Abstract:
One aspect of the present invention is a method for imaging an organ of a patient that includes steps of: scanning a volume of a patient&#39;s body including an organ of the patient with a computed tomographic (CT) imaging system having a radiation source and detector coupled to a rotating gantry, the detector array having a z-direction parallel to an axis of rotation of the gantry and an x-direction transverse to the z-direction; acquiring attenuation data from a plurality of staggered half detector segments of the detector array; and reconstructing an image including the patient&#39;s organ using the acquired attenuation data.

Description:
BACKGROUND OF THE INVENTION 
     This invention relates generally to methods and apparatus for computed tomographic cardiac imaging systems, and more particularly to methods and apparatus specialized for cardiac imaging with substantial component reuse. 
     In at least one known computed tomography (CT) imaging system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the “imaging plane”. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile. 
     In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a “view”. A “scan” of the object comprises a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector. In an axial scan, the projection data is processed to construct an image that corresponds to a two dimensional slice taken through the object. One method for reconstructing an image from a set of projection data is referred to in the art as the filtered back projection 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. 
     More particularly, and referring to  FIGS. 5 and 6 , a computed tomograph (CT) imaging system  10  is shown as including a gantry  12  representative of a “third generation” CT scanner. Gantry  12  has an x-ray source  14  that projects a beam of x-rays  16  toward a detector array  18  on the opposite side of gantry  12 . Detector array  18  is formed by detector elements  20  which together sense the projected x-rays that pass through an object  22 , for example a medical patient. Each detector element  20  produces an electrical signal that represents the intensity of an impinging x-ray beam and hence the attenuation of the beam as it passes through patient  22 . During a scan to acquire x-ray projection data, gantry  12  and the components mounted thereon rotate about a center of rotation  24 . Detector array  18  may be fabricated in a single slice or multi-slice configuration. In a multi-slice configuration, detector array  18  has a plurality of rows of detector elements  20 , only one of which is shown in FIG.  2 . 
     Rotation of gantry  12  and the operation of x-ray source  14  are governed by a control mechanism  26  of CT system  10 . Control mechanism  26  includes an x-ray controller  28  that provides power and timing signals to x-ray source  14  and a gantry motor controller  30  that controls the rotational speed and position of gantry  12 . A data acquisition system (DAS)  32  in control mechanism  26  samples analog data from detector elements  20  and converts the data to digital signals for subsequent processing. An image reconstructor  34  receives sampled and digitized x-ray data from DAS  32  and performs high speed image reconstruction. The reconstructed image is applied as an input to a computer  36  which stores the image in a mass storage device  38 . 
     Computer  36  also receives commands and scanning parameters from an operator via console  40  that has a keyboard. An associated cathode ray tube display  42  allows the operator to observe the reconstructed image and other data from computer  36 . The operator supplied commands and parameters are used by computer  36  to provide control signals and information to DAS  32 , x-ray controller  28  and gantry motor controller  30 . In addition, computer  36  operates a table motor controller  44  which controls a motorized table  46  to position patient  22  in gantry  12 . Particularly, table  46  moves portions of patient  22  through gantry opening  48 . 
     In a multislice imaging system  10 , detector array  18  comprises a plurality of parallel detector rows, wherein each row comprises a plurality of individual detector elements  20 . An imaging system  10  having a multislice detector  18  is capable of providing a plurality of images representative of a volume of object  22 . Each image of the plurality of images corresponds to a separate “slice” of the volume. The “thickness” or aperture of the slice is dependent upon the thickness of the detector rows. 
     For example, and referring to  FIGS. 7 and 8 , a multislice detector array  18  includes a plurality of detector modules  50 . Each detector module  50  has a plurality of detector elements  20 . Particularly, each x-ray detector module  50  includes a plurality of photodiodes  52 , a semiconductor device  54 , and at least one flexible electrical cable  56 . Scintillators  58 , as known in the art, are positioned above and adjacent photodiodes  52 . Photodiodes  52  may be individual photodiodes or a multidimensional photodiode array. Photodiodes  52  are optically coupled to scintillators  58  and generate electrical outputs on lines  60  representative of light generated by scintillators  58 . Each photodiode  52  produces a separate electrical output  60  that is a measurement of the beam attenuation for a specific detector element  20 . In one known embodiment, photodiode output lines  60  from each detector module  50  are located at the top and bottom of the photodiode array. 
