WO2013012809A1

WO2013012809A1 - Radiation detector modules based on multi-layer cross strip semiconductor detectors

Info

Publication number: WO2013012809A1
Application number: PCT/US2012/046936
Authority: WO
Inventors: Yonggang Cui; Ralph James; Anwar Hossain
Original assignee: Brookhaven Science Associates, Llc
Priority date: 2011-07-15
Filing date: 2012-07-16
Publication date: 2013-01-24

Abstract

A multi-layer radiation detector is described based on cross-strip detector design for imaging tissue or an inanimate object that is configured to provide a more compact design, reduce the number of readout channels, reduce the inactive volume between each thin detector, and provide better interconnect density and thermal management than conventional multi-pixel and cross-strip detectors.

Description

RADIATION DETECTOR MODULES BASED ON MULTI-LAYER CROSS STRIP SEMICONDUCTOR DETECTORS

CROSS-REFERENCE TO A RELATED APPLICATION

[0001] This application claims priority to U.S. Provisional Application Serial No.

61/508,1 13, filed on July 15, 201 1 , and entitled "Radiation Detector Modules Based on Multi- Layer Cross Strip Semiconductor Detectors," the content of which is incorporated herein in its entirety.

STATEMENT OF GOVERNMENT LICENSE RIGHTS

[0002] The present invention was made with government support under contract number

DE-AC02-98CH10886 awarded by the U.S. Department of Energy. The United States government may have certain rights in this invention.

BACKGROUND

I. FIELD OF THE INVENTION

[0002] The present invention relates to the field of radiation imaging. In particular, the present invention relates to a multi-layer stacked semiconductor-based apparatus for the detection of high-energy electromagnetic radiation.

II. BACKGROUND OF THE RELATED ART

[0003] Semiconductor nuclear radiation detectors have experienced a rapid development in the last decade. They are now used in a large variety of fields, including nuclear physics, X- ray and gamma ray astronomy, and nuclear medicine. Their good energy- and spatial- resolution, and their capability to be integrated into compact systems are very attractive features, in comparison with other types of detectors, such as gas detectors and scintillator detectors, for example, ordinarily used in positron emission tomography (PET) systems.

[0004] In recent years, a substantial effort has been invested in developing a range of compound semiconductors with wide band-gap and high atomic number for X-ray and gamma ray detection. These compound semiconductors are generally derived from elements of groups III and V, e.g. GaAs, and groups II and VI, e.g. CdTe, of the periodic table. Besides binary compounds, ternary materials have been also produced, e.g. CdZnTe and CdMnTe. Among these compound semiconductors, cadmium telluride (CdTe) and cadmium zinc telluride

(CdZnTe or CZT) are two of the most promising materials for radiation detectors with good energy resolution, high detection efficiency, and room-temperature operation. Moreover, compared to scintillator-based detectors, the semiconductor-based detectors are easier to fabricate.

[0005] One of the drawbacks of the semiconductor based detectors, however, is their low charge-carrier mobility that causes a slow signal response and low timing resolution, especially in thick semiconductor detectors. This is the main reason that the semiconductor detectors, e.g. , CZT detectors, have not yet been adopted for coincident-measurement applications, for example, in PET systems where timing resolution is critical to the system performance.

[0006] A variety of approaches have been explored to address the shortcomings of the semiconductor detectors. Most efforts that focus on thick detectors attempt to extract timing information from cathode signals and simultaneously use the amplitude ratio of the anode to cathode to obtain the depth of interaction of incident gamma-ray in the detector and correct the charge loss. However, the cathode signals can be very weak, especially when the interaction points are close to the anode. Furthermore, the cathode is an unsuitable choice for timing extraction due to the lower mobility of holes. Another method to obtain timing from the cathode is via waveform digitization. Again, this is unsuitable because the signal-processing circuit can be very complicated and the on-line reconstruction of coincident events can be cumbersome and problematic.

[0007] Another approach is to use thin semiconductor detectors with anode and cathode developed on the two opposite large-area surfaces. By applying high voltage biasing on the two electrodes, the charge carriers (holes and electrons) generated by incident X-ray or gamma-ray photons can be swept quickly to the cathode and the anode. The induced signals are then collected by the attached readout electronics circuits. Because the detector is thin, such signal collecting process takes much shorter time and allows better timing resolution. The cathode and anode in this approach are normally developed into specific patterns to obtain position resolution along the electrode plane. For example, the patterns include a strip array or pixel array.

[0008] The thin semiconductor detectors, however, exhibit relatively low detection efficiency. To increase the detection volume, multiple detectors can be stacked on top of each other (with a spacer if needed) and signals from each detector are read out separately. A drawback of this arrangement is that the total number of readout channel is prohibitively large. Specifically, in order to readout multiple detector layers in the detector module, the system requires at least 2*N layer of readout channels to readout an N-layer detector module. This creates additional challenges in connection with data acquisition bandwidth, interconnect density and thermal management. Hence, it is highly desirable to develop a multi-layer stacked semiconductor detector with reduced number of readout channels that can be used in an ultrahigh resolution PET detection system(s). SUMMARY

[0009] A multi-layer semiconductor detector is disclosed comprising two or more superimposed semiconductor detectors arranged to provide a more compact design, reduce the number of readout channels, reduce the inactive volume between each thin detector, and provide better interconnect density and thermal management.

