) = fθ (n) −pθ (N ₂ −n) (Equation 8) gθ ′ (n) = fθ ′ (n) + pθ (n−N ₂ ) (Equation 9) (Equation 8) And (Equation 9) gives gθ (N ₂ ) = gθ '(N ₂ ).

The following procedure is executed for each projection image. 1. By embedding zero according to (Equation 2), fθ The original half FO called (n)
Obtain V projection data. 2. Set of rays in the opposite direction fθ '(N) is obtained and zero is embedded according to (Equation 3). 3. According to (Equation 8) and (Equation 9), a smoothing function is applied based on the step error d (θ) (where d (θ) = fθ (N ₂ ) −fθ). '(N _2)). 4. gθ (N) and gθ '(N) and a set of N detector data, ie, hθ (N) is formed. Here, when N ₂ <n <N, hθ (n) = gθ (n), and when 1 <n <N ₂ , hθ (n) = g
θ '(N). 5. The conventional filtered backprojection is applied to hθ (n). 6. Steps 1 to 5 are repeated for all projection angles.

The above procedure is valid for any CT scanner that can acquire rays in the opposite direction by interpolating other projection data from various angles. In the case of a 2D fan beam, where the data of another half FOV can always be approximately calculated from the redundant fan beam projection data, this is a perfect case. If this technique is extended to 3D-VCT, complete interpolation can only occur on the mid-plane when using circular orbits.

According to our simulations, if we apply the same technique to VCT using a circular orbit, this technique will work for cone angles in the range of ± 1.5 degrees (for the position of the central ray). It has been shown to perform better than conventional smoothing techniques (using more than 20 additional detectors). However, this is not the case for large cone angles. This is because the angular orientation of the ray in the opposite direction is considerably different from the measured data when a perfect detector is used. To address this situation, iterative techniques can be used to improve image quality. The procedure is shown below.

1. An initial three-dimensional image is obtained according to the above steps 1 to 6. 2. In the second iteration of the method, “forward rays” for each set of half-FOV projection data are acquired using forward projection, and these are obtained according to steps 3-6 outlined above. Combine into a complete set of projection data. 3. Step 2 is continued until the process converges, that is, until the quality of the image has not improved.

It should be noted that the present invention has been discussed with respect to particular embodiments. Therefore, the present invention is not limited to these embodiments. For example, consider 3
By way of example, this does not mean that it covers all the methods that can obtain the proper operation mode of the VCT system using the trade-off relationship of the above parameters. These examples are discussed to illustrate the concepts of the present invention, as well as the manner in which these fundamental parameters are in a trade-off relationship to achieve a proper scanning protocol. Furthermore, this trade-off is not limited to one scan protocol,
That is, the trade-off relationship can be applied to both the axial scan protocol (without moving the patient table during the scan) and the helical scan protocol. One of ordinary skill in the art would understand how to utilize and extend these concepts to enable other area detector scan protocols useful for particular applications.

[Brief description of the drawings]

FIG. 1 is a block diagram of a CT system according to the present invention.

FIG. 2 is a diagram illustrating the excursion of a detector utilized in accordance with the method of the present invention.

FIG. 3 is a block diagram illustrating the method of the present invention according to a preferred embodiment.

[Procedure amendment]

[Submission date] September 27, 2001 (2001.9.27)

[Procedure amendment 1]

[Document name to be amended] Statement

[Correction target item name] Full text

[Correction method] Change

[Contents of correction]

[Title of Invention] devices and methods used in computer tomography system utilizing a reduced size detector that covers only half shooting range

[Claims]

Wherein said detector includes a plurality of detector elements, for each of the plurality of projection view, a detector element value Va related to one detector element is selected by the computer, the detection vessel element value Va is a value representing the line integral of the line attenuation coefficient of the rays connecting the test can elements associated with the source through the subject the detector associated with the detector element value computer was or Estimating another detector element value Vb for the element , and in each projection view, the detector element value Va corresponds to the measured projection data, and the detector element value Vb corresponds to the estimated projection data; 2. The method of claim 1, further comprising applying a smoothing function to remove differences between the detector element values Va and Vb after the detector element values Va and Vb are obtained for each projection view. The three-dimensional computer described Data tomography system.

3. A three-dimensional computed tomography system according to claim 2, wherein said detector is a high-resolution area detector .

4. A method for weighting a smoothing function prior to applying the smoothing function and applying the weighted smoothing function to the measured projection data and the estimated projection data to obtain a detector at a central ray position of the detector. 4. A stereoscopic computer according to claim 3, wherein the difference between the element values Va and Vb is reduced and the difference between the detector element values Va and Vb is smoothed in a region near the center of the field of view of the VCT system. Tomography system.

