Abstract:
The present invention is directed to a CT detector array having uniform cross-talk. Discontinuities in cross-talk between adjacent CT detectors of a CT detector array are minimized by increasing the cross-talk at the boundaries of adjacent CT detectors. Discontinuities throughout a CT detector contribute to artifact presence in a final reconstructed image, therefore, it is preferred that cross-talk throughout the CT detector array be relatively uniform. Reducing the width of reflector material between adjacent CT detectors increases the cross-talk between the CT detectors. This increase in cross-talk offsets the reduced cross-talk that typically occurs between scintillators, optical epoxy layers, and photodiodes at the CT detector interface. Cross-talk may also be increased by reducing the amount of chrome deposited in the reflector between CT detectors or reducing the levels of titanium oxide typically used in reflector layers.

Description:
BACKGROUND OF INVENTION  
         [0001]    The present invention relates generally to diagnostic imaging and, more particularly, to a CT detector array having uniform cross-talk. More particularly, the invention is directed to a CT array constructed such that the optical cross-talk through the reflector between CT detectors is purposely increased so as to offset reduced electrical cross-talk and coupler layer cross-talk typically present at the interface of CT detectors. Increasing the cross-talk through the reflector between adjacent CT detectors reduces cross-talk discontinuities that contribute to artifact presence in a final reconstructed diagnostic image.  
           [0002]    Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage.  
           [0003]    Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.  
           [0004]    Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom.  
           [0005]    Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.  
           [0006]    “Cross-talk” between detector cells of a CT detector is common. “Cross-talk” is generally defined as the communication of data between adjacent cells of a CT detector. Generally, cross-talk is sought to be reduced as cross-talk leads to artifact presence in the final reconstructed CT image and contributes to poor spatial resolution. Typically, four different types of cross-talk may result within a single CT detector. X-ray cross-talk may occur due to x-ray scattering between scintillator cells. Optical cross-talk may occur through the transmission of light through the reflectors that surround the scintillators. Known CT detectors utilize a contiguous optical coupling layer(s), typically epoxy, to secure the scintillator array to the photodiode array. Cross-talk, however, can occur as light from one cell is passed to another through the contiguous layer. Electrical cross-talk can occur from unwanted communication between photodiodes.  
           [0007]    Cross-talk between adjacent elements of a CT detector is relatively uniform throughout the single CT detector. However, at the junction of one CT detector to another, there is generally a drop in cross-talk. The drop in cross-talk results from increased reflector material and anti-cross-talk matter disposed between CT detectors or modules as well as reduced electrical cross-talk and coupler layer cross-talk customarily present at the interface or boundaries of adjacent CT detectors. While cross-talk reduction is generally preferred between elements of single CT detector, a drop of cross-talk between CT detectors results in cross-talk discontinuities throughout the CT detector array. These discontinuities negatively affect the final reconstructed image.  
           [0008]    Additionally, backlit photodiodes are commonly being used in CT detectors. Backlit diodes are particularly susceptible to electronic cross-talk. As a result, at the interface or junction of CT detectors, there is a greater discontinuity in cross-talk as backlit photodiodes intrinsically are discontinuous in cross-talk at the boundaries of adjacent CT detectors. The implementation of backlit photodiodes, however, is preferred as backlit photodiodes have better tileability and improved interconnectivity than traditional photodiodes.  
           [0009]    CT detector array construction typically results in a greater reflector thickness between CT detectors than that found between detector elements of the CT detector. This results in a very finite space that limits the amount of scintillator surface area. If the junction between CT detectors was constructed to match or be less than the width between detector elements of a single CT detector, more of the array space can be devoted to scintillator surface area thereby improving x-ray detection. Additionally, reducing the thickness of reflector material disposed between CT detectors provides improved module to module tolerances for detector manufacturability.  
           [0010]    Therefore, it would be desirable to design a CT detector array having uniform cross-talk so as to reduce discontinuities typically found at the interface of adjacent CT detectors. It would also be desirable to design a CT detector array with reduced reflector width between CT detectors thereby improving manufacturing and detector element placement.  