[0010] The multi-layer semiconductor radiation detector includes a plurality of thin semiconductor detectors superimposed on top of each other. In one embodiment, the number of semiconductor detectors is between 2 and 100. In a more preferred embodiment the number of semiconductor detectors is between 2 and 16. Each semiconductor detector has a semiconductor wafer commonly referred to as a "slab" that can have a length and a width between 1 mm and 100 mm and a height (thickness) between 0.1 mm and 20 mm. In one embodiment, all semiconductor slabs in the multi-layer semiconductor radiation detector have the same thickness. In another embodiment, the semiconductor slabs may have different thicknesses.

[0011] The multi-layer semiconductor radiation detector further includes a plurality of electrode layers having a cross-strip pattern. The cross-strip pattern has one set of linear array of electrodes on one side of the semiconductor slab, and another set of linear array of electrodes on the opposing side of the semiconductor slab, oriented perpendicular to the direction of the opposing array (see FIG. 1A). In this configuration, signals are read out from both sides, and the coincidence of signals from these two sides indicates the interaction position of the gamma-ray photon inside the detector, thus providing desired spatial information while using fewer electronic channels compared to standard square pixel array designs, e.g., 2n vs. n².

[0012] Each electrode layer of the multi-layer semiconductor radiation detector makes conductive contacts with either side of each semiconductor slab. The contacts can be fabricated using electroless plating, sputtering, evaporating or surface alloying. Alternatively, the contacts can be Ohmic connections or Schottky connections. As applicable to all present embodiments, the adjacent semiconductor slabs share a common electrode layer between the slabs, thus reducing the number of the required electrode layers and the total readout channels by 50%. The multi-layer semiconductor radiation detector maintains good timing performance and spatial resolution. The position resolution along the z direction can be determined by the thickness of the semiconductor slabs, while the position resolution in x-y plane can be determined by the pitch of the cross strip pattern.

[0013] The semiconductor slab of the multi-layer semiconductor radiation detector is preferably constructed from a solid-state semiconductor capable of operating as a photon-to- charge direct conversion device. In particular, such semiconductors may be derived from, but not limited to, elements of groups III and V (e.g. GaAs), groups II and VI (e.g. CdTe), and group IV (e.g. Si) of the periodic table. Among these semiconductors and their alloys, cadmium zinc telluride (CdZnTe) and silicon (Si) are the most preferred.

[0014] Each electrode layer of the multi-layer semiconductor radiation detector has a cross-strip pattern of electrode strips that make conductive contacts with either side of each semiconductor slab. In one embodiment, the number of electrode strips in each electrode layer can range between 1 and 1000, with the number of electrode strips between 2 and 500 being preferred and between 2 and 100 being more preferred. In one embodiment, the number of electrode strips is inversely proportional to the width of each strip. For example, if the electrode has n number of strips with width m, to accommodate 2n strips, the width of each strip would have to be reduced by at least m 2. While the number of electrode strips in the electrode layer may be the same or similar to the number of electrode strips in the electrode layer on the opposite side of the slab, the number of strips can also be different. Similarly, the width, length and pitch of the strips in each electrode layer typically would be determined by the specific applications for the detector, such as the desired spatial resolution in a PET nuclear-medical imaging application. For instance, due to a small-pixel effect, which mitigates the cathode-signal deficiency due to the slow-moving and easily trapped holes, the number of electrode strips in the electrode layer that function as an anode may be significantly greater than the number of electrode strips in the electrode layer that function as a cathode. For example, the cathode layer may have 4 to 10 cathode strips, while the anode layer may have 3 to 10 times as many, e.g., 30 to 50 anode strips. However, in some embodiments where the spatial resolution on x- and/or y direction is not needed, the electrode strips on either the cathode layer or the anode layer can merge together to form a single planar electrode.

[0015] The conductive connections from detector electrode(s) to inputs of the readout circuits can be provided by wire bonding, which routes individual wires from each electrode to the inputs of readout circuits; by flexible cable, which has all the conductor wires pre-routed and embedded on a soft substrate; and, by directed connections to a printed circuit board (PCB) where the readout circuit resides. The physical connection between the detector electrode and the signal carrier (wire, flexible cable or PCB board) can be established by a conductive epoxy, an indium-bump-bond process, or a soldering process.

[0016] The cathode or the anode layer can be made from gold, platinum, indium, indium- tin-oxide, rhodium and other conductive metal and non-metallic materials. Those skilled in the art will recognize that nearly any conductor may be used for the electrodes. In one embodiment, the cathode and anode electrodes can use the same metal, e.g. , gold, to make ohmic connections. In another embodiment, the cathode and anode electrodes use different electrodes, e.g., platinum and gold, to make Schottky connections in order to reduce the leakage current.