5. A smoothing function, the detector three-dimensional computed tomography system according to claim 4, wherein the weighted by a half of the difference between the element values Va and Vb are added to the measured projection data and estimated projection data .

6. A smoothing function, the detector element values Va and steric computer tomography is the請 Motomeko 4, wherein the weighting by the half the difference is subtracted from the measured projection data and estimated projection data between Vb system.

7. A method for acquiring projection data of a subject using a computed tomography (CT) system, the method comprising: projecting X-rays from the X-ray source to the subject; Receiving X-rays projected in a projection view at a detector of a CT system, said detector comprising a plurality of detector elements;
Generating an electrical signal in response to X-rays incident on the detector; digitizing the electrical signal; and for each projection view, determining the location of the central ray on the detector of the CT system. a step of selecting a detector element values Va close inspection can element to the detector element associated with the detector element value Va obtained by the selection, by interpolation of the projection data collected from the opposite direction, or in the same direction Estimating a detector element value Vb for said detector element by forward projection of the reconstructed data of said detector element; and selecting a smoothing function capable of removing a difference between Va and Vb . weight in a projection view of each detector element value Va is corresponds to the measured projection data, and the detector element value Vb corresponds to the estimated projection data; and smoothing function And using the weighted smoothing function to reduce the difference between the combined measured and estimated projection data to generate a smooth transition region .

8. A detector element value Va obtained by the selection is a value representing the line integral of the line attenuation coefficient of the rays connecting the detector element associated with the source through the subject the respect projection view of each 8. The method of claim 7 , wherein after obtaining the detector element values Va and Vb, utilizing the weighted smoothing function is performed.

9. The method of claim 8, wherein the smoothing function is weighted by adding half the difference between the detector element values Va and Vb to the measured projection data and the estimated projection data .

10. The method of claim 8, wherein the smoothing function is weighted by subtracting half the difference between the detector element values Va and Vb from the measured projection data and the estimated projection data .

11. A X-ray source for projecting X-rays to a subject, see contains a plurality of detector elements, receives the X-rays projected from the X-ray source, an electrical signal in response to incident X-rays a detector for generating, by reading the electric signal from the detector, which includes a data collection component that converts the Digi tal signal, three-dimensional computer tomography for obtaining projection data of the subject ( VCT) system , wherein the detector is shifted by half the width of the detector with respect to a center position corresponding to the projection of the center of rotation of the VCT system onto the detector. , in the device receives the digital signal from the data acquisition component, a computer capable of executing the reconstruction algorithm, the reconstruction algorithm by running Yes to that device a computer to reconstruct the image when processing the digital signal.

12. For each of a plurality of projection views, the computer selects a detector element value Va associated with a detector element, the detector element value Va passing through the subject and the source and the source element. a value representing the line integral of the line attenuation coefficient of the rays connecting the associated detector element, the computer is also have One detector element associated with the detector element value to estimate the further detector element value Vb, In each projection view, the detector element value Va corresponds to the measured projection data, and the detector element value Vb corresponds to the estimated projection data, and the computer also determines, for each projection view, the detector element value The apparatus of claim 11 , wherein after the values Va and Vb are obtained, a smoothing function is applied to remove differences between the detector element values Va and Vb .

Wherein said detector 12. Symbol mounting device is the area detector of high resolution.

14. The computer, prior to applying a smoothing function, weights the smoothing function and applies the weighted smoothing function to the measured projection data and the estimated projection data to determine the center ray of the detector. reducing the difference between the detector element values Va and Vb at the position, according to claim 13, wherein smoothing the difference between the detector element child values Va and Vb in the region near the center of the shooting range of the VCT system Equipment.

15. smoothing function, the detector device 請 Motomeko 14 wherein weighted by half of the difference between the element values Va and Vb are added to the measured projection data and estimated projection data.

16. The apparatus of claim 14, wherein the smoothing function is weighted by subtracting half the difference between the detector element values Va and Vb from the measured projection data and the estimated projection data .

17. embodied on a computer readable medium, and computer tomography (CT) A computer program for acquiring projection data of the object using the system, detection of CT system A first code segment for selecting the detector element value Va of the detector element closest to the position of the center ray on the detector, and a detector element associated with the selected detector element value Va in opposite directions. A second code segment for estimating the detector element value Vb for the detector element by interpolation of the projection data collected from or from the forward projection of the reconstructed data in the same direction ; a third code segment for selecting a smoothing function capable of removing the difference in projection view of each detector element value Va measurement Corresponding to the projection data, and the detector element value Vb corresponds to the estimated projection data, and a third code segment, the fourth code segment to weight the smoothing function, a smoothing function that the weighting by using the measured projection data and to reduce the difference when the combination of the estimated projection data, the fifth code segment and a comprise computer programs that generates a smooth transition region.