         BRIEF DESCRIPTION OF INVENTION  
         [0011]    The present invention is directed to a CT detector array having uniform cross-talk overcoming the aforementioned drawbacks. Discontinuities in cross-talk between adjacent CT detectors of a CT detector array are minimized by increasing the cross-talk at the boundaries of adjacent CT detectors. Discontinuities throughout a CT detector contribute to artifact presence in a final reconstructed image, therefore, it is preferred that cross-talk throughout the CT detector array be relatively uniform. Reducing the width of reflector material between adjacent CT detectors increases the cross-talk between the CT detectors. This increase in cross-talk offsets the reduced cross-talk that typically occurs between scintillators, optical epoxy layers, and photodiodes at the CT detector interface. Cross-talk may also be increased by reducing the amount of chrome deposited in the reflector between CT detectors or reducing the levels of titanium oxide typically used in reflector layers.  
           [0012]    Therefore, in accordance with the present invention, a CT detector array includes a plurality of CT detectors arranged to receive x-rays impinged by a subject and output electrical signals to a data acquisition system. Each of the CT detectors includes a plurality of detector elements. Each element includes a scintillator and a photodiode optically coupled to one another. The CT detectors are aligned relative to one another such that cross-talk throughout the plurality of CT detectors is substantially uniform.  
           [0013]    In accordance with another aspect of the present invention, a CT system includes a rotatable gantry having a bore centrally disposed therein. A table is provided and is movable fore and aft through the bore and configured to position a subject for CT data acquisition. A high frequency electromagnetic energy projection source is positioned within the rotatable gantry and is configured to project high frequency electromagnetic energy toward the subject. A detector array is disposed within the rotatable gantry and configured to detect high frequency electromagnetic energy projected thereat and impinged by the subject. The detector array is configured to be absent substantial cross-talk discontinuities.  
           [0014]    According to another aspect of the invention, a method of manufacturing a CT detector array includes the steps of forming a plurality of CT detectors in an array and connecting the plurality of detectors to one another such that cross-talk throughout the connected plurality of CT detectors is substantially uniform. Each CT detector includes a plurality of detector elements.  
           [0015]    Various other features, objects and advantages of the present invention will be made apparent from the following detailed description and the drawings. 
       
    
    
     BRIEF DESCRIPTION OF DRAWINGS  
       [0016]    The drawings illustrate one preferred embodiment presently contemplated for carrying out the invention.  
         [0017]    In the drawings:  
         [0018]    [0018]FIG. 1 is a pictorial view of a CT imaging system.  
         [0019]    [0019]FIG. 2 is a block schematic diagram of the system illustrated in FIG. 1.  
         [0020]    [0020]FIG. 3 is a perspective view of one embodiment of a CT system detector array.  
         [0021]    [0021]FIG. 4 is a perspective view of one embodiment of a detector.  
         [0022]    [0022]FIG. 5 is illustrative of various configurations of the detector in FIG. 4 in a four-slice mode.  
         [0023]    [0023]FIG. 6 is a perspective view of adjacent CT detectors of a CT system detector array similar to that shown in FIG. 3.  
         [0024]    [0024]FIG. 7 is a pictorial view of a CT system for use with a non-invasive package inspection system. 
     
    
     DETAILED DESCRIPTION  
       [0025]    The operating environment of the present invention is described with respect to a four-slice computed tomography (CT) system. However, it will be appreciated by those skilled in the art that the present invention is equally applicable for use with single-slice or other multi-slice configurations. Moreover, the present invention will be described with respect to the detection and conversion of x-rays. However, one skilled in the art will further appreciate that the present invention is equally applicable for the detection and conversion of other high frequency electromagnetic energy. The present invention will be described with respect to a “third generation” CT scanner, but is equally applicable with other CT systems.  
         [0026]    Referring to FIGS. 1 and 2, a computed tomography (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 the gantry  12 . Detector array  18  is formed by a plurality of detectors  20  which together sense the projected x-rays that pass through a medical patient  22 . Each detector  20  produces an electrical signal that represents the intensity of an impinging x-ray beam and hence the attenuated beam as it passes through the 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 .  
         [0027]    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 an 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 detectors  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 reconstruction. The reconstructed image is applied as an input to a computer  36  which stores the image in a mass storage device  38 .  
         [0028]    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  and gantry  12 . Particularly, table  46  moves portions of patient  22  through a gantry opening  48 .  
         [0029]    As shown in FIGS. 3 and 4, detector array  18  includes a plurality of scintillators  57  forming a scintillator array  56 . A collimator (not shown) is positioned above scintillator array  56  to collimate x-ray beams  16  before such beams impinge upon scintillator array  56 .  
