Abstract:
The present invention is directed to a scintillator array having an integrated air gap. By integrating an air gap within the reflector, light collection efficiency is improved while simultaneously lowering cross-talk between scintillators. That is, implementing a reflector without chromium oxide (Cr 2 O 3 ) increases light reflectivity and an air gap lowers cross talk through the reflector. To further improve the reflectivity, the base reflector material may be coated with a low index material and a reflective material such as silver.

Description:
BACKGROUND OF INVENTION  
         [0001]    The present invention relates generally to diagnostic imaging systems and, more particularly, to a reflector for a scintillator array having an integrated air gap. Specifically, the scintillator array is constructed such that a uniform air gap or void exists between adjacent scintillators.  
           [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. 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.  
           [0003]    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.  
           [0004]    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.  
           [0005]    Scintillator arrays typically incorporate a reflector layer between adjacent scintillators to limit cross-talk between the scintillators thereby improving light collection efficiency of the corresponding photodiodes. Generally, the reflector is formed of a material comprising chromium oxide or other types of optically absorbent material. Because chromium oxide operates as a good absorbent of light, the relative reflectivity of the reflector is reduced. As such, incorporating a reflector layer that includes chromium oxide, a trade-off in CT detector design is made between lower cross talk and reflectivity. If the reflector layer is fabricated without chromium oxide or other optically absorbent materials, cross talk between scintillators increases. In contrast, implementing optically absorbent materials reduces cross talk but lowers the reflectivity of the reflector. Reduced reflectivity degrades low signal performance and increased cross talk affects spatial resolution. Low signal performance is a function of noise generated in the CT detector. As reflectivity falls the light output of the scintillator also falls. Noise, however, is relatively constant, therefore, decreases in light output increases the ratio of noise to functional light output. Additionally, known CT detectors are constructed such that the reflector material is disposed such that it fills any spaces that exist between adjacent scintillators. This also contributes to increased cross-talk between scintillators as there is a constant interface between the scintillators.  
           [0006]    It would therefore be desirable to design a CT detector having a reflector with integrated air gaps to improve light collection efficiency and lower cross-talk between scintillators.  
         BRIEF DESCRIPTION OF INVENTION  
         [0007]    The present invention is directed to a scintillator array having an integrated air gap overcoming the aforementioned drawbacks. By integrating an air gap within the reflector, light collection efficiency is improved while simultaneously lowering cross-talk between scintillators. That is, implementing a reflector without chromium oxide (CR 2 O 3 ) increases light output and an air gap reduces cross talk through the reflector. To further improve the reflectivity, the reflector material may be coated with a low index material and a reflective material such as silver.  
           [0008]    Therefore, in accordance with one aspect of the present invention, a CT detector includes a scintillator array of scintillators arranged to receive x-rays from an x-ray projector source and output light in proportion to the x-rays received. The detector further includes a cast reflector integrally disposed between adjacent scintillators. An air gap is disposed within the cast reflector such that a space or void is formed between adjacent scintillators.  
           [0009]    In accordance with another aspect of the present invention, a CT detector having a scintillator array optically coupled to a photodiode array is provided. The CT detector is formed by creating voids between adjacent scintillators of the scintillator array and disposing a cast reflector within each of the voids. Air gaps are then created in the cast reflector disposed within the voids. A photodiode array is then coupled to the scintillator array to form a CT detector.  
           [0010]    According to another aspect of the present invention, a CT system includes a rotatable gantry having a bore centrally disposed therein and a table movable fore and aft through the bore and configured to position a subject for CT data acquisition. The CT system further includes a high frequency electromagnetic energy projection source positioned within the rotatable gantry and configured to project high frequency electromagnetic energy toward the subject. A detector array is provided and disposed within the rotatable gantry and configured to detect high frequency electromagnetic energy projected by the projection source and impinged by the subject. The detector array includes a scintillator array having a plurality of scintillators and wherein the scintillator array includes a reflector between adjacent scintillators. The reflector is configured to have an integrated air gap. The detector array further includes a photodiode array coupled to the scintillator array and configured to produce electrical signals in response to light emitted by the scintillator array.  