[0017] The multi-layer semiconductor detector(s) described herein may be employed in a high-resolution PET system for radiation imaging of an object or tissue of interest. A plurality of multi-layer semiconductor detectors is arranged in a z direction facing the axis of a gantry. Each pair of multi-layer detectors are placed facing each other along the circumference of the gantry at a predefined radial distance. Depending on the application, the individual multi-layer detectors can have different orientations along the circumference of the PET system. For example, the detector modules can be arranged so the x- (or y-) direction faces the axis of the gantry.

[0018] In addition to this traditional ring-like imaging application, the multilayer cross strip detector modules can be used in Compton gamma cameras. In this application, a gamma ray interacts with one or more detector modules for two or more times. Each interaction position, timing and the deposited energy are measured by signals read out from cathode and anode surrounding the interaction point. All the information can be used by software algorithm to reconstruct the original direction of the gamma ray.

[0019] In addition to imaging applications, multi-layer cross strip detector modules can also be used in spectroscopic applications to measure the energy spectra of incident photons.

[0020] A method for radiation imaging of an object or tissue of interest is also provided using the multi-layer semiconductor detector(s) described herein. The method includes locating an object of interest, for instance, a cancerous tissue; positioning a multi-layer radiation imaging detector near the object of interest; and detecting the x-ray or gamma ray radiation emitted by the absorbed tracers within the object utilizing the multi-layer radiation detector. In particular, with reference to the PET system, the step of positioning can encompass positioning a plurality of multi-layer radiation imaging detectors in a circumference surrounding the object of interest at a predetermined distance, and the step of detecting can include absorbing radiation emitted by the radiation tracers in x-, y- and z- directions, storing/recording a plurality of images provided by each multi-layer radiation imaging detector, and reconstructing a 2D or 3D image of the object of interest.

[0021] Numerous medical-, industrial-, scientific-, environmental cleanup-, and national security-applications exist for the present multi-layer detector(s). Some of the most prominent are imaging systems for detecting and localizing tumors and other abnormalities in the body, hand-held instruments to detect the trafficking of nuclear materials, and portable field instruments for environmental monitoring and remediation.

BRIEF DESCRIPTION OF THE DRAWINGS

[0022] FIG. 1A illustrates a cross-strip detector.

[0023] FIG. IB illustrates a multi-layer detector module with stacked cross-strip detectors with adjacent detectors sharing either an anode or a cathode strips between them.

[0024] FIGs. 2A and 2B illustrate a spatial resolution determination of incident position in a multi-layer detector along z direction (A) or in the x-y plane for the detector (B) shown in FIG. IB.

[0025] FIG. 3A and 3B illustrate exemplary approaches of connecting two adjacent cross strip detectors by sharing their electrodes (A) using a conductive epoxy or using a flexible cable (B).

[0026] FIG. 4 illustrates an exemplary setup of a signal fan-out from multi-layer detector module to preamplifiers based on direct PCB trace connections, flexible cable connections, and/or wire-bonded connections. [0027] FIG. 5 illustrates multi-layer detector modules in PET imaging.

DETAILED DESCRIPTION

[0028] The present multi-layer semiconductor radiation detectors provide a more compact design, reduces the number of readout channels, reduces the inactive volume, and provide better interconnect density and thermal management compared to conventional cross- strip detectors.

[0029] As illustrated in FIGs. 1A and IB, the present multi-layer semiconductor radiation detector 101 includes a plurality of detectors 100 superimposed on top of each other. In one embodiment, the number of detectors 100 is between 2 and 100, while having between 2 and 16 detectors 100 is more preferred. Each detector 100 has a thin two-dimensional semiconductor slab 108. Each semiconductor slab 108 can have the length and width between 1 mm and 100 mm and the thickness (height) between 0.1 mm and 20 mm. In one embodiment, the thicknesses of the semiconductor slabs 108 are the same in all detectors 100 in the multilayer detector 101. In another embodiment, the thicknesses of the semiconductor slabs 108 are different between detectors 100 in the multi-layer detector 101. In yet another embodiment, the thicknesses of the semiconductor slabs 108 are the same in some detectors 100 and different in other detectors 100 in the multi-layer detector 101.

