Abstract:
A monolithic solid-state detector using a staggered arrangement of pixels in multiple rows improves spatial resolution without requiring reduction in pixel size. Parallelogram shapes of CZT monolith allow tiling in one dimension without inefficient zones between monoliths. A scanning device using linear array of detectors with non-rectangular shape and staggered rows of detection elements such that no dead zones occur within a scan field.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
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   STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT 
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   BACKGROUND OF THE INVENTION 
   The present invention relates generally to x-ray detectors, and in particular, to a cadmium zinc telluride (CZT) detector used for quantitative x-ray imaging. 
   Measurements of the x-ray absorption by an object at two different x-ray energies can reveal information about the composition of that object as decomposed into two selected basis materials. In the medical area, the selected basis materials are frequently bone and soft tissue. The ability to distinguish bone from surrounding soft tissue allows x-ray images to yield quantitative information about in vivo bone density for the diagnosis of osteoporosis and other bone disease. 
   Selecting different selected basis materials allows dual energy x-ray measurements to be used for other purposes. For example, dual energy x-ray measurements can be used for the analysis of body composition by distinguishing between fat and lean tissue, or for baggage scanning by distinguishing between explosive and non-explosive materials. 
   Cadmium zinc telluride (CZT) detectors may be used to measure x-rays passing through a measured object in dual energy x-ray systems. Such CZT detectors release an electrical charge for each incident photon proportional to the photon energy and thus allow separate measurement of high and low energy x-rays as sorted by pulse height. 
   Generally, a CZT detector employs a number of separate crystals of CZT, each having a front and rear surface electrode to detect x-rays within a pixel defined by the area of the crystal. Constructing a CZT detector requires the assembly of many separate CZT crystals which can be difficult. High-resolution detectors having smaller pixel sizes require smaller crystals, exacerbating the problem of assembly. 
   SUMMARY OF THE INVENTION 
   The present invention provides a high resolution CZT detector constructed of a monolithic crystal of CZT having multiple electrodes placed on one face to define multiple pixels. The monolithic design eliminates the assembly problems caused by the use of many separate small crystals. However, regions between pixels are known to be inefficient when counting x-rays absorbed between adjacent pixels (“gutter” regions) due to mutual sharing of deposited charge. Additionally in slot-scanning applications it is more efficient to cover a significant area by use of an extended linear array of monoliths. This necessarily implies that multiple crystals must be butted against each other end-to-end resulting in dead zones between crystals. For these reasons, the present invention has multiple rows of staggered pixels on each crystal. Scanning using staggered rows of pixels allows subsequent rows of detector elements to cover inefficient regions of previous rows. To enable tiling of multiple monolithic elements, without interruption of pixel pitch along rows or loss of efficiency due to gaps between monoliths, monoliths are fabricated into parallelogram shapes. 
   These particular features, objects and advantages may apply to only some embodiments falling within the claims and thus do not define the scope of the invention. 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
       FIG. 1  is a cross-sectional view of a monolithic CZT detector according to the present invention showing charge carrier migration from the gutter regions into adjacent pixel regions.  FIG. 1  also shows the detection circuitry used for initialization of the bias on the monolithic detector and for interpolation within the defined pixels; 
       FIG. 2  is a top planar view of the rear surface of a monolithic CZT detector showing the placement of the steering electrodes in a grid pattern and showing the location of the anodes in a staggered parallelogram configuration for improved sampling in a scanning x-ray machine; and 
       FIG. 3  is a figure similar to that of  FIG. 2  showing an alternative staggered configuration of electrodes using rectangular detector elements. 
   

   DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
   Referring now to  FIG. 1 , a solid-state, dual energy x-ray detector system  10  may include a monolithic CZT crystal  12  having a front surface  14  normally facing a source of x-ray photons  16  and  18  and a rear surface  20  on the opposite side of the CZT crystal from the front surface. Alternatively, other detector materials such as CdTe and HgI 2  may be used. 
   A cathode  22  is applied to the front surface  14  of the CZT crystal  12 , and an anode  24  is applied to the rear surface  20  of the CZT crystal  12  to provide a biasing electrical field between them. Generally, the cathode  22  will cover the entire front surface  14 , but the anode will cover only a small area centered on the rear surface  20 . Both the cathode  22  and anode  24  may be applied directly to the CZT crystal  12 , for example, by sputtering, and are preferably formed of a conductive metal such as platinum. The front surface  14  of the CZT crystal  12  may also be protected by a light-opaque, x-ray transparent material such as aluminized Mylar. 
   The anodes  24  are separated by a gutter region  25 . In one embodiment of the invention, the anodes  24  are approximately 1.5 by 2.5 millimeters in area and the gutter regions  25  are approximately 150–200 microns wide. The gutter regions  25  serve to electrically isolate the anodes  24  to permit independent measurement of bursts of charge released between the cathode  22  on front surface  14  and the anodes  24  on the rear surface  20  along axis  23  for each pixel region  15 . Weak electric fields in this inter-pixel (gutter) region are responsible for inefficient charge collection. Although the preferred embodiment may use steering electrodes (not shown), there is always a region (typically 0.1–0.2 mm) in which charge is split between two pixels, due to the finite width of charge deposition created by x-ray absorption. 
