Patent Abstract:
A computed tomography imaging system includes a source of a conical beam of radiation and a two-dimensional detector array arranged on opposite sides of an axis of rotation. Projection data is acquired in a conventional manner as the source and detector array make a full rotation about an object. A conventional half-scan image reconstruction algorithm is applied to the projection data at a plurality of different center-view angles to produce a plurality of sub-images. Image segments, which are centered in each sub-image along the axis of the respective center-view angle, are selected and combined to form a cross-sectional image of the object. The regions of each sub-image preferably are defined by a weighting function.

Full Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     Not applicable. 
     STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT 
     Not applicable. 
     BACKGROUND OF THE INVENTION 
     The present invention relates to computed tomography (CT) imaging apparatus; and more particularly, to the acquisition of data from the separate x-ray detectors in 2D detector arrays. 
     In a current computed tomography system, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system, termed the “imaging plane.” The x-ray beam passes through the object being imaged, such as a medical patient, and impinges upon a row, or one-dimensional array of radiation detectors. The intensity of the transmitted radiation is dependent upon the attenuation of the x-ray beam by the object and each detector produces a separate electrical signal that is a measurement of the beam attenuation. The attenuation measurements from all the detectors are acquired separately to produce the transmission profile. 
     The source and detector array in a conventional CT system are rotated on a gantry within the imaging plane and around the object so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements from the detector array at a given angle is referred to as a “view” and a “scan” of the object comprises a set of views made at different angular orientations during one revolution of the x-ray source and detector. In a 2D scan, data is processed to construct an image that corresponds to a two dimensional slice taken through the object. The prevailing method for reconstructing an image from 2D data is referred to in the art as the filtered backprojection technique. This process converts the attenuation measurements from a scan into integers called “CT numbers” or “Hounsfield units”, which are used to control the brightness of a corresponding pixel on a cathode ray tube display. 
     In one type of multi-slice CT imaging system  10  as shown in FIG. 1, the x-ray beam  14  also fans out along the z-axis producing a “cone-beam”. The detectors  22  are arranged in a two dimensional array  20  that has multiple rows of detectors to acquire attenuation measurements in a plurality of slices disposed along the z axis. The backprojected slice image is reconstructed from the volumetric data using the Feldkamp algorithm. 
     The reconstructed volume for an axial scan is a cylindrically shaped region. FIG. 2 illustrates a cross sectional view of the volume coverage at the iso-channel plane (a plane passing through both the x-ray focal spot  16  and the detector iso-channels and parallel to the z-axis). The desired reconstructed volume is represented by a dashed rectangle  50 , having a height D that is equal to the height of the detector at the iso-center along the z-axis. From an image reconstruction point-of-view, every voxel, or element, in the image needs to be sampled by all projections to ensure artifact free reconstruction. However, the volume that satisfies this condition is depicted by a cross-hatched hexagon  52 . In practice, the reconstruction volume for each scan is limited to the thick line rectangle  54  enclosed inside the cross-hatched hexagon  52  in order to obtain a continuous reconstruction volume with multiple axial scans. As a result, the distance between the adjacent axial scans cannot be larger than the dimension of this inner rectangle along the along the z-axis. 
     If the distance from the source to iso-center is denoted by S and the radius of the reconstruction field of view (x-y) is denoted by R, the distance t between adjacent axial scans for continuous coverage is given by the expression:              t   ≤       (       S   -   R     S     )        D             (   1   )                                
     Thus, to obtain continuous coverage of an organ within the patient being scanned, the distance between adjacent scans is limited to about half of the detector coverage in the z-axis at the iso-center. This significantly reduces the volume coverage capability of the scanner. Therefore it is desirable to increase the reconstruction volume to the full distance D. However doing so necessitates extensive extrapolation of the projection data and thus introduces significant image artifacts because the stippled triangular areas  56  and  58  at each corner of the desired reconstruction volume  50  are not fully scanned. As seen in FIG. 2, when the emitter  12  and detector array  20  are oriented as illustrated, the corner areas  58  closest to the detector array are within the x-ray beam  14 . However, in this orientation, the triangular corner areas  56  closest to the emitter  12  are outside the x-ray beam  14 . When the emitter and detector assembly has rotated 180°. i.e. have reversed the illustrated positions, corner areas  58  are outside the x-ray beam and corner areas  56  lie within the beam. As a consequence, the triangular corner areas  56  and  58  are only partially scanned during a complete rotation of the emitter and detector array. This requires extrapolation of the data for these areas which generates artifacts in the reconstructed image. 