[0030] The semiconductor slab of the multi-layer semiconductor radiation detector described herein is generally derived from, but not limited to, elements of groups III and V (e.g. GaAs), groups II and VI (e.g. CdTe), and group IV (e.g. Si) of the periodic table. Instead of binary compounds or single element semiconductors (e.g. Si), ternary materials also may be used as the compound semiconductors capable of operating as photon-to-charge direct conversion devices, e.g., Cdi__xZn_xTe, Cdi__xMn_xTe, Hg_xBri__xl2, and Hg_xCdi__xl2 where 0<x<l . It is common practice to omit the fractional subscripts when referring to the alloy families; such practice is followed in describing the present multi-layer semiconductor radiation detector(s). Among these semiconductors and their alloys, a specific examples of the semiconductors that can be used in the present multi-layer semiconductor radiation detector include cadmium telluride (CdTe), cadmium zinc telluride (CdZnTe), cadmium manganese telluride (CdMnTe), zinc telluride (ZnTe), silver gallium telluride (AgGaTe₂), mercury cadmium telluride (HgCdTe), gallium telluride (Ga₂Te₃), gallium arsenide (GaAs), zinc selenide (ZnSe), lead(II) oxide (PbO), silicon carbide (SiC), gallium selenide (Ga₂Se₃), mercury zinc selenide (HgZnSe), thallium arsenic selenide (Ti₃As₂Se₃), thallium gallium selenide (TlGaSe₂), indium phosphide (InP), bismuth tri- iodide (B¾), indium iodide (Inl₂), lead(II) iodide (Pbl₂), thallium lead iodide (TlPbl3), thallium mercury iodide (TLtHgle), mercury bromine iodide (HgBrI₂), mercury cadmium iodide (HgCdI₂), thallium bromide (TIBr), silicon crystal (Si) and mercuric iodide (Hgl₂). However, it will be appreciated and understood by those skilled in the art that any compound or element may be used in the present multi-layer semiconductor radiation detector(s) as long as it is capable of operating as photon-to-charge direct conversion device. In a particularly preferred embodiment, the compound semiconductor crystal used for the plurality of radiation detector elements is made from cadmium zinc telluride (CdZnTe or CZT) crystals. One skilled in the art will appreciate that the semiconductor may be larger or smaller and vary in shape depending upon the design specifications.

[0031] As illustrated in FIGs. 2 A and 2B, each detector 100 further has two electrode layers (110, 120) positioned on each side of the slab 108 having a cross-strip pattern. The cross- strip pattern has one set of linear array of electrodes (111 or 121) on one side of the

semiconductor slab 108, and another set of linear array of electrodes (111 or 121) on the other side of the semiconductor slab 108, orienting perpendicular to the direction of the former array. In this embodiment, for example, the electrode 110 is an anode, and the electrode 120 is a cathode. Each electrode (110, 120) can be made from gold, platinum, indium, indium-tin-oxide, rhodium and other conductive metal materials. Those skilled in the art will recognize that nearly any conductor may be used for the electrodes and not limited to the specific materials specified above.

[0032] An example of a double-sided orthogonal strip detector that may be used in the present multi-layer semiconductor radiation detector is described in J.C. Lund et al. in

"Miniature Gamma-Ray Camera for Tumor Localization", published by Sandia National Laboratories (SAND97-8278, August 1997), which is incorporated by reference in this application in its entirety. For example, as illustrated in FIG. 2A, the orientation of the electrode strips 111 in the electrode layer 110 is along y-axis, while the orientation of the electrode strips 121 in the electrode layer 120 is along x-axis. In this configuration, signals are read out from both sides, and the coincidence of signals indicates the interaction position of the gamma-ray photon inside the detector, thus providing desired spatial information while using fewer electronic channels compared to standard square pixel designs, .e.g. , 2n vs. n . In particular, in the orthogonal strip design, rows and columns of parallel electrical contacts (strips) are placed at right angles to each other on opposite sides of a piece of semiconductor slab (wafer). Radiation detection on the detector plane is determined by scoring a coincidence event between a column and a row. Readout electronics transmit the received signals to processing and analyzing equipment in a known manner. Each electrode layer (110 or 120) of the detector 100 has contacts on the semiconductor slab 108. These contacts can either be fabricated by electroless plating, evaporation, surface alloying, sputtering, or other known means to deposit thin layers of materials suitable for forming electrical contacts. Alternatively, the contacts can be Ohmic or Schottky connections.

[0033] The number of electrode strips (111, 121) in each electrode layer (110, 120) can range between 1 and 1000. In a more preferred embodiment, the number of electrode strips (111, 121) is between 2 and 500, while in an even more preferred embodiment, the number of electrode strips (111, 121) is between 2 and 100. While all electrode strips can have the same dimensions, i.e., width, length and pitch, it is also envisioned to have some electrode strips to have the same dimensions and some have different dimensions. It is also envisioned to have all strips being different depending on the overall design of the multi-layer detector 101. Those skilled in the art could determine the width, length and pitch of the strips in each electrode layer based on the applications of the detector.

[0034] While the number of electrode strips (111, 121) in the electrode layer (110, 120) that functions as anodes or as cathodes may be the same or similar, the number of strips can also be different. For instance, due to a small-pixel effect, which mitigates cathode-signal deficiency due to slow-moving and easily trapped holes, the number of electrode strips in the electrode layer that functions as anodes may be significantly more than the number of electrode strips in the electrode layer that function as a cathodes. For example, the cathode layer may have 4 to 10 cathode strips, while the anode layer may have 30 to 50 anode strips. However, in some embodiments where the spatial resolution on x- or y- direction is not needed, the electrode strips on the corresponding surfaces can merge together to form a single planar electrode.