   To promote efficient collection of charge deposited in the crystal  12 , a bias voltage from bias voltage source  31  is applied across the opposed cathode  22  and anodes  24  of each pixel region  15  producing an electrical field  32 . X-ray photons  16  passing through cathode  22  on the front surface  14  enter the monolithic crystal  12  to liberate charge carriers  34  (shown here as electrons) which are then collected by anodes  24  on the rear surface  20  and conducted via separate leads  36  for each pixel region  15  to a ground referenced charge integrator  38 . The amount of charge liberated by each photon  16  is indicative of the energy of the x-ray photon  16 . Outputs from the charge integrators  38  are received by a processing computer  40  that may produce a quantitative image of the x-ray photons  16  according to techniques well known in the art. 
   In contrast to x-ray photons  16  striking within the pixel regions  15 , x-ray photons  18  passing into the monolithic crystal  12  at gutter region  25  will produce charge carriers  39  that may migrate into a pixel region  15  to be collected by anode  24  on the rear surface  20 . These charge carriers  39  degrade the quantitative accuracy and spatial resolution of a monolithically designed detector system  10 , adding an effective noise component to the charge collected from x-ray photons  16 . 
   Referring now also to  FIG. 2 , generally the x-ray detector system  10  may provide for multiple detector elements on a single CZT crystal  12 . In this case, multiple anodes  24  will be placed on the CZT crystal  12 , each surrounded by steering electrodes  30 , may be interconnected and covered by a single cathode  22 . 
   The steering electrodes  30  surrounding each anode  24  describe by their perimeter a pixel region  15 . The pixel regions  15  describe areas which may independently detect x-ray photons  16  to produce a quantitative detection value that will be mapped to individual pixels in a resultant image. 
   In the embodiment shown in  FIG. 2 , the pixel regions  15  are generally parallelograms tiling in rows and slanted columns. In this embodiment, each parallelogram pixel region  15  has a first base  52  generally perpendicular to a scan direction  54  in which the x-ray detector system  10  will be scanned to collect information over an area of the patient. Sidewalls  56  of the parallelogram and the pixel regions  15  are angled such that the centers of the pixel regions  15  defined approximately by the center of the anode  24  for a first row of pixel regions  15 , follow paths  60  that interleave with paths  62  followed by centers of the pixel regions  15  of a second row of pixel regions  15 . In this way, larger pixel regions  15  may provide higher spatial resolution sampling to improve the resultant image. Further the data lost from the gutter areas in one row are regained in the next staggered row. 
   Referring now to  FIG. 3 , in an alternative embodiment, the pixel regions  15  may be rectangular with the pixel regions  15  of a first row staggered with respect to the second row to provide interleaved paths  60  and  62  as before. The rectangular pixel regions  15  of  FIG. 2  provide the advantage of a more compact detection region limiting the effective size of a convolution kernel (a function of the project width of the pixel regions  15  on a line perpendicular to the scan direction  54 ) that can make a resultant image less distinct. 
   Referring still to  FIG. 3 , a convenient form factor for the x-ray detector system  10  has two rows each having eight pixel regions  15 . Multiple detector systems  10  of this or similar form factors may be ganged edgewise to provide arbitrary continuations of the rows. For an x-ray detector system  10  having rectangular pixel regions  15 , pixel regions  15   a  and  15   b  at a first and second row of a right edge of the x-ray detector system  10  may be cut at an angle with respect to the scan direction  54  to equally reduce the area of the pixel regions  15   a  and  15   b . Similarly reduced pixel regions  15   c  and  15   d  at a first and second row of a left edge of a next x-ray detector system  10 ′ may be placed in close proximity to their counterpart pixel regions  15   b  and  15   a . The area of each pixel region  15   a–d  is reduced by half the width of the joint gap between x-ray detector system  10  and  10 ′, which then preserves the regular lateral of the other pixel regions  15 . In another embodiment, the area of each pixel region  15   a – 15   d  is reduced to slightly less than half to accommodate the joint gap between x-ray detector system  10  and  10 ′. This provides two virtual pixel regions, the first being a combination of the signals from pixel regions  15   a  and  15   d , and the second being a combination of the pixel regions  15   b  and  15   c . The slightly reduced detection area of these detectors&#39; virtual pixel regions may be corrected mathematically by a weighting factor applied by the computer receiving the signals. 
   The present invention is applicable not only to polygonal electrode regions, but other shapes as well. 
   It is specifically intended that the present invention not be limited to the embodiments and illustrations contained herein, but include modified forms of those embodiments including portions of the embodiments and combinations of elements of different embodiments as come within the scope of the following claims.