     Therefore, although it is desirable to increase the size of the reconstruction volume as much as possible, ideally to distance D, doing so with conventional processes introduces significant artifacts into the reconstructed image. As a consequence, an alternative reconstruction process which reduces such artifacts is desired. 
     SUMMARY OF THE INVENTION 
     A computed tomography imaging system, includes a source of a conical beam of radiation and a multi-row detector array arranged on opposite sides of an axis of rotation. That imaging system employs an image reconstruction method which comprises rotating the source and detector about the axis of rotation. While that rotating occurs, x-ray attenuation data samples are collected from the multi-row detector array at a plurality of projection angles β thereby producing a set of projection data. 
     A half-scan image reconstruction technique is applied to the set of projection data at a plurality of different center-view angles β 0  to produce a plurality of sub-images. The plurality of sub-images then are combined, such as by superimposition for example, to form a cross-sectional image of the object. 
     The preferred reconstruction technique includes weighting the set of projection data to produce a set of weighted data. For example, the weighting process applies a first weight to data samples within a first region centered about the respective center-view angle β 0 . A second weight is applied to data samples within predefined second regions on either side of the first region. A third weight is applied to data samples in the remainder of the set of projection data. The weighted data is filtered using a conventional filter for single slice scans. The filtered data are backprojected by known 3-D backprojection algorithms to produce a plurality of sub-images which are then combined by weighting functions to formulate the final image. 
    
    
     BRIEF DESCRIPTION OF THE OF THE DRAWINGS 
     FIG. 1 is a pictorial view of a gantry of a CT imaging system; 
     FIG. 2 illustrates the reduced volume coverage of a conventional cone-beam, multiple slice CT imaging system; 
     FIG. 3 is a block schematic diagram of a CT imaging system in which the present invention may be employed; 
     FIGS. 4,  5  and  6  depict three sub-images which are reconstructed according to the present invention; 
     FIG. 7 depicts a full scan image which is assembled from the sub-images; and 
     FIG. 8 graphically illustrates a weighting function employed in one techniques for producing a sub-image. 
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     With reference to FIGS. 1 and 3, a CT imaging system  10  includes an x-ray source  12  oriented to project a cone beam of x-rays  14  from a focal spot  16  through a patient  18  to be received by a two-dimensional detector array  20 . The two-dimensional detector array  20  includes a number of detector elements  22  arranged over the area of the detector array  20  in generally perpendicular columns and rows to detect a projected image of the x-rays  14  passing through the patient  18 . 
     The x-ray source  12  and the two-dimensional detector array  20  are mounted on either side of a gantry  24  so as to rotate about an axis of rotation  26  generally positioned within the patient  18 . The axis of rotation  26  forms the z-axis of a Cartesian coordinate system having its origin centered within the cone beam  14 . The plane defined by the x and y axes of this coordinate system thus defines a plane of rotation, specifically the gantry plane  28  of the gantry  24 . 
     Rotation of the gantry  24  is measured by angle β from an arbitrary reference position within the gantry plane  28 . Angle β varies between 0 and 2 π radians (360°). The x-rays of the cone beam  14  diverge from the gantry plane  28  by angle φ and diverge along the gantry plane  28  by angle φ. The two-dimensional detector array  20  is arranged as a section of the surface of a sphere having a center at the focal spot  16 , and its array of detector elements  22  is arranged to receive and make intensity measurements along the rays of the cone beam  14  throughout the angles of φ and θ of the cone beam  14 . the detector array  20  is comprised of a 2D array of detector elements  22  arranged in rows which extend along an in-slice dimension. Each row may include, for example, 1,000 separate detector elements, and the array  20  may include 16 rows disposed along the slice dimension. The detectors  22  may be gas or, solid state detectors which produce an electrical signal proportional to the x-ray flux received over the sample period. 
     Referring to FIG. 3, the control subsystem of the CT imaging system  10  has gantry associated control modules  30  which include: x-ray controller  32 , which provides power and timing signals to the x-ray source  12 , gantry motor controller  34 , which controls the rotational speed and position of the gantry  24 . A data acquisition system (DAS)  36  receives projection data from the two-dimensional detector array  20  and converts the data into digital form for later computer processing, while preserving the values of φ, θ, and the gantry angle β at which the data was taken. The x-ray controller  32 , the gantry motor controller  34  and the data acquisition system  36  are connected to computer  38 . The computer  38  also governs operation of a table motor control  37  which drives a motor that moves the patient table  39  along the z-axis  26 . 