[0035] As illustrated in FIG. IB, the adjacent semiconductor detectors 100 in the multi-layer detector 101 share an electrode layer(s) (either 110 or 120) interposed between semiconductor slabs 108 of each detector 100, thus reducing the total readout channels by 50%. Thus, the present multi-layer semiconductor radiation detector essentially has a shared electrode structure. The multi-layer semiconductor radiation detector 101 maintains good timing performance and spatial resolution. The position resolution along the z direction can be determined by the thickness of the semiconductor slabs. The event position between two adjacent layers is solved by the coincident signal readout from both anode and cathode channels.

[0036] For example, FIG. 2 A shows that if an incident photon deposits its energy in an area covered by anode 110 and cathode 120(k) inside detector 100(k), it will generate signals on anode output 112 and cathode output 122(k). If the incident event happens in an area covered by anode 110 and cathode 120(k+l) inside detector 100(k+l), it will generate signals on anode output 112 and a cathode output 122(k+l). On the other hand, the position resolution in x-y plane can be determined by the pitch of the cross-strip pattern. For example, as illustrated in FIG. 2B for an ease of understanding only one detector 100 is shown of the multi-layer detector 101. In this example, each electrode (110, 120) has seven (7) electrode strips (111, 121) with an anode output 112 and a cathode output 122 that generates a pair of coincident signals. Based on the coincident signal, it is possible to determine whether the incident event is within the cube (voxel) 102 defined by the overlap volume of cross strips 111 and 121.

[0037] In one embodiment, the number of the interposed electrode layers (either 110 or

120) is one between the adjacent semiconductor detectors 100 in the multi-layer detector 101. The signal is fanned-out from the electrode layer (110 or 120) to preamplifiers (113, 123) through direct connections using a wire or traces on a flexible cable or printed circuit board (PCB).

[0038] In an alternative embodiment illustrated in FIGs. 3A and 3B, each detector 100 is a double-sided cross-strip detector having two electrode layers (110 and 120). FIGs 3A and 3B show two adjacent detectors 100(k) and 100(k+l). The detector 100(k) has electrode strips on both sides of the semiconductor slab, oriented perpendicular to each other. Electrode strips lll(k) are aligned along the y-directions, and the electrode strips 121(k) are aligned along x- direction. The adjacent detector 100(k+l) has the same setup as 100(k) except the detector is rotated by 180° to align the interposed electrode strips 121(k) and 121(k+l) on the same side in the same direction, along x- and y- axes. With reference to FIG. 3A, the electrode strips 121(k) and 121(k+l) can be connected together by any known approaches such as gluing them together by conductive epoxy 131 or similar adhesives. In this embodiment, the epoxy 131 can be connected to a wire 132 for signal fan-out. With reference to FIG. 3B, if the signal readout uses flexible cable 134, a thin flexible cable 134 is inserted between the two adjacent electrode strips 121(k) and 121(k+l) and attached between the electrode strips by, for example, a conductive epoxy 131 to the contacts 133 on the flexible cable 134. As illustrated in FIG. 4, the signal fan- out from the electrode layers (110 or 120, i.e., anodes and cathodes) to the preamplifiers 113, 123 can use direct connections with the traces 136 on a printed circuit board (PCB) 130, use wire -bonding 132 or use flexible cable(s) 134 with traces 135. In one embodiment, at least one detector 100 is affixed to a mounting frame 130 or to the preamplifier 113, 123., while the other stacked detectors 100 in the multi-layer detector 101 are connected to the same or to a different board 130 via wire bonding 132 or flexible cables 134. When a flexible cable is used, the preamplifier 113, 123 can reside on the flexible cable or the PCB board that the cable is connected to.

[0039] In one embodiment, the detector further has a collimator that is adapted to be positioned substantially parallel to the stack of multi-layer detectors. The collimator is fabricated of a radiation absorbing material, but has a plurality of closely arranged apertures, e.g., holes or pinholes. The apertures on the collimator allow only the radiation of interest to pass to the detector. Specifically, the radiation beams emitting from the object, if not absorbed or scattered by body tissue, exit the object along a straight-line trajectory. The collimator blocks or absorbs radiation beams that are not parallel to the axes of the apertures (openings in the collimator). Radiation beams traveling parallel to the apertures are detected by the radiation detector elements of the radiation detector. In one embodiment, the apertures of the collimator are uniform and may be perpendicular or skewed to the plurality of radiation detectors. In another embodiment, the apertures in the collimator are not uniform.

[0040] Preferably, the non-uniform collimator has a diverging fan-beam pattern of the apertures, which brings a wider field-of-view to the detector that may be useful for imaging large objects and for edge detection. Alternatively, the fan-beam pattern of the apertures can be used to obtain initial low resolution images. The low resolution information can be used to define the range of control parameters for subsequent high-resolution imaging. Since a wider FOV allows for faster imaging, it can be useful at the start of the imaging process. In some embodiments, the lower resolution system can serve as a guide for adaptation of a higher-resolution system to the specific application and object of interest. In contrast, the non-uniform collimator can also have a focus pattern of the apertures, which gives higher spatial resolution as the collimator magnifies the imaged object. For example, by using a convergent collimator, it is possible to reduce the spatial resolution of rotatable radiation imaging probe to 100 microns or less. In yet another alternative, a collimator can have an interwoven multi-aperture configuration for 3-dimensional radiation imaging applications, which is disclosed in an International Patent Application No. PCT/US2010/029409 assigned to Brookhaven Science Associates, which is incorporated by reference herein in its entirety. For use with interwoven multi-aperture collimation, two or more independent sets of cross-strip patterns on the cathode and anode sides would be formed; these sets of cross-strip electrodes would be matched to the interwoven multi-aperture configuration. Under such circumstances, three-dimensional images of an object can be obtained. The number of independent sets of cross-strip electrode geometries would be determined by the number of independent parallel pin-holes used to configure the inter-woven collimator.