         [0030]    In one embodiment, shown in FIG. 3, detector array  18  includes 57 detectors  20 , each detector  20  having an array size of 16×16. As a result, array  18  has 16 rows and 912 columns (16×57 detectors) which allows 16 simultaneous slices of data to be collected with each rotation of gantry  12 .  
         [0031]    Switch arrays  80  and  82 , FIG. 4, are multi-dimensional semiconductor arrays coupled between scintillator array  56  and DAS  32 . Switch arrays  80  and  82  include a plurality of field effect transistors (FET) (not shown) arranged as multi-dimensional array. The FET array includes a number of electrical leads connected to each of the respective photodiodes  60  and a number of output leads electrically connected to DAS  32  via a flexible electrical interface  84 . Particularly, about one-half of photodiode outputs are electrically connected to switch  80  with the other one-half of photodiode outputs electrically connected to switch  82 . Additionally, a reflector layer (not shown) may be interposed between each scintillator  57  to reduce light scattering from adjacent scintillators. Each detector  20  is secured to a detector frame  77 , FIG. 3, by mounting brackets  79 .  
         [0032]    Switch arrays  80  and  82  further include a decoder (not shown) that enables, disables, or combines photodiode outputs in accordance with a desired number of slices and slice resolutions for each slice. Decoder, in one embodiment, is a decoder chip or a FET controller as known in the art. Decoder includes a plurality of output and control lines coupled to switch arrays  80  and  82  and DAS  32 . In one embodiment defined as a 16 slice mode, decoder enables switch arrays  80  and  82  so that all rows of the photodiode array  52  are activated, resulting in 16 simultaneous slices of data for processing by DAS  32 . Of course, many other slice combinations are possible. For example, decoder may also select from other slice modes, including one, two, and four-slice modes.  
         [0033]    As shown in FIG. 5, by transmitting the appropriate decoder instructions, switch arrays  80  and  82  can be configured in the four-slice mode so that the data is collected from four slices of one or more rows of photodiode array  52 . Depending upon the specific configuration of switch arrays  80  and  82 , various combinations of photodiodes  60  can be enabled, disabled, or combined so that the slice thickness may consist of one, two, three, or four rows of scintillator array elements  57 . Additional examples include, a single slice mode including one slice with slices ranging from 1.25 mm thick to 20 mm thick, and a two slice mode including two slices with slices ranging from 1.25 mm thick to 10 mm thick. Additional modes beyond those described are contemplated.  
         [0034]    Referring now to FIG. 6, a pair of adjacent CT detectors is shown. The CT detectors are similar to that which was described with respect to FIG. 4 and, accordingly, like numbers will be used where appropriate. Further, as the construction of each CT detector is similar, suffixes “a” and “b” will be used in the description of the CT detectors of FIG. 6. It should be noted that the illustrated CT detectors comprise only a portion of the CT detector array illustrated and described with respect to FIG. 3.  
         [0035]    As shown in FIG. 6, CT detectors  20   a ,  20   b  are positioned and aligned relative to one another. Additional CT detectors (not shown) are then used to form a CT detector array, FIG. 3. Each CT detector  20   a ,  20   b  is formed of detector elements, each of which includes a scintillator array  56   a ,  56   b  and a photodiode array  52   a ,  52   b  that are optically coupled to one another via an optically coupling layer (not shown). Reflector elements  86   a ,  86   b  are disposed between adjacent scintillators to form a reflector layer. The reflector elements are typically doped with chromium oxide (chrome) to reduce cross-talk emissions and titanium dioxide to improve light reflectivity.  
         [0036]    Typically, the CT detectors are formed individually and later aligned and oriented to form a CT detector array. As a result, a reflector element extends along each side of the CT detector. When two CT detectors are then aligned next to one another, a composite reflector wall  88  is formed having a thickness or width that exceeds the width of the reflector elements  86  positioned between adjacent scintillators. This increase in reflector thickness, which is typically doped with anti-cross-talk particles, causes a decrease in cross-talk between CT detector  20   a  and CT detector  20   b  relative to the cross-talk between detector elements. For example, the increased reflector width results in larger amount of anti-cross-talk particles, such as chromium oxide, between the CT detectors. The anti-cross-talk particles absorb optical cross-talk introduced in the reflector from each CT detector  20   a ,  20   b . The increased cross-talk absorption together with reduced electrical cross-talk and reduced optical layer cross-talk causes a discontinuity of cross-talk at the interface of adjacent CT detectors.  