     Semiconductor device  54 , in one embodiment, includes two semiconductor switches  62  and  64 . Switches  62  and  64  each include a plurality of field effect transistors (FETs) (not shown) arranged as a multidimensional array. Each FET includes an input line electrically connected to a photodiode output  60 , an output line, and a control line (not shown). FET output and control lines are electrically connected by flexible cable  56 . Particularly, one-half of photodiode output lines  60  are electrically connected to each FET input line of switch  62  with the remaining one-half of photodiode output lines  60  electrically connected to FET input lines of switch  64 . 
     Flexible electrical cable  56  includes a first end (not shown), a second end (not shown) and a plurality of electrical wires  66  traveling therebetween. Cable  56  may, for example, be a single cable having multiple first ends  68  and  70  or in another known embodiment, may include multiple cables (not shown) each having a first end (not shown). FET output and control lines are electrically connected to cable  56  by wire bonding. Cable first ends  68  and  70  are secured to detector module  50  using mounting brackets  72  and  74 . Detector modules  50  are secured to detector array  18  using rails  76  and  78 . 
     One known detector array  18  is arcuate. However, and referring to  FIG. 9 , detector arrays are represented in simplified drawings by a flat, two-dimensional representation of the area exposed to radiation beam  16 . In such representations, the axis of rotation of gantry  12  defines a z-direction of detector array  18 . A transverse direction, i.e., the direction in which each row of detector elements  20  extends, defines an x-direction. In  FIG. 9 , rows (not separately shown) of detector elements  20  extend linearly in the plane of the paper, but each row, in reality, follows the arc of detector array  18 . Centerline  80  on  FIG. 9  represents an imaginary line of a radiation beam  16  passing through an axis of rotation of gantry  12 . Detector array  18  is at least approximately symmetric about centerline  80 , i.e., it is operationally insignificant if there is a slight asymmetry in the number of detector cells  20  on each side of centerline  80 . 
       FIG. 9  is not drawn to scale. In addition, only a few detector modules  50  are represented in FIG.  9 . In one known imaging system, fifty-seven detector modules  50 , each having 16 rows of 16 elements, are assembled in detector array  18 . 
     One problem with known imaging systems  10  is that they do not have detector arrays  18  that provide a sufficient number of rows of detector elements  20  to image a heart or other organ of patient  22  in a single revolution of radiation source  14  and detector array  18 . Thus, known cardiac CT imaging methods require multiple revolutions and a substantial amount of time (relative to a cardiac cycle). 
     It would, in principle, be possible to image an entire heart in a single revolution using a larger detector array  18  that had a sufficient number of detector rows to capture attenuation data from all parts of the heart. A CT imaging system  10  having such a detector array  18  would provide the advantage of reducing a patient&#39;s total radiation dosage during a cardiac scan. However, to provide acceptable resolution for diagnostic purposes, detector array  18  would have to generate massive amounts of data from a large total number of detector elements. Providing a data acquisition system  32  capable of handling such a large amount of data would be costly. 
     It is known to selectively combine data from a plurality of adjacent detector rows (i.e., a “macro row”) to obtain images representative of slices of different selected thicknesses, which also reduces the amount of data that must be handled by data acquisition system  32  during a scan. If a detector array  18  large enough to image an entire heart during a single revolution were provided, rows could be combined to reduce the amount of data generated. Alternately, detector elements  20  could simply be made larger to provide increased coverage without providing massive amounts of data. However, either of these alternatives runs a significant risk of reducing resolution to unacceptable levels. 
     It would therefore be desirable to provide methods and apparatus to provide satisfactory CT cardiac imaging with a minimum number of revolutions of an x-ray source and detector. It would further be desirable if such imaging could be accomplished with a single revolution. It would also be desirable to reduce the amount of data collected during such a cardiac CT scan without making unacceptable sacrifices in image quality and resolution. 