           [0011]    In accordance with yet a further aspect of the present invention, a scintillator array includes a plurality of scintillators arranged to receive x-rays and generate light in response thereto. The scintillator array further includes a reflector disposed between adjacent scintillators and a plurality of voids integrated in the reflector along at least one dimension. The reflector is formed of a material absent chromium oxide.  
           [0012]    In accordance with another aspect of the present invention, a method of manufacturing a scintillator array includes the steps of creating voids between adjacent scintillators. The method further includes disposing a cast reflector within each of the voids and creating air gaps in the cast reflector.  
           [0013]    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  
       [0014]    The drawings illustrate one preferred embodiment presently contemplated for carrying out the invention.  
         [0015]    In the drawings:  
         [0016]    [0016]FIG. 1 is a pictorial view of a CT imaging system.  
         [0017]    [0017]FIG. 2 is a block schematic diagram of the system illustrated in FIG. 1.  
         [0018]    [0018]FIG. 3 is a perspective view of one embodiment of a CT system detector array.  
         [0019]    [0019]FIG. 4 is a perspective view of one embodiment of a CT detector.  
         [0020]    [0020]FIG. 5 is illustrative of various configurations of the detector in FIG. 4 in a four-slice mode.  
         [0021]    [0021]FIG. 6 is a cross-sectional schematic of a CT detector in accordance with one embodiment of the present invention.  
         [0022]    [0022]FIG. 7 is a cross-sectional schematic of a CT detector in accordance with another embodiment of the present invention. 
     
    
     DETAILED DESCRIPTION  
       [0023]    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.  
         [0024]    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 .  
         [0025]    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 .  
         [0026]    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 .  
         [0027]    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 .  
         [0028]    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 .  
         [0029]    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 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 .  
         [0030]    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.  
         [0031]    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.  
         [0032]    Referring now to FIG. 6, a cross-sectional schematic of a CT detector  20  in accordance with one embodiment of the present invention is shown. CT detector  20 , as described previously, includes a scintillator array  56  comprising a plurality of scintillators or scintillation elements  57 . Coupled to the scintillator array  56  is a photodiode array  52 . Coupling photodiode array  52  to scintillator array  56  is an optical coupler  86  that typically is in the form of an optical epoxy.  
         [0033]    Still referring to FIG. 6, a cast reflector  88  is typically used to coat each of the scintillators  57 . Preferably, the cast reflector is formed from an epoxy loaded with titanium dioxide (TiO 2 ). The cast reflector is generally opaque and is designed to prevent light emissions from each of the scintillators. That is, the cast reflector operates to confine the light generated by each of the scintillators to be within the respective scintillators. As such, light is not translated between adjacent scintillators. Since the photodiode array  52  is designed to detect light emissions from each of the scintillators  57 , the cast reflector is used to improve the convergence of light toward the photodiode array  52 .  
         [0034]    The cast reflector is also designed to absorb light emissions from the scintillators to assist with preventing cross-talk between the scintillators. As such, the reflector is preferably fabricated without chromium oxide and other absorbing materials used to improve the cross-talk characteristics of the reflector. As illustrated in FIG. 6, detector  20  is further constructed such that an air gap or void  90  is formed between adjacent scintillators  57 . Constructing the detector  20  in such a manner  50  as to incorporate an air gap  90  between adjacent scintillators  57  improves overall reflectivity of the reflector. As shown in FIG. 6, the air gap/reflector combination results in a U-shaped channel being formed between each of the scintillators  57 . The air gap  90  is constructed such that it does not extend to the optical coupler  86 . That is, the reflector  88  forms a cast bridge  92  between adjacent scintillators  57 . As such, the cast reflector  88  is positioned adjacent to the optical coupler  86 . However, air gap  90  may be formed so as to extend to the optical coupler.  