[0041] In a preferred embodiment, the collimator may be constructed from a radiation- absorbing material known as the "high-Z" materials that have high densities and/or high atomic masses. Examples of such materials include, but not limited to, lead (Pb), tungsten (W), gold (Au), molybdenum (Mo), copper (Cu) and the composites thereof, such as a composite containing 70-99% of tungsten. The selection of the radiation-absorbing material and the thickness of the radiation-absorbent material should be determined so as to provide efficient absorption of the incident radiation, and would normally depend on the type of incident radiation and the energy level of the radiation when it strikes the surface plane of the collimator. The type of incident radiation and the energy level of the radiation depends on the particular imaging application, e.g., medical or industrial, or may be designed to be used in any of several different applications by using a general purpose radiation-absorbing material. In medical applications, for instance, in one embodiment, Indium- 1 1 1 (^{i n}In; 171 keV and 245 keV) and Technetium- 99m (^yymTc; 140 keV) are used as radioactive tracers for imaging of the prostate and other organs. In such applications, it is envisioned that the collimator may comprise copper, molybdenum, tungsten, lead, or gold.

[0042] In another embodiment for medical applications, Palladium- 103 (¹⁰³Pd; 21 keV) is used as a radioactive implant seed for treatment of the early stage prostate cancer. In such applications, it is envisioned that the collimator may be fabricated from copper, molybdenum, tungsten, lead, or gold. In one preferred embodiment, the collimator is fabricated from copper. In another preferred embodiment, the collimator is fabricated from tungsten. In yet another preferred embodiment, the collimator is fabricated from gold. The collimator body defining the surface plane may be fabricated of a solid layer of radiation-absorbing material of a

predetermined thickness, in which the plurality of apertures may be machined in any known manner according to optimized specifications. For example, a solid layer of radiation-absorbing material of a predetermined thickness may be machined or fabricated in a known manner, e.g. , using precision lasers, to achieve a collimator with the appropriate aperture parameters and aperture distribution pattern.

[0043] The collimator body containing the plurality of apertures may also be fabricated by laterally arranging septa of radiation-absorbing material so as to form predetermined patterns of radiation-guiding conduits or channels. In addition, the collimator body having a plurality of apertures may be manufactured by vertically stacking multiple layers of radiation-absorbing material with each layer having predetermined aperture cross-sections and distribution patterns so as to collectively form radiation-guiding conduits or channels. For example, multiple layers of lead, gold, tungsten, or the like may be vertically stacked to provide enhanced absorption of stray and scattered radiation to thereby ensure that only radiation with predetermined wavelengths is detected. In the case of vertically stacking multiple layers, the collimator may be formed by stacking repeating layers of the same radiation-absorbing material, or by stacking layers of different radiation-absorbing materials.

[0044] In the collimator, the aperture parameters such as aperture diameter and shape, aperture material, aperture arrangement, number of apertures, focal length, and acceptance angle(s) are not limited to specific values, but are to be determined subject to optimization based on required system performance specifications for the particular system being designed, as will be understood by those skilled in the art. Extensive patent and non-patent literature providing optimal configurations for apertures such as pinholes and parallel holes is readily available. Examples of such documentation are U.S. Patent No. 5,245,191 to Barber et al., entitled

Semiconductor Sensor for Gamma-Ray Tomographic Imaging System, and non-patent literature article entitled "Investigation of Spatial Resolution and Efficiency Using Pinholes with Small Pinhole Angled by M. B. Williams, A. V. Stolin and B. K. Kundu, IEEE TNS/MIC 2002, each of which is incorporated herein by reference in its entirety.

[0045] In a further embodiment, the collimator and the multi-layer radiation detector can have a side shielding surrounding the four (4) side surfaces of the collimator and detector combination. In a further preferred embodiment, back shielding can also be added so that the detectors are covered fully from all directions. In these embodiments, it is envisioned that the shielding has a uniform rectangular shape. However, in other embodiments, other shapes are also envisioned.