     BRIEF SUMMARY OF THE INVENTION 
     There is therefore provided, in one embodiment of the present invention, a method for imaging an organ of a patient. The method includes steps of: scanning a volume of a patient&#39;s body including an organ of the patient with a computed tomographic (CT) imaging system having a radiation source and detector coupled to a rotating gantry, the detector array having a z-direction parallel to an axis of rotation of the gantry and an x-direction transverse to the z-direction; acquiring attenuation data from a plurality of staggered half detector segments of the detector array; and reconstructing an image including the patient&#39;s organ using the acquired attenuation data. 
     The above-described embodiment is capable of providing satisfactory CT cardiac imaging with a minimum number of revolutions of an x-ray source and detector of the CT imaging system, without unacceptable sacrifices in image quality and resolution. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a simplified view of an embodiment of a detector array of the present invention as seen looking towards the detector from the x-ray source. 
         FIG. 2  is a pictorial view of a center detector array module of the detector array embodiment of FIG.  1 . 
         FIG. 3  is a simplified view of a detector array embodiment having a staggered/offset rail joining two detector arrays. 
         FIG. 4  is a simplified view of one half detector segment of the embodiment of FIG.  1 . (Some of the half detector segments are mirror images of the one represented in  FIG. 4. ) 
         FIG. 5  is a pictorial view of a prior art CT imaging system embodiment. 
         FIG. 6  is a block schematic diagram of the prior art system illustrated in FIG.  1 . 
         FIG. 7  is a perspective view of a prior art CT system detector array embodiment. 
         FIG. 8  is a perspective view of a prior art 16×16 detector module of the detector array embodiment of FIG.  7 . 
         FIG. 9  is a simplified view of the prior art CT system detector array embodiment illustrated in FIG.  7 . 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring to the simplified representation of  FIG. 1 , a detector array embodiment  82  of the present invention is used in a CT imaging system such as imaging system  10  to image a heart or other organ of patient  22 . ( FIG. 1  is simplified similarly to  FIG. 9. ) Detector array  82  replaces, or is provided as original equipment in imaging system  10  of  FIG. 5  in place of detector array  18 . Detector array embodiment  82  makes advantageous use of the fact that data from only one half of a detector arc rotated around patient  22  is necessary and is equivalent to half scan sampling and image reconstruction. (Each detector cell  20  in half scan sampling receives and measures radiation from at least a 180° arc around patient  22 .) Detector array  82  comprises a plurality of half detector segments  84  staggered on left and right sides of centerline  80  (more precisely, on the positive and the negative x-directions from centerline  80 ). In one CT imaging system  10  embodiment, centerline  80  is defined as an imaginary line of a radiation beam  16  passing through an axis of rotation of gantry  12 . Without reference to imaging system  10 , a centerline  80  of a staggered detector array  82  can be defined as an imaginary line parallel to the z-direction that bisects detector array  82  in the x-direction. Half detector segments  84  abut one another in regions about centerline  80 . 
     Another embodiment of detector array  82  comprises a plurality of half-detector segments  84  on the same side of centerline  80 . (In other words, half-detector segments  84  are not staggered.) However, the staggered embodiment of  FIG. 1  provides spaces  86  between half-detector segments. Spaces  86  provide space for prior art detector modules  50  to be used by providing room for rails  88 , and  90  to hold modules  50  in array  82 . In addition, spaces  86  allow flexible electrical cables  56  to be run out from detector array  82 . In half-detector segments  84 , modules  50  have four edges, and are each abutted by at most two other modules. 
     Referring to  FIG. 2 , center detector module  92  of detector array  82  is constructed differently from prior art modules  20 . Because of the limited space available for flexible cables  56  and the presence of adjacent modules in the z-direction, center detector modules  92  are configured so that its flexible electrical cable  56  runs in the x-direction rather than the z-direction, as mounted. To accommodate this construction, electrical output lines  60  (not shown in  FIG. 2 ) and a semiconductor switch  62  are located at one side of detector module  92 , in contrast to prior art detector module  50 , on which they are located at the top and bottom of the module. In the embodiment of  FIG. 2 , all signals are handled by one flexible electrical cable  56  extending in one direction and one semiconductor switch  62 . This permits the use of detector modules  92  in each half detector segment simply by orienting it in an appropriate direction, as each detector module  92  has butt joints on three other edges. In one embodiment, detector module  92  has a wider electrical cable  56  than detector modules  50  that wraps tightly around the free edge of module  92  so as not to interfere with flexible electrical cables  56  of other modules  50 . In one embodiment, cable  56  is shaped with a pre-formed right angle bend. Also in one embodiment, center detector modules  92  at x-extremities of detector array  82  have an extra mounting flange (not shown) for mounting to a collimator rail rather than to a third butt joint. 