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention covers only half of the field of view, thereby increasing (or substantially without increasing) artifacts in an area detector.
(volumetric computed tomography) using a reduced-size area detector that can reduce the size and cost of (area detector)
The present invention relates to a method and apparatus used in a mography (VCT) system.

[0002]

BACKGROUND OF THE INVENTION In computed tomography (CT), generally X-raying a patient, collecting digital X-ray projection data of a part of the patient's body, and processing and back-projecting the digital X-ray projection data It is then necessary to display the image on the display monitor of the CT system. CT systems typically include a gantry, table, X-ray tube, X-ray detector array, computer, and display monitor. The computer sends a command to the gantry control device, and the gantry sends X
The tube and / or the detector array are rotated at a certain rotational speed.

In the third generation of CT systems, there is a relative rotational movement between the gantry and the patient's body, part of which is constituted by a detector array and an X-ray tube. As this relative rotational movement occurs, the computer controls the data acquisition process performed by the x-ray tube and detector array to acquire a digital radiograph. The computer then processes and backprojects the digital radiographic data by executing a reconstruction algorithm, and displays the reconstructed CT image on a display monitor.

[0004] Many CT systems currently in use utilize a single row of detectors in the gantry, and this single row detector is commonly referred to as a linear array of detector elements. A further improved CT system uses a linear array of two to four rows of detectors to create a multi-row detector. Both detector arrangements can be used in a helical scan protocol. However, the multi-row detector facilitates scanning of the patient by increasing the helical pitch of the CT system so that the specified axial coverage of the patient can be scanned more quickly. The helical pitch is typically the gantry 1 of the table supporting the patient.
It is defined as the ratio of displacement during rotation to detector pitch. For example, a helical pitch of 1 means that during one revolution of the CT gantry of the CT system, the patient table translates by an amount equal to the detector pitch.

[0005] Typically, the X-ray source emits X-
The entire imaging area for the line fan beam is covered. In other words, X-rays that have passed through the subject being scanned (which may or may not be a patient) or illuminated that region of the subject are absorbed by the detector array.

[0006] In some CT imaging systems, it is desirable to reduce the size of the detector array, and in some cases, the size of the detector array. For example, recent developments in CT technology have used area detector arrays composed of multiple rows of linear detector arrays for CT data collection. At present, detector panels that cover the entire field of view, ie the entire area of the patient being imaged, are not yet available. In addition, some systems that use linear detector arrays provide an extremely large field of view compared to the patient being scanned. Even in this situation, it is desirable to reduce the size and cost of the detector array.

One technique that has been used to overcome these limitations is to increase its size.
Translation of the detector array, which has been made smaller and half its width. For example, The original size of the detector array is 80c to cover the desired imaging area of the patient. Assume that it has a width equal to m. A width equal to half the original detector width (in this case,
Smaller detectors, such as having a height of 40 cm) can be used. This
Detector covers approximately half of the field of view of the CT imaging system.
To shift by half the detector width (20 cm in this example). According to this example
If the same field of the patient is acquired by a detector with a width equal to half the original value
can do.

[0008] Furthermore, it is possible to enlarge the imaging area with a system having a fixed width detector. Typically, the projection of the center of rotation of the CT imaging system coincides with the center of the detector panel. The rotation center of CT imaging is X around it.
The physical location of the point at which the source and detector array are rotating. But,
By displacing the detector by half the detector width with respect to the original position, the field of view (FOV) of the system can be expanded. Projection onto the detector of the rotation center of the imaging system, is linear of shifting, in the vicinity of the edge of the multiple lines or area-type detector. However, the detector still measures projection data from X-rays passing through the physical center of rotation of the imaging system (ie, the isocenter location). On the other hand, this arrangement effectively doubles the imaging area of the imaging system from its original configuration, thereby greatly increasing the imaging area of the imaging system. A system configuration that shifts a detector by half its width is typically referred to as a semi-detector shift.