     The computer  38  is a general purpose minicomputer programmed to acquire and manipulate projection data as will be described in detail below. The computer  38  is connected to an image reconstructor  40  which performs high speed image reconstruction according to methods known in the art. 
     The computer  38  receives commands and scanning parameters via operator console  42  which is generally a CRT display and keyboard that enables an operator to enter parameters for the CT scan and to display the reconstructed image. A mass storage device  44  provides a means for storing operating programs. 
     During data acquisition, the CT imaging system  10  functions as a conventional cone-beam system in gathering data. Specifically, the table  39  is held stationary as the x-ray emitter  12  and detector array  20  make a complete revolution around the gantry  24  about the axis of rotation  26 . At each of a plurality of angular positions β, the attenuation data from all the detectors  22  in array  20  are stored in the mass memory  44 . Upon completion of a full rotation, the computer commands the table motor control  37  to advance the table  39  to another position along the z-axis  26  and another rotational scan of the patient  18  is preformed. This process is repeated until the desired portion of the patient  18  has been fully scanned. 
     Then image reconstruction commences. The essence of the present invention in the reconstruction of a plurality of sub-images each representing a wedge of the circular slice image. Each sub-image is produced using a half-scan reconstruction technique at a different center-view angle β 0  spaced around a full 2π radian rotation. Each sub-image comprises a wedge portion of the half-scan reconstructed image that contains the least amount of artifacts. The details of the present invention will be explained using an exemplary procedure that employs three sub-images produced at center-view angles of π/3 radians, π radians, and 5π/3 radians. However, a greater number of sub-images can be utilized with commensurately smaller intervals between the center-view angles. For example, four sub-images having π/2 radian wedges could be produced utilizing four center-view angles β 0  of 0, π/2, π, and 3π/2 radians. The artifact reduction improves as the number of sub-images wedges increases, however, that also increases the amount of computation time to reconstruct the full slice image. The separate sub-images then are patched together to form the full slice image representing a cross section through the patient. 
     With reference to FIG. 4, a first half-scan reconstruction is performed at the center-view angle β 0 =π/3 radians. The projection data stored in mass storage  44  is weighted by the following function prior to filtering and application of the three-dimensional back-projection algorithm:            θ   n          (     γ   ,   β   ,     β   o       )       =     {                 β   -     β   o     +     π   /   2     +     γ   m           2        γ   m       -     2      γ                                β   o     -     π   /   2     -     γ   m       ≤   β   &lt;       β   o     -     π   /   2     +     γ   m     -     2      γ                     1   ,               β   o     -     π   /   2     +     γ   m     -     2      γ       ≤   β   &lt;       β   o     +     π   /   2     -     γ   m     -     2      γ                       β   o     +     π   /   2     +     γ   m     -   β         2        γ   m       +     2      γ                                β   o     +     π   /   2     -     γ   m     -     2      γ       ≤   β   &lt;       β   o     +     π   /   2     +     γ   m                     0   ,                      OTHERWISE                
            w   n          (     γ   ,   β   ,     β   o       )         =       3          θ   n   2          (     γ   ,   β   ,     β   o       )         -     2          θ   n   3          (     γ   ,   β   ,     β   o       )                                      
     where n is the detector row index (see FIG.  2 ), β O  is the center-view angle of a half scan reconstruction, y is the angle between a given detector ray and the iso-ray  27  (a line from the x-ray focal spot  16  to the detector array  20  which passes through the axis of rotation  26 , see FIG.  3 ), and β is the projection angle. The weighting of the projection data accounts for the greater likelihood of artifacts occurring the farther the data point is from the center-view angle. It will be understood that when other numbers of sub-images are employed, other weighting functions may be utilized. For example, helical interpolation or helical extrapolation algorithms developed for single slice CT can be utilized as weighting functions. 
     The weighted projection data is then filtered along the γ direction as is commonly done for single slice scans or for the Feldkamp reconstruction for a cone beam. 