     Center detector modules  92  need not have the same number of detector elements as detector modules  50 , and are provided to reduce image center artifacts. Thus, in one embodiment, detector modules  92  straddle centerline  80  in each half detector segment  84  and have sixteen detector cells  20  in the z-direction and fourteen in the x-direction. Also in one embodiment, detector cells  20  are paired (i.e., two are combined by hardwiring to produce a single output) in the x-direction. The two-cell “space” in the x-direction (i.e., fourteen detector cells  20  rather than sixteen) provides space for photodiode  52  signal routing and flexible cable  56  termination. Detector modules  92  having a greater or lesser number of detector cells  20  in the x-direction are used in other embodiments. The number of cells  20  is selected to ensure that the center of the field of view of imaging system  10  is adequately sampled. 
     As shown in  FIG. 1 , two types of rails  88 , and  90  are used in the construction of detector array  82  and form a portion of a post-patient collimator. Detector modules  50  of detector array  82  are mounted on rails  88  and  90  in a manner similar to that of prior art detector module  18 , for example, by screws passing through detector modules  50  into threaded holes in the rails. Rail  88  is unremarkable, and extends across an entire length of a half detector segment  84  in the x-direction. Rails  90  extend across most of the length of a half-detector segment except for a portion at which it abuts a detector module  92  of an adjacent half detector segment  84 . At this point, as shown in phantom, it extends beneath center detector module  92  diagonally in one embodiment, and continues as a mounting rail for another half detector segment  86 . Center modules  92  are mounted either by adhering them to rails  90  as they run under modules  92 . In another embodiment of detector array  82 , they are mounted on a free edge (i.e., the edge having flexible cable  56  attached). Thus, rails  88  and  90  are mounted in front of detector modules  50 , and rails  90  extend behind center modules  92   
     In one embodiment, detector modules  50  and  92  of detector array  82  are removable and replaceable. 
     Post-patient collimator plates  102  are used in one embodiment. Plates  102  are conventional except over center detector modules  92 , where each extends the full z-direction thickness of detector  82 , i.e., between both rails  88  and over a plurality of center detector modules  92 . (Conventional plates  102  extend in the z-direction over only a single detector module  50 .) Wires  104  of the post-patient collimator extend transverse to the post-patient collimator plates and present no special construction difficulty. Only a few post patient collimator plates  102  and wires  104  are represented in FIG.  1 . 
     In one embodiment, center modules  92  sit flush over rails  90 . 
     The mounting arrangement shown in  FIG. 1  is only exemplary. Moreover, detector array embodiments of the present invention are scalable, for example, in that they can use any number of staggered half-detector segments  84 .  FIG. 3  represents another detector array embodiment  94  having a slightly different form than detector array embodiment  82  of FIG.  1 . This embodiment has only two half detector segments  84 . In addition, rail  96  is wide enough to support two center detector modules  92 . 
     One embodiment of cardiac CT imaging system  10  utilizing detector array  82  instead of multislice detector array  18 . This embodiment produces a volume of data output similar to that of a standard eight slice imaging system  10  using prior art multislice detector array  18 . 
     For example, in one embodiment and referring to  FIGS. 2 and 4 , detector modules  92  have 16 cells in Z and 7 paired cells in X, for a total of 112 outputs per module. Detector elements  50  in region  98  adjacent to center detector modules  92  have a minimum of 112 detector cells  20  in an x-direction and a 13.04 cm field of view (FOV). A half field of view (FOV) at 541 mm to center of rotation  58  is 6.52 degrees, or about 0.0618 degrees per cell. A total FOV of a heart of patient  22  is therefore 13 cm. 