A fan beam CT system, which is a CT system in which the X-ray source is a point and irradiates only the detector panel and emits X-rays having a shape similar to a fan at a certain angle aperture, is a CT gantry. Need to collect projection data for a fraction of the full rotation of. Specifically, while the gantry is rotating in a circular orbit around the patient over equal angular intervals to a value obtained by adding the fan angle to 180 degrees, it is necessary to collect projection data. Again, the fan angle is a measure of the angular aperture of the x-rays such that only the detector array is illuminated in the axial plane of the imaging system. Obviously, some of the projection data is redundant because the measurement of the projection does not necessarily have to be made during the full 360 degree rotation of the gantry around the patient.

In a semi-detector shift configuration of a CT system, data is collected for a full 360 degree rotation of the gantry. At each view angle of the gantry, only half of the projection data is measured. Data from other views of the gantry is used to complete projection data for a given view angle. Method for carrying out this process for linear detector array are well known in the art. But,
If the measured projection data covering half of the field of view of the imaging system is combined with data provided by other views of the gantry, the resulting projection data may not match well near the center of the projection data. These mismatches, if not reduced or eliminated, can cause undesirable artifacts in the reconstructed image.

One technique currently used to mitigate artifacts caused by discontinuities in projection data within the field of view is to use a weighting function to smooth the discontinuities in the transition region. ing. This technique requires that the detector have some additional detector elements that extend beyond the projection of the center of rotation of the imaging system onto the detector. When the gantry rotates 360 degrees around the patient, the area of the shifted detector panel that extends in both directions slightly beyond the projection of the center of rotation on the detector is called the "transition area" . . The actual data is measured by the detector in half of the transition area. Also, within the second half of the transition area, data can be created from another view of the gantry. The measured and combined data in the transition region is multiplied by a weight factor to smooth discontinuities between the data . In general, a large transition area improves the image quality, but the imaging area of this system configuration is slightly smaller than the imaging area provided by the half-detector shift configuration, which leads to an increase in system cost. In this alternative, it is necessary to have some additional detector elements extends beyond the projection of the center of rotation of the Imaging System onto the detector, which further increased the cost of the system.

The integration of the measured data with the data generated in the transition area so that the entire imaging area can be realized with this detector array or, alternatively, the number of individual detector elements can be reduced. Need to be improved.

[0013] In multi-line or three-dimensional CT system utilizing a half-detector shift configuration, while measuring the projection data in half the imaging area of the imaging system, rays in the opposite direction the remaining half of the projection data ( that is, there is a need to create or synthesized from ray (ray)) obtained from the angular position of the gantry to provide a similar line integral data of linear attenuation coefficients of the object being photographed. However, the projection data measured at another projection angle of the CT gantry is twice the original width and has the same azimuth as the ray when measured with a detector having no offset. Absent. This, X-rays are the source or we release has both radial angle (i.e., the axial orientation of the rays in the plane)及 beauty elevation (i.e., the orientation of the rays of the outer axial plane) That's why . Therefore, there is a need for a VCT system that utilizes an area detector in a semi-detector shift configuration, realizes its benefits, and overcomes the above problems.

[0014]

SUMMARY OF THE INVENTION The present invention provides a computed tomography (CT ) system for acquiring projection data of a subject . The CT system includes an X-ray source and a detector. The detector has been shifted by half its width relative to a center position corresponding to the projection of the center of rotation of the CT system onto the detector. In accordance with the method of the present invention, for each projection view , a processed detector element value representing a line integral of a line attenuation value of a ray connecting the source and the detector element through the subject , wherein as the position, the processed detector element value Va which indicates the signal from the nearest detector element to the projection position of the rotation center of the detector on to the image in g systems (CT system) is selected You. The detector element value Vb is then estimated for the selected detector element from the opposite direction or from the same direction of the forward projection. Next, a smoothing function that can remove the difference between Va and Vb is selected. The difference between Va and Vb is then removed by applying this smoothing function .

[0015]

DETAILED DESCRIPTION OF THE INVENTION Prior to describing the method and apparatus of the present invention, the VC of the present invention will be described with reference to FIG.
Here is a general discussion of the T system. FIG. 1 is a block diagram of a stereoscopic CT scanning system suitable for implementing the method and apparatus of the present invention. This stereoscopic CT scanning system will be discussed for use in reconstructing images for patient anatomical features, but the invention is limited to imaging any particular object Please understand that it is not. Further, as will be appreciated by those skilled in the art, the present invention may be used for industrial processes. Furthermore, the invention is not limited to medical CT devices, but rather an industrial system in which the geometry of the X-ray source and detector is fixed and the subject rotates during the scan time. Is also included.