     A three-dimension backprojection technique for a cone beam then is applied to the filtered projection data to create a first sub-image  60  depicted in FIG.  4 . The first sub-image  60  has an approximately 2π/3 radian wedge shaped first primary segment  62  opposite the center-view angle and centered on the axis  63  of that view angle which passes through the iso-center, i.e. the center of the wedge is at angle β 0 +π. The first primary segment  62  has a relatively minimal amount of artifacts and thus contains voxels which have not been attenuated by the weighting function. The image elements, or voxels, in smaller wedge shaped segments  64  on each side of the first primary segment  62  are attenuated by amounts that increase with increase in the angular distance from that first primary segment. Thus the image intensity in these smaller wedge segments  64  gradually decreases. The voxels in the remainder of the sub-image (the non-crosshatched portion  66 ) have been assigned a zero value by the weighting function. 
     There are two approaches which can be employed to produce the wedge shaped sub-image. The first is to use a conventional three-dimension backprojection technique to produce a full image representing the entire 2π slice region. Then another weighting function, depicted in FIG. 8, is applied to the full reconstructed image to suppress information that does not fall within the desired of the sub-image region. The weight applied to a given voxel is a function of the voxel&#39;s angular position in the image. The weighting function is centered opposite to the center-view angle, at angle β 0 +π. A weight of one is applied to the voxels in the primary segment  62  and the weight decreases in the smaller border wedge segments  64  the farther the voxel is from the primary segment  62 . A weight of zero is applied to voxels that are in region  66  in the full image. 
     The second approach produces a sub-image directly in the backprojection process. Here the backprojection is preformed only in a 2/3 π region corresponding to the sub-image area. Although this results in a more complex backprojection process, it could result in less computation as region  66  is not backprojected. 
     Next, this novel half-scan back projection process is applied a second time to the projection data using a center-view angle β 0  of π. The projection data is weighted, filtered and a second sub-image is reconstructed using the half-scan algorithm. The second sub-image  67 , shown in FIG. 5, has a second primary segment  68  which has a wedge shape of approximately 2π/3 radians opposite to the center-view angle and centered on the axis  65  of that view angle. The image intensity of the voxels in the second primary segment has not been attenuated by the weighting function. The second primary segment  68  is flanked by a pair of smaller wedge shaped segments  69  in which the image intensity gradually decreases going away from the second primary segment. 
     The half-scan back projection process is applied a third time to the stored projection data to produce a third sub-image  70  shown in FIG. 6 that is reconstructed at a center-view angle B 0 =5π/3. The third sub-image  72  has a third primary segment  72  which is a 2π/3 radian wedge centered on the axis  73  of and opposite to the center-view angle. The image intensity of the voxels in the third primary segment has not been attenuated by the weighting function. The third primary segment  72  is flanked by a pair of smaller wedge shaped segments  71  in which the image intensity gradually decreases going away from the third primary segment. The weighting functions for the three sub-images have the property that the summation of the three functions is unity within the reconstruction region of interest. 
     Once all three sub-images  60 ,  67  and  70  have been formed, the computer combines them to form the final full slice image  74  depicted in FIG.  7 . In essence the three sub-images are superimposed over each other. Because of the manner in which the data was weighted, the result is as thought the primary segments  62 ,  68  and  72  were cropped and placed adjacent one another into the final full slice image. The superimposition overlays the smaller wedge segments  64 ,  69  and  71  which feathers the transition between the primary segments  62 ,  68  and  72 . 
     It should be understood that the quality of each reconstructed sub-image is best at a location that corresponds to the iso-ray of the center-view. Image quality degrades gradually from that center line. Therefore, image quality will improve by using greater numbers of sub-images, each having a smaller wedge angle. For example, the image quality of the resultant image produced by five sub-images will be greater than that produced by three sub-images. However, this involves a trade-off between image quality and computation time. 
     The foregoing description was primarily directed to a preferred embodiment of the invention. Although some attention was given to various alternatives within the scope of the invention, it is anticipated that one skilled in the art will likely realize additional alternatives that are now apparent from disclosure of embodiments of the invention. For example, instead of the half scan weights, helical weighting functions can be used to produce sub-images. In addition, the above scheme can be applied only to the image slices near the outer rows of the detector. For the regions that are close to the detector center plane, simple row interpolation can be used. Accordingly, the scope of the invention should be determined from the following claims and not limited by the above disclosure.

Technology Classification (CPC): 0