Method and apparatus for reducing polarization within an imaging device

A method of reducing polarization within an image detecting device includes coupling at least one blocking contact to the image detecting device, and heating the image detecting device to facilitate reducing polarization within the image detecting device.

BACKGROUND OF THE INVENTION

This invention relates generally to imaging systems using pixilated detectors, and more particularly to pixilated semiconductor detectors in imaging systems.

Imaging devices, such as gamma cameras and computed tomography (CT) imaging systems, are used in the medical field to detect radioactive emission events for example, gamma rays in the range of 140 keV emanating from a subject, such as a patient and to detect transmission x-rays not attenuated by the subject, respectively. An output, typically in the form of an image that graphically illustrates the distribution of the sources of the emissions within the object and/or the distribution of attenuation of the object is formed from these detections. An imaging device may have one or more detectors that detect the number of emissions, and may have one or more detectors to detect x-rays that have passed through the object. Each of the detected emissions and x-rays may be referred to as a “count,” but the detected emissions may also be counted together as a ‘signal current’. The detector also determines the number of counts received at different spatial positions. The imager then uses the position dependent count tallies to determine the distribution of the gamma sources and x-ray attenuators, typically in the form of a graphical image having different colors or shadings that represent the processed count tallies.

A pixilated semiconductor detector, for example, fabricated from cadmium zinc telluride (CZT), may provide an economical method of detecting the gamma rays and x-rays. Specifically, at least one known imaging system includes a Room Temperature Semiconductor Radiation Detector (RTD) that is utilized to produce an image having a higher image quality. During operation, the RTD converts radiation photons to an electric charge (Q) using at least one of the photoelectric effect, the Compton effect, and/or electron-electron scattering. Converting photons directly to an electric charge facilitates eliminating the steps of light production and light detection and their corresponding inefficiencies that occur in the known scintillator technology. However, to operate at room temperature, RTD's must have sufficiently large Band Gap Energies (BG) to decrease the quantity of free charge carriers (N) in the material and allow the application of a higher bias voltage (Bias High Voltage HV). This allows the detection of signal pulses without producing a background electric current, referred to herein as a dark current (Id). During operation, the dark current can saturate the readout electronics, and/or reduce the signal to noise ratio (SNR) when measuring the signal electric charge (Q). To measure the signal electric charge (Q), detection electrodes and electronics are applied to surfaces of the RTD. Provided the charge mobility (μ) and carrier recombination lifetime (τ) are high enough, the bias high voltage causes the detection of the electric charge (Q) on the electrodes and electronics.

However, known detectors that are fabricated using a Cadmium Zinc Telluride (CZT) material may have a dark current (Id) that is not sufficiently controlled by the larger band gap. Accordingly, at least some known imaging systems include a cooling system to facilitate reducing the free charge carriers (N) and/or reducing the dark current (Id). For example, at least one known imaging system includes a cooling system that utilizes liquid nitrogen to facilitate reducing the free charge carriers (N) and/or reducing the dark current (Id). However, using a liquid nitrogen system is generally impractical for use in a commercial imaging system. Another known system uses circulating chilled water to control CZT and electronics temperatures, but this is also a significant cost in terms of engineering and safety. Moreover, at least one known imaging system utilizes a Peltier element to facilitate reducing the free charge carriers (N) and/or reducing the dark current (Id) which facilitates avoiding the adverse increase of the dark current (Id) that may be generated due to the heat of nearby objects, such as electronics, for example.

Accordingly, while known cooling systems may have a positive effect on reducing the dark current (Id), the cooling systems may have an adverse effect on the charge mobility (μ) and carrier recombination lifetime (τ). For example, when the quantity of impurities and band edge states within the intrinsic semiconductor device increases, i.e. reduced grade detectors and/or doped semiconductor devices, cooling may decrease the charge mobility (μ) and carrier recombination lifetime (τ) by increasing the interaction of the electric charge (Q) with these localized states, referred to as shallow and deep traps. More specifically, when the semiconductor device is fabricated using a Cadmium Zinc Telluride (CZT) materiel, where the charge mobility (μ) and carrier recombination lifetime (τ) product is marginal, such traps may be a limiting factor.

BRIEF DESCRIPTION OF THE INVENTION

In one aspect, a method of reducing polarization within an image detecting device is provided. The method includes coupling at least one blocking contact to the image detecting device, and heating the image detecting device to facilitate reducing polarization within the image detecting device.

In another aspect, an image detecting device is provided. The image detecting device includes a substrate, a blocking contact coupled to the substrate, and a heat source configured to increase a temperature of said substrate to facilitate reducing polarization within the image detecting device.

In a further aspect, an imaging system is provided. The imaging system includes a radiation source configured to emit a flux of photons, an image detecting device configured to receive the flux of photons and generate a response based on the flux of photons, wherein the image detecting device includes a substrate fabricated using a cadmium zinc telluride (CZT), a blocking contact coupled to the substrate, and a heat source configured to increase a temperature of the substrate to facilitate reducing polarization within said image detecting device.