         [0035]    Referring now to FIG. 7, a CT detector  20   a  constructed in accordance with another embodiment of the present invention is shown. Similar to the detector of FIG. 6, detector  20   a  includes a plurality of scintillators  57   a  arranged in an array  56   a  that is coupled to a photodiode array  52   a  with an optical coupler  86   a . A cast reflector  88   a  is used to coat each of the scintillators and is designed to absorb cross-talk emissions between scintillators  57   a . In contrast to the detector of FIG. 6, detector  20   a  incorporates a thin layer of low index material  94  coating reflector  88   a . A reflective layer  96   a  is then affixed to the low index material layer  94 . The low index layer  94  and the reflective material layer  96  operate to improve the reflectivity and reduce cross talk between scintillators  57   a . Silver is one material that is well suited for implementation as the reflective layer  96 . Detector  20   a  also is constructed with an air gap  90   a  being disposed between each of the scintillators. Air gap  90   a  improves the reflectivity of light between scintillators  57   a.    
         [0036]    A number of fabrication techniques may be used to construct a CT detector similar to that shown in FIGS. 6 and 7. One such method includes the casting of titanium dioxide laden epoxy between scintillators. Depending upon the particular detector, the reflector may be disposed along one or two dimensions. That is, the epoxy may be cast such that a reflector is fabricated along an x-axis, a z-axis, or both. Once the cast reflector is deposited between the scintillators, the cast reflector is diced to created air gaps. These air gaps or spaces may be made with any number of cutting techniques. In one embodiment, the reflector is left with the dissected air gaps to form the CT detector of FIG. 6. Alternately, the air gaps may be subsequently coated with a low index material or resin followed by a coat or film of reflective material to form the CT detector of FIG. 7. Both of these methodologies produce a reflector having improved light collection efficiency by the photodiode and lower cross-talk. As stated above, it is preferred that the reflector be fabricated from a material absent chromium oxide to improve cross-talk absorption.  
         [0037]    Referring now to FIG. 8, 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  12  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.  
         [0038]    Therefore, in accordance with one embodiment of the present invention, a CT detector includes a scintillator array of scintillators arranged to receive x-rays from an x-ray projector source and output light in proportion to the x-rays received. The detector further includes a cast reflector integrally disposed between adjacent scintillators. An air gap is disposed within the cast reflector such that a space or void is formed between adjacent scintillators.  
         [0039]    In accordance with another embodiment of the present invention, a CT detector having a scintillator array optically coupled to a photodiode array is provided. The CT detector is formed by creating voids between adjacent scintillators of the scintillator array and disposing a cast reflector within each of the voids. Air gaps are then created in the cast reflector disposed within the voids. A photodiode array is then coupled to the scintillator array to form a CT detector.  
         [0040]    According to another embodiment of the present invention, a CT system includes a rotatable gantry having a bore centrally disposed therein and a table movable fore and aft through the bore and configured to position a subject for CT data acquisition. The CT system further includes a high frequency electromagnetic energy projection source positioned within the rotatable gantry and configured to project high frequency electromagnetic energy toward the subject. A detector array is provided and disposed within the rotatable gantry and configured to detect high frequency electromagnetic energy projected by the projection source and impinged by the subject. The detector array includes a scintillator array having a plurality of scintillators and wherein the scintillator array includes a reflector between adjacent scintillators. The reflector is configured to have an integrated air gap. The detector array further includes a photodiode array coupled to the scintillator array and configured to produce electrical signals in response to light emitted by the scintillator array.  
         [0041]    In accordance with yet a further embodiment of the present invention, a scintillator array includes a plurality of scintillators arranged to receive x-rays and generate light in response thereto. The scintillator array further includes a reflector disposed between adjacent scintillators and a plurality of voids integrated in the reflector along at least one dimension. The reflector is formed of a material absent chromium oxide.  
         [0042]    In accordance with another embodiment of the present invention, a method of manufacturing a scintillator array includes the steps of creating voids between adjacent scintillators. The method further includes disposing a cast reflector within each of the voids and creating air gaps in the cast reflector.  
         [0043]    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.