[0046] The present multi-layer semiconductor radiation detector(s) further encompass a method for radiation imaging of an object or tissue of interest. As illustrated in FIG. 5, in one embodiment, the method employs a PET imaging system with an annular detection gantry 107, which includes a plurality of multi-layer radiation detectors 101 distributed throughout the gantry 107 facing a patient's body 104. The number of multi-layer radiation detectors can range between 4 and 100. When radiopharmaceuticals are administrated into the patient's body 104 positioned on a movable bed 103, the radioactive tracer will concentrate in the specific tissues inside the target organ. The tracer will decay and emit gamma-ray photons in all directions with known energy {e.g. 140-keV gamma rays for Tc-99m, 27-36 keV gamma rays for 1-125, 171 keV and 245 keV gamma rays for In- 111 , and 364 keV gamma rays for 1-131). These gamma radiation photons will ionize the compound semiconductor and generate electron-hole pairs that are separated and guided to the contacts by the internal electric field inside the detectors. The generated electronic signals are readout by the preamplifier, signal processing circuit 105, and collected by a computer 106. The coincidence of the anode and cathode signals determines the position of gamma-ray interaction point inside the detector. Finally the position information is used for image reconstruction, and generated images are displayed on a display device 106. In one embodiment, the stack direction of multi-layer cross-strip detector with shared electrodes is pointing to the rotation center of the gantry of the imaging system. In another embodiment, the stack direction of the detector is parallel to the axis of the gantry. Yet, in another embodiment, the stack direction of the detector is parallel to the rotation direction of the gantry.

[0047] Numerous medical-, industrial-, scientific-, environmental cleanup-, and national and homeland security-applications exist for the detectors of the present multi-layer

semiconductor radiation detector(s). Some of the most prominent are imaging systems for detecting and localizing tumors and other abnormalities in the body, hand-held instruments to detect the trafficking of nuclear materials, and portable field instruments for environmental monitoring and remediation.

[0048] All publications and patents mentioned in the above specification are incorporated by reference in this specification. Various modifications and variations of the described detector(s) and its components will be apparent to those skilled in the art without departing from the scope and spirit of the invention. Although the disclosure has been described in connection with specific preferred embodiments, it should be understood that the invention as claimed should not be unduly limited to such specific embodiments. Indeed, those skilled in the art will recognize, or be able to ascertain using no more than routine experimentation, many equivalents to the specific embodiments of the invention described herein. Such equivalents are intended to be encompassed by the following claims.

Claims

CLAIMS:

1. A multi-layer semiconductor radiation detector comprising:

a plurality of thin detectors superimposed on top of each other, each detector having a semiconductor slab having a first and a second opposing surface and a linear array of electrodes positioned on both of said surfaces of the slab, oriented perpendicular to each other, wherein the linear arrays of electrodes interposed between two adjacent semiconductor slabs have the same orientation and form conductive contacts there between to share one or more readout channels.

2. The multi-layer semiconductor radiation detector of claim 1 , wherein the number of the linear arrays of electrodes interposed between adjacent semiconductor slabs is one.

3. The multi-layer semiconductor radiation detector of claim 1 , wherein the number of the linear arrays of electrodes interposed between adjacent semiconductor slabs is two or more.

4. The multi-layer semiconductor radiation detector of claim 1 , wherein the linear arrays of electrodes interposed between adjacent semiconductor slabs are connected by conductive epoxy.

5. The multi-layer semiconductor radiation detector of claim 1, wherein the linear arrays of electrodes interposed between adjacent semiconductor slabs are connected by cold- pressing of indium to form conducting bumps.

6. The multi-layer semiconductor radiation detector of claim 1 , wherein the linear arrays of electrodes interposed between adjacent semiconductor slabs further comprise a flexible cable positioned between the linear arrays of electrodes.

7. The multi-layer semiconductor radiation detector of claim 6, wherein the flexible cable forms connections with the detector electrodes and the connections are made with a conductive epoxy, cold-pressing of indium, or bump-bonding.

8. The multi-layer semiconductor radiation detector of claim 1 , wherein each semiconductor slab is a crystal having a defined length, width and height, wherein the length of the crystal is between 5 mm and 100 mm, the width of the crystal is between 5 mm and 100 mm, and the height of the crystal is between 0.1 mm and 20 mm.

9. The multi-layer semiconductor radiation detector of claim 8, wherein the length and the width of the crystal is about 40 mm and the height of the crystal is about 1 mm.

10. The multi-layer semiconductor radiation detector of claim 1 , wherein the number of semiconductor slabs is between 2 and 16.

11. The multi-layer semiconductor radiation detector of claim 1 , wherein each semiconductor slab is selected from the elements of groups III and V, groups II and VI, and group IV of the periodic table.

12. The multi-layer semiconductor radiation detector of claim 1 1, where the semiconductor slab is a single element crystal, a binary compound, a ternary compound or a ternary alloy.

13. The multi-layer semiconductor radiation detector of claim 1 1, wherein the semiconductor slab comprises CdZnTe (Cadmium Zinc Telluride), CdTe (Cadmium Telluride), CdMnTe (Cadmium Manganese Telluride), Hgl₂ (Mercury Iodide), TIBr (Thallium Bromide) or Si (Silicon).

14. The multi-layer semiconductor radiation detector of claim 13, wherein the semiconductor slab comprises CdZnTe (Cadmium Zinc Telluride).

15. The multi-layer semiconductor radiation detector of claim 1 , wherein the array of electrodes comprises 1 to 1000 strip electrodes.