         [0037]    Reducing the thickness of reflector wall  88  results in a reduction of the number of cross-talk absorption particles between CT detectors. As a result, cross-talk between adjacent CT detectors increases. The reduction in reflector wall  88  thickness is such that the cross-talk between CT detectors is similar to the cross-talk between individual detector elements of a CT detector. The uniformity of cross-talk throughout the CT detector array that results from implementation of a thinner reflector wall  88  between CT detectors minimizes any cross-talk discontinuities which reduce artifacts in the final reconstructed image.  
         [0038]    Reducing the thickness of reflector wall  88  is only one example whereupon uniform cross-talk in the CT detector array may be achieved. For example, standard reflector wall  88  thickness may be used but doped with less cross-talk absorption components to increase the cross-talk between CT detectors. Alternately, the amount of titanium dioxide used in the reflector may be reduced such that the reflectivity of reflector wall  88  is reduced. A reduction in the reflectivity characteristics of reflector wall  88  increases cross-talk between detectors.  
         [0039]    To achieve a uniform or consistent cross-talk level across the CT detector array, the reflector thickness, reflectivity, and/or cross-talk absorption is controlled such that the amount of cross-talk that results as a result of changes in the reflector between adjacent CT detectors offsets the reduction in optical cross-talk and electrical cross-talk typically experienced between the CT detectors at the boundary or interface of adjacent detectors. Other advantages of this construction include more overall uniform spacing across detector boundaries, increased scintillator area for improved x-ray quantum detection efficiency, generally referenced, “detector QDE”. That is, reducing the width of the reflector wall between CT detectors allows for larger scintillator x-ray reception surface. As a result, more x-rays are received by scintillation material and may be used for the imaging process. Thinner reflector elements between adjacent CT detectors also allows for more detector to detector tolerances that improve detector manufacturability.  
         [0040]    Referring now to FIG. 7, package/baggage inspection system  100  includes a rotatable gantry  102  having an opening  104  therein through which packages or pieces of baggage may pass. The rotatable gantry  102  houses a high frequency electromagnetic energy source  106  as well as a detector assembly  108  having scintillator arrays comprised of scintillator cells similar to that shown in FIGS.  6  or  7 . A conveyor system  110  is also provided and includes a conveyor belt  112  supported by structure  114  to automatically and continuously pass packages or baggage pieces  116  through opening  104  to be scanned. Objects  116  are fed through opening  104  by conveyor belt  112 , imaging data is then acquired, and the conveyor belt  112  removes the packages  116  from opening  104  in a controlled and continuous manner. As a result, postal inspectors, baggage handlers, and other security personnel may non-invasively inspect the contents of packages  116  for explosives, knives, guns, contraband, etc.  
         [0041]    Therefore, a CT detector array includes a plurality of CT detectors arranged to receive x-rays impinged by a subject and output electrical signals to a data acquisition system. Each of the CT detectors includes a plurality of detector elements. Each element includes a scintillator and a photodiode optically coupled to one another. The CT detectors are aligned relative to one another such that cross-talk throughout the plurality of CT detectors is substantially uniform.  
         [0042]    A CT system includes a rotatable gantry having a bore centrally disposed therein. A table is provided and is movable fore and aft through the bore and configured to position a subject for CT data acquisition. A high frequency electromagnetic energy projection source is positioned within the rotatable gantry and is configured to project high frequency electromagnetic energy toward the subject. A detector array is disposed within the rotatable gantry and configured to detect high frequency electromagnetic energy projected thereat and impinged by the subject. The detector array is configured to be absent substantial cross-talk discontinuities.  
         [0043]    A method of manufacturing a CT detector array includes the steps of forming a plurality of CT detectors in an array and connecting the plurality of detectors to one another such that cross-talk throughout the connected plurality of CT detectors is substantially uniform. Each CT detector includes a plurality of detector elements.  
         [0044]    The present invention has been described in terms of the preferred embodiment, and it is recognized that equivalents, alternatives, and modifications, aside from those expressly stated, are possible and within the scope of the appending claims.