In a stereoscopic CT scanning system, the gantry rotates around a subject, such as a patient, and projection data is collected. The computer 1 controls the operation of the stereoscopic CT scanning system. In this specification, the term gantry rotation is intended to mean the rotation of the X-ray tube 2 and / or the rotation of the detector 3 (preferably a high-resolution area detector). The X-ray tube 2 and the area type detector 3 are included in a gantry. The control devices 4A and 4B are three-dimensional CT
It is controlled by a computer 1 of the scanning system and is coupled to an X-ray tube 2 and a detector 3, respectively. The X-ray tubes 2 and / or the control devices 4A and 4B
Alternatively, the detector 3 is given an appropriate relative rotational movement. The control devices are not necessarily individually required. The gantry can also be rotated using a single controller component. It should also be noted that the computer 1 controls image scan time variation, image resolution and / or axial coverage to implement the method of the present invention.

The computer 1 controls the data acquisition process by instructing the data acquisition system 6 when to sample the detector 3 and controlling the speed of the gantry. In addition, the computer 1 instructs the data acquisition system 6 to set the resolution of the radiograph obtained by the area type detector 3, thereby making it possible to change the resolution of the system. The data collection system 6 includes readout electronics as shown.

The area type detector 3 is configured by an array (not shown) including detector elements. Each detector element measures an intensity value corresponding to the respective element, which is related to the amount of X-ray energy incident on the detector element. By incorporating the apparatus and method of the present invention into a stereoscopic CT scanning system, a new stereoscopic CT scanning system is created. Therefore, the present invention can provide a novel stereoscopic CT scanning system.

It should be noted that the present invention is not limited to any particular computer in order to perform data collection and perform processing according to the present invention. As used herein, the term "computer" is intended to refer to any device capable of performing the calculations and computations necessary to perform the processes according to the present invention. Intention to do so. Thus, the computer utilized to execute the control algorithm 10 according to the present invention can be any device capable of performing the required processing.

[0020] For the present invention, by a method, eliminating the need to use additional detector elements to cover the transition region by alternative data smoothing scheme. Further , one alternative uses an iterative algorithm to estimate the projection data that would be measured if the data were collected using a large detector array that covers the entire field of view. . Iterative algorithm, the step of reconstructing CT data, estimating the error between the data generated by the measured projection data and the forward projection routine, a step of smoothing and / or reducing error Repeating the above processing until the stop requirement is satisfied. The stop requirement can be a threshold on the metric measuring the difference of the reconstructed data between two iterations of the method. One example of such a metric is the well-known least squares metric . The forward projection routine, by the subject or along a ray connecting the location of the X-ray source to the detector for calculating a line integral of the linear attenuation coefficient in the CT reconstruction of a patient, a method for simulating the experimental data collection It is. Since the errors in the transition region are handled in a similar manner, the two schemes will be discussed below within the same context.

In the technique of the present invention , either by forward-projecting the reconstructed data obtained from the preceding iteration step, or by interpolating redundant detector data obtained from a set of oppositely directed rays, Form _a set of X-ray projections {P _a }. Here, another set of projection data in the direction opposite to that of {P _a } is formed for rays in the opposite direction. As mentioned earlier, the technique of forward projection is the process of emitting rays from a virtual x-ray source, which rays traverse the reconstructed volume towards individual detector elements. The linear attenuation values reconstructed along the ray are summed along the ray, and this is referred to as the linear integral of the linear attenuation coefficient. It may utilize techniques known have several to implement the forward projection process.

Iterative techniques for forward projection of reconstructed data (referred to as IFPT ) are generally suitable for producing projection data corresponding to larger cone angles (ie, as used in VCT systems), while The technique of interpolating redundant projection data (referred to as RPDT ) is more suitable for projection data located closer to the mid-plane (ie, a plane with a cone angle of zero degrees) . Like the fan angle, the cone angle refers to the angular range of X-rays emitted from the X-ray source in a direction perpendicular to the fan angle direction. This can be called the X-ray elevation angle. Due to the difference between the detector estimate obtained using either IFPT or RPDT and the original value that would have been obtained if the data were actually measured , the CT system on the detector which may cause distortion in an image near the projection (position) of the center of rotation of the. For simplicity, the position of the projection of the center of rotation of the CT system onto the detector is referred to as the central ray position. RPDT techniques are useful with respect to the center line of the area detector or a plurality of rows detector, also IFPT techniques area detector in cone angle RPDT technique is not valid or to each row in a multiline detector Useful for.

[0023] In order to reduce distortion or artifacts in the reconstructed image, formic method of the present invention utilizes a smoothing function. The smoothing function can be described as follows with reference to FIG.