     In one embodiment, 7 detector modules  50  each having an array of 16×16 detector cells  20  is used in region  98  adjacent to the center of gantry rotation to provide a 13.04 cm FOV. Detector cells  20  in this embodiment provide 1.25 mm resolution in the z-direction. Cells in the x-direction are paired (i.e., their electrical outputs are connected together) so that there are only 128 distinct outputs per detector module  50 . Pairing cells  20  in the x-direction allows a standard detector module  50  to be used throughout detector array  82  with only minor modification. For example, detector modules  50  in one embodiment are hard wired in pairs. In another embodiment, FET arrays  62  and  64  are used in place of hard wiring so that the gain of all pixels can be calibrated. The construction of detector modules  50  is otherwise similar to that of such modules in known multi-slice imaging systems. 
     Hard wired pairing (or otherwise combining outputs of detector modules  50 ) reduces the number of DAS  32  data inputs required for processing cardiac images. In addition, summing in the x-direction gives more isotropic voxels in image space when using detector cells  20  dimensioned as described herein. The total number of detector cells  20  in region  98  is thus 7×128=896 cells per array. 
     A second region  100  of half detector segment  84  supports reconstruction of an entire FOV. However, region  100  is not required to provide cardiac imaging details, and thus can provide much lower detector cell sampling than does region  98 . For example, in one embodiment, a 48 cm FOV at 541 mm to isocenter  24  is 26.34 degrees, or 0.0618 degrees per detector cell  20 . However, data from cells  20  in each module  50  in region  100  is combined so that each module provides a single output for each row. In other words, all the cells in each module are combined in the x-direction, but the z-direction resolution is still 1.25 mm. Thus, each module  50  provides 16 outputs. In one embodiment, summation of cells in the x-direction is performed within modules  60 . In another embodiment, summation is performed in a backplane of DAS  32 . In either of these two embodiments, a total of 426 detector cells are to the left of the center of gantry rotation, or 426 cells/16 cells per module=26.63 or twenty-seven total modules  50 . Thus, there are twenty modules  50  in region  100 , because seven modules  50  are used in region  98 . With twenty modules  50  and only sixteen outputs per module  50 , there are effectively 320 cell outputs in region  100 . 
     By combining detector cells  20  in the manner described above, detector array  82  provides a relatively higher spatial resolution near centerline  80  and a relatively lesser spatial resolution distal to centerline  80 . 
     A known DAS  32  from an eight-slice CT imaging system includes 48 boards having 128 channels per board, which provides sufficient capability for processing 48×128=6144 detector cells  20 . Thus, the known DAS  32  provides sufficient processing capability for 4.63 half detector segments  84  (6144 cells/1328 cells per detector array=4.63 detector arrays). In one embodiment, however, a detector array  82  having five half detector segments  84  is provided for imaging to provide cardiac coverage of 13 cm (X) by 10 cm (Z). Thus, only a few additional DAS  32  boards are required for the additional channels needed. In another embodiment, additional cells  20  in a portion of region  98  adjacent to region  100  are summed to further reduce the amount of data output from the detector array without significantly sacrificing image quality. This embodiment also requires a few additional DAS 32 channels beyond that provided by the known 8 slice CT imaging system. In another embodiment having less cell sampling in regions  100  and/or fewer overlapping cells  20  in regions  98 , no additional DAS  32  boards or channels are required. 
     A modified bowtie can reduce x-ray dosage in the outer low resolution portion of the patient. 
     In one embodiment of the present invention, to image an organ of patient  22 , a volume of the body of patient  22  including the organ of interest is scanned with a computerized imaging system  10  that uses, instead of detector array  18 , a detector array  82  of the present invention. Attenuation data is acquired from a plurality of staggered half-detector segments  84  of detector array  82 , and an image of the organ of patient  22  is reconstructed using the acquired attenuation data. 
     It is thus clear that the embodiments described herein provide satisfactory CT cardiac imaging with a minimum number of revolutions of an x-ray source and detector, or, in some embodiments, with only a single revolution. Moreover, the amount of data collected during such a cardiac CT scan is reduced to levels that can be handled by known data acquisition systems with little or no augmentation, without making unacceptable sacrifices in image quality and resolution. 
     While the invention has been described in terms of various specific embodiments, those skilled in the art will recognize that the invention can be practiced with modification within the spirit and scope of the claims.