16. The multi-layer semiconductor radiation detector of claim 15, wherein each strip electrode has the length approximately equal to the length of the semiconductor slab.

17. The multi-layer semiconductor radiation detector of claim 1 , wherein the radiation detected by the multi-layer semiconductor radiation detector is x-ray or gamma-ray radiation.

18. The multi-layer semiconductor radiation detector of claim 1 , wherein the linear arrays of electrodes interposed between adjacent semiconductor slabs are Ohmic contacts or Schottky contacts.

19. A high-resolution PET system comprising a gantry having a plurality of multilayer semiconductor radiation detectors oriented in a circumference around a central radial point of the gantry, a signal processing unit attached to the plurality of multi-layer semiconductor radiation detectors and a display unit connected to the processing unit,

wherein the multi-layer semiconductor radiation detector comprises a plurality of thin detectors superimposed on top of each other, each detector having a semiconductor slab having a first and a second opposing surface and a linear array of electrodes positioned on both of said surfaces of the slab, oriented perpendicular to each other, wherein the linear arrays of electrodes interposed between two adjacent semiconductor slabs have the same orientation and form conductive contacts there between to share one or more readout channels.

20. The high-resolution PET system of claim 19, wherein a planar surface of the semiconductor slabs of each multi-layer semiconductor radiation detector is perpendicular to a radial vector of the gantry.

21. The high-resolution PET system of claim 19, wherein a planar surface of the semiconductor slabs of each multi-layer semiconductor radiation detector is parallel to a radial vector of the gantry.

22. The high-resolution PET system of claim 19, wherein a planar surface of the semiconductor slabs of each multi-layer semiconductor radiation detector is perpendicular to a rotation center of the gantry.

23. The high-resolution PET system of claim 19, wherein a planar surface of the semiconductor slabs of each multi-layer semiconductor radiation detector is parallel to a rotation center of the gantry.

24. The high-resolution PET system of claim 19, wherein a planar surface of the semiconductor slabs of each multi-layer semiconductor radiation detector is perpendicular to an axis of the gantry.

25. The high-resolution PET system of claim 19, wherein a planar surface of the semiconductor slabs of each multi-layer semiconductor radiation detector is parallel to an axis of the gantry.

26. The high-resolution PET system of claim 19, wherein each multi-layer semiconductor detector has sufficient detection efficiency for 511-keV gamma-rays.

27. The high-resolution PET system of claim 19, wherein the dimension of each multi-layer semiconductor detector along the radius of the gantry is at least 5 mm to provide sufficient detection efficiency for 51 1-keV gamma-rays.

28. The multi-layer semiconductor radiation detector of claim 19, wherein the electrodes are Ohmic contacts or Schottky contacts.

29. The multi-layer semiconductor radiation detector of claim 19, wherein the linear arrays of electrodes interposed between adjacent semiconductor slabs are connected by conductive epoxy.

30. The multi-layer semiconductor radiation detector of claim 19, wherein the linear arrays of electrodes interposed between adjacent semiconductor slabs are connected by cold- pressing of indium to form conducting bumps.

31. The multi-layer semiconductor radiation detector of claim 19, wherein the linear arrays of electrodes interposed between adjacent semiconductor slabs further comprise a flexible cable positioned between the linear arrays of electrodes.

32. The multi-layer semiconductor radiation detector of claim 31 , wherein the connections between the flexible cable and the detector electrodes are made with conductive epoxy, cold-pressing of indium, or bump-bonding.

33. The multi-layer semiconductor radiation detector of claim 19, wherein each semiconductor slab is a crystal having a defined length, width and height, wherein the length of the crystal is between 5 mm and 100 mm, the width of the crystal is between 5 mm and 100 mm, and the height of the crystal is between 0.1 mm and 20 mm.

34. The multi-layer semiconductor radiation detector of claim 19, wherein the length and the width of the crystal is about 10-50 mm, and the height of the crystal is about 1-5 mm.

35. The multi-layer semiconductor radiation detector of claim 19, wherein the number of semiconductor slabs is between 2 and 16.

36. The multi-layer semiconductor radiation detector of claim 19, wherein each semiconductor slab is selected from the elements of groups III and V, groups II and VI, and group IV of the periodic table.

37. The multi-layer semiconductor radiation detector of claim 19, wherein the semiconductor slab is a single element crystal, a binary compound, a ternary compound or a ternary alloy.

38. The multi-layer semiconductor radiation detector of claim 19, wherein the semiconductor slab comprises CdZnTe (Cadmium Zinc Telluride), CdTe (Cadmium Telluride), CdMnTe (Cadmium Manganese Telluride), Hgl₂ (Mercuric Iodide), TIBr (Thallium Bromide) or Si (Silicon).

39. The multi-layer semiconductor radiation detector of claim 19, wherein the semiconductor slab comprises CdZnTe (Cadmium Zinc Telluride).

40. The multi-layer semiconductor radiation detector of claim 19, wherein the array of electrodes comprises 1 to 1000 strip electrodes.