1. For each projection view, the known detector element closest to the center ray position is selected (21). Hereinafter, this is represented by Va. 2. An estimate is obtained (22) for the same detector element ( via interpolated projection data from alternate views (RPDT) or by forward projection of reconstructed data (IFPT) ). Hereinafter, this is represented by Vb. 3. Generate an appropriate smoothing function (23). 4. The smoothing function is weighted by half the difference between Va and Vb, and the function is added (or subtracted) to the measured data and the composite data when needed . The smoothing function reduces the difference between Va and Vb at the center ray position and smooths this difference stepwise in an area near the center of the imaging area of the imaging system (24).

This can be understood from the following considerations. That is, the amount of discontinuity of the projection data at the center of the shooting area is set to d = Va−Vb. As an example, one of the smoothing functions that can be used to smooth this difference stepwise can be an exponential function defined by:

V = 0.5 ^de -ax ₀ (Equation 1) In the above equation, x ₀ is the absolute value of the distance from the detector position corresponding to the detector element value Va, and a is related to this smoothing function. This is a coefficient for controlling the slope of the curve. The weighted exponential is added (subtracted) lower to the projection value located on one side of the central ray position (the detector position corresponding to the projection of the center of rotation of the imaging system onto the detector). Increase (decrease) the (high) estimate,
This exponential function is added (subtracted) to the projection value on the other side of the center ray position to lower (increase) the higher (lower) original value. In other words, by combining the true projection data and the estimated projection data, it is possible to provide a method of reducing the mismatch of the projection data at the position of the center ray. Thus, this process can provide a smooth data transition region with reduced or eliminated artifacts (25). Using the above-described method, the amount of overlap of the detector at the center ray position in the minimal, it is possible to provide a cost-effective design.

[0027] At present, the prior art does not know how to use an area detector in a semi-detector shift configuration in a stereoscopic CT (VCT) system. VCT system,
Having described above various known techniques for eliminating projection data discontinuities in the field of view of an imaging system by creating a transition region, another aspect of the present invention will now be described. To Accordingly, the above description can be abstracted and created in a general manner.

The variable fθ And fθ ′ To represent two functions obtained at the source angle θ and representing the signal intensities of the forward ray and the opposite ray (the same angle and traversal but opposite directions), respectively. To

Fθ (n) = 0 (when N ₂ <n <N) (Equation 2) fθ ′ (n) = 0 (when 1 <n <N ₂ ) (Equation 3) where the variable n is Used to point to a location within the detector array. For example, if the detector array has 1000 detectors, then 1 <n <1000. N is the number of horizontal detector elements, N ₂ is the position of the center ray in the detector. Since the above two persons is across the same portion of the subject, ideally, f [theta] (N ₂ ) is fθ '(N ₂ ). However, this cannot happen for the following reasons.

(A) The actual shape of each ray is a shallow tetrahedron starting from the source and ending at the detector element, except that the subject is entirely homogeneous and rotationally symmetric. No two rays cross exactly the same location in (B) The subject / patient movement during the scan cycle introduces additional errors in each forward / backward ray pair. (C) no efficient complete interpolation scheme has been developed so far; That is, errors are introduced by the interpolation process.

Fθ at source angle position θ (N ₂ ) and fθ Assuming that the difference of '(N ₂ ) is d (θ), if d (θ) is completely random, the error introduced into the reconstructed image is probably related to other random errors in CT. It must be covered by quantum noise. However, if the error is somewhat systematic,
Obvious artifacts will be introduced in the reconstructed image. For this reason, it is necessary to use a smoothing process in the transition region. In other words, this smoothing function is fθ And fθ ′ Have been developed to be smaller around the central ray position represented by detector element N ₂ .

Fθ And fθ ′ Will be represented using W and W ′. In deriving W and W ', certain criteria must be considered, as those skilled in the art will understand. Further, those skilled in the art will appreciate that a variety of smoothing functions other than those specifically listed herein are suitable for this purpose, such as the following examples. .

W (n) + W ′ (n) = 1 (for all n) (Equation 4) δW / δn = δW ′ / δn = 0 (at the position of n = N ₂ ± Δn) (Equation 5) )

In the above equation, Δn sets the degree of smoothness of W and W ′, and δ is a differential operator. Further, it should be noted that the conventional smoothing function for this type of application is also commonly known as a feathering function.
We use W and W 'to find the transition region between the forward and the opposite ray. Note that Δn must be an integer greater than zero for the smoothing function to work properly. In fact, the greater Δn, the better the smoothing function works. However, making Δn too large requires the use of additional detector elements to extend beyond the central ray detector element N ₂ . Thus, in the selection of Δn, it is necessary that Δn be large enough, but not so large as to require the addition of additional detector elements to the transition region. From these facts, fθ And fθ 'Has the following limit:

Fθ (n) = 0 (when N ₂ + Δn <n <N) (Equation 6)
) Fθ '(N) = 0 (when 1 <n <N ₂ −Δn) (Equation 7)
)

The actual detector signal used in the backprojection process is Wfθ And W'fθ '. Since a wider transition region tends to remove much of the mismatch error between the forward and opposite rays, it is necessary to construct opposite rays for each half field of view (FOV) projection data. Note further that there is no. In other words, zero is embedded in each half-FOV data (and each value of the additional Δn detector) to obtain detector data of length N, followed by performing a conventional filtered projection procedure. No interpolation is needed.

Increasing the number of detector rows in a CT system, ie, in an area detector, motivates the additional detector elements to be more than (Δn times the number of rows), resulting in Δn
There is a need to improve upon conventional approaches by devising methods and apparatus that can minimize.

In the method according to the present invention, Δn is reduced to 1. On the other hand, computer simulations show that Δn needs to be approximately 20 in the conventional smoothing method in order to achieve an equivalent artifact level. This advantage becomes even more important in VCT applications where the number of detector elements required in the transition area of the area detector is three orders of magnitude higher than the number of elements required in a linear array.

There may be systematic errors in the forward ray as well as in the opposite ray obtained by interpolation or forward projection. fθ (N ₂ ) and fθ The stepwise error can be measured by detecting the difference between '(N ₂ ). here,
d (θ) = fθ (N ₂ ) −fθ '(N ₂ ), d (θ) becomes fθ (N ₂
) And fθ '(N ₂ ) as a step error (where θ is the angular position of the X-ray source).

In this method, for example, a step-like error is smoothed using an exponential function described in (Equation 1) (where a is a control coefficient for controlling the smoothing of the exponential function).
A function of the forward and opposite rays, fθ (N) and fθ '(N) is the projection data gθ (N). Here, 1 <n <N.

When 1 <n <N ₂ gθ (n) = fθ (n) −pθ (N ₂ −n) (Equation 8) When N ₂ <n <N gθ (n) = fθ ′ (n) + Pθ (n−N ₂ ) (Equation 9) From (Equation 8) and (Equation 9), when n = N ₂ , gθ (N) becomes equal.

The following procedure is executed for each projection image. 1. By embedding zero according to (Equation 2), fθ The original half FO called (n)
V projection dataMeasurementI do. 2. Set of rays in the opposite direction fθ '(N) and pad it with zeros according to (Equation 3).
Inset. 3. According to (Equation 8) and (Equation 9), smoothing is performed based on the step-like error d (θ).
Apply a function (where d (θ) = fθ (N_Two) -Fθ '(N_Two)). 4.gθ Create (n) to form a set of N detector data. 5.gθ (N)To the conventional filter-corrected backprojection. 6. Steps 1 to 5 are repeated for all projection angles.

As one method of the present invention, the above procedure may be performed on any CT that can obtain rays in the opposite direction by interpolating other projection data from various angles.
Valid for scanners. In the case of a two-dimensional fan beam in which data for another half FOV can always be approximately calculated from the redundant fan beam projection data, this is a perfect case. If this technique is extended to 3D-VCT, complete interpolation can only occur on the mid-plane when using a circular orbit (ie, the x-ray plane with zero elevation) .

According to the results of the present invention, when applying the same technique to a VCT using a circular orbit, this technique is useful for cone angles in the range of ± 1.5 degrees or elevation angles as used herein . Operation is shown to be superior to conventional smoothing techniques (using 20 additional detectors beyond the location of the center ray). However, this is not the case for large cone angles. This is because the angular orientation of the ray in the opposite direction is considerably different from the measured data when a perfect detector is used. To understand this , assume a ray that passes through the center of rotation of the CT scanner and deviates from the detector mid- plane, and thus has a non-zero elevation angle. Therefore, the point between the source and the detector is rotated by 180 degrees. Obviously, in that case ray does not pass through the same region of the virtual patient. To address this situation, iterative techniques can be used to improve image quality. The procedure is shown below.

1. An initial three-dimensional image is obtained according to the above steps 1 to 6. As described immediately above, it is highly possible that artifacts will occur in the reconstructed image. 2. In the second iteration of the method, the forward projections are used to obtain "opposite ray" for each set of half-FOV projection data, and these are obtained according to steps 3-6 outlined above. in combination with measurement data obtained by the other half FOV projection to create a complete set of projection data. The smoothing function can be implemented for each row in the set of projection data . 3. Until reconstruction converges, that is, until or not improved to the quality of the image does not change, continue to step 2 of the process.

It should be noted that the present invention has been discussed with respect to particular embodiments. Therefore, the present invention is not limited to these embodiments. For example, a flat smoothed function discussed are not shall not meant to exclude other functions of the type may be utilized for this purpose. Furthermore, the method is not limited to a scanning protocol. That is, the method is applicable to both axial (axial) scan protocols (the patient table does not move during the scan ) and helical (spiral) scan protocols . One of ordinary skill in the art would understand how to utilize and extend these concepts to enable other area detector scan protocols useful for particular applications.

[Brief description of the drawings]

FIG. 1 is a block diagram of a CT system according to the present invention.

FIG. 2 is a diagram illustrating the excursion of a detector utilized in accordance with the method of the present invention.

FIG. 3 is a block diagram illustrating the method of the present invention according to a preferred embodiment.

[ Description of Signs ] 1 Computer 2 X-ray tube 3 Detector 4A Controller 4B Controller 6 Data acquisition system 10 Control algorithm

[Procedure amendment 2]

[Document name to be amended] Drawing

[Correction target item name] Figure 3

[Correction method] Change

[Contents of correction]

FIG. 3

──────────────────────────────────────────────────の Continued on the front page (51) Int.Cl. ⁷ Identification symbol FI Theme coat ゛ (Reference) G06T 1/00 420 G06T 1/00 420A 5/20 5/20 C H04N 1/04 G01T 1/161 C / / G01T 1/161 H04N 1/04 E (81) Designated country EP (AT, BE, CH, CY, DE, DK, ES, FI, FR, GB, GR, IE, IT, LU, MC, NL, PT, SE), OA (BF, BJ, CF, CG, CI, CM, GA, GN, GW, ML, MR, NE, SN, TD, TG), AP (GH, GM, KE, LS, MW) , SD, SL, SZ, TZ, UG, ZW), EA (AM, AZ, BY, KG, KZ, MD, RU, TJ, TM), AE, AL, AM, AT, AU, AZ, A, BB, BG, BR, BY, CA, CH, CN, CR, CU, CZ, DE, DK, DM, EE, ES, FI, GB, GD, GE, GH, GM, HR, HU, ID , IL, IN, IS, JP, KE, KG, KP, KR, KZ, LC, LK, LR, LS, LT, LU, LV, MA, MD, MG, MK, MN, MW, MX, NO, NZ, PL, PT, RO, RU, SD, SE, SG, SI, SK, SL, TJ, TM, TR, TT, TZ, UA, UG, US, UZ, VN, YU, ZA, ZW (72 ) Inventor Ahmad, Ishark N. United States, 12065, New York, Clifton Park, Kins Road, No. 629 No., 55th F term (reference) 2G088 EE02 FF02 GG19 GG20 JJ05 KK33 LL12 LL13 4C093 AA22 CA13 CA32 EB16 EB19 FD01 FD07 FD12 FE14 FE26 FF41 5B047 AA11 AA17 AA11 AA17 AB02 BA02 BB10 BA23 CH05 ACO5 AA01 BA17 CA20 EA10 UA20 VA01 XA10

Claims (2)

[Claims] 1. A three-dimensional computed tomography (VCT) system for acquiring projection data of a subject, comprising: an X-ray source for projecting X-rays onto the subject; and an X-ray projected from the X-ray source. And generating an electrical signal in response to incident X-rays, the width of the detector being referenced to a center position corresponding to the projection of the center of rotation of the CT system onto the detector. A data acquisition system for reading the electrical signal from the detector and converting it to a digital signal, receiving the digital signal from a data acquisition component, and executing a reconstruction algorithm. A computer capable of reconstructing an image while causing a data acquisition component to process the digital signal; Stereoscopic computer tomography system comprising a computer, a. 2. A method for acquiring projection data of a subject using a computed tomography (CT) system, comprising: projecting X-rays from the X-ray source onto the subject; Receiving X-rays projected in a projection view at a detector of a CT system, said detector comprising a plurality of detector elements;
Generating an electrical signal in response to X-rays incident on the detector; digitizing the electrical signal; for each projection view, a detector element closest to an isocenter of the CT system; Selecting a value Va for the detector element; estimating a value Vb for the detector element by interpolation from the opposite direction or forward projection in the same direction for the selected detector element; Selecting a smoothing function that can remove the difference between Va and Vb; applying the smoothing function to remove the difference between Va and Vb; Generating a smooth transition region by smoothing a step when combining the projection data of the above and the estimated projection data.