DETAILED DESCRIPTION OF THE INVENTION

FIGS. 1 and 2illustrate an exemplary imaging system, for example, a Computed Tomography (CT) imaging system. System10is shown as including a gantry12representative of a “third generation” CT imaging system. Gantry12has an x-ray tube14(also called x-ray source14herein) that projects a beam of x-rays16toward a detector array18on the opposite side of gantry12. Detector array18is formed by a plurality of detector rows (not shown) including a plurality of detector elements20which together sense the projected x-rays that pass through an object, such as a medical patient22between array18and source14. Each detector element20produces an electrical signal that represents the intensity of an impinging x-ray beam and hence can be used to estimate the attenuation of the beam as it passes through object or patient22. During a scan to acquire x-ray projection data, gantry12and the components mounted therein rotate about a center of rotation24.FIG. 2shows only a single row of detector elements20(i.e., a detector row). However, multi-slice detector array18includes a plurality of parallel detector rows of detector elements20such that projection data corresponding to a plurality of quasi-parallel or parallel slices can be acquired simultaneously during a scan.

Rotation of components on gantry12and the operation of x-ray source14are governed by a control mechanism26of CT system10. Control mechanism26includes an x-ray controller28that provides power and timing signals to x-ray source14and a gantry motor controller30that controls the rotational speed and position of components on gantry12. A data acquisition system (DAS)32in control mechanism26samples analog data from detector elements20and converts the data to digital signals for subsequent processing. An image reconstructor34receives sampled and digitized x-ray data from DAS32and performs high-speed image reconstruction. The reconstructed image is applied as an input to a computer36, which stores the image in a storage device38. Image reconstructor34can be specialized hardware or computer programs executing on computer36.

Computer36also receives commands and scanning parameters from an operator via console40that has a keyboard. An associated display42, for example, a cathode ray tube or other suitable display device, allows the operator to observe the reconstructed image and other data from computer36. The operator supplied commands and parameters are used by computer36to provide control signals and information to DAS32, x-ray controller28, and gantry motor controller30. In addition, computer36operates a table motor controller44, which controls a motorized table46to position patient22in gantry12. Particularly, table46moves portions of patient22through gantry opening48.

In one embodiment, computer36includes an instruction reading or receiving device50, for example, a floppy disk drive, CD-ROM drive, DVD drive, magnetic optical disk (MOD) device, or any other digital device including a network connecting device such as an Ethernet device for reading instructions and/or data from a computer-readable medium52, such as a floppy disk, a CD-ROM, a DVD or another digital source such as a network or the Internet, as well as yet to be developed digital means. In another embodiment, computer36executes instructions stored in firmware (not shown). Computer36is programmed to perform functions described herein, and as used herein, the term computer is not limited to just those integrated circuits referred to in the art as computers, but broadly refers to computers, processors, microcontrollers, microcomputers, programmable logic controllers, application specific integrated circuits, and other programmable circuits, and these terms are used interchangeably herein. Although the specific embodiment mentioned above refers to a third generation CT system, the methods described herein equally apply to fourth generation CT systems (stationary detector—rotating x-ray source) and fifth generation CT systems (stationary detector and x-ray source). Additionally, it is contemplated that the benefits of the invention accrue to imaging modalities other than CT. Additionally, although the herein described methods and apparatus are described in a computed tomographic imaging system, it is contemplated that the benefits of the invention accrue to other medical and non-medical imaging systems such as a gamma camera, and/or those systems typically employed in an industrial setting or a transportation setting for Non-Destructive Testing, such as, for example, but not limited to, a baggage scanning system for an airport or other transportation center.

FIG. 3is a cross-sectional elevation view of an exemplary imaging device detector100, that can be used with imaging system10, in accordance with one embodiment of the present invention and includes a plurality of pixilated semiconductor detector elements102that may be used in connection with, for example, localizing a radiation interaction event in the detector, and a detector substrate104. Detector100may be formed of a radiation responsive semiconductor material, for example, cadmium zinc telluride (CZT) crystals. Detector elements102may be formed on substrate104by pixelating a corresponding plurality of pixel electrodes108, also referred to herein as anodes, coupled to a first surface110of detector substrate104(shown as a lower surface). A cross-sectional size and shape of pixel electrodes108and a spacing between each of the pixel electrodes108facilitates determining a location and size of each pixilated detector element102. Specifically, each pixilated detector element102is located proximate a second surface112(shown as an upper surface) of detector substrate104in substantial alignment with a longitudinal axis114of a corresponding pixel electrode108. An intrinsic spatial resolution of detector100may be defined by the size of and the spacing between each pixilated detector element102.

Moreover, in the exemplary embodiment, second surface112may be substantially covered by a single cathode electrode154. First surface110has a rectangular (or hexagonal or otherwise) array of small, for example between about one millimeters squared (mm2) and about ten mm2, generally square pixel electrodes108configured as anodes. A voltage difference applied between pixel electrodes108and cathode154during operation generates an electric field (detector field) in substrate104. The detector field may be, for example, about one kilovolts per centimeter to three kilovolts per centimeter. Although pixel electrodes108are described in the exemplary embodiment as being generally square, this shape should not be understood to be limiting, in that other shapes of pixel electrodes108are contemplated, in particular, round, or area filling shapes.

In operation, an uncollimated flux of photons144, for example emission gammas and transmission x-rays144, from a source140are directed towards second surface112. When a photon is incident on substrate104, it generally loses all its energy in substrate104by absorption and subsequent ionization and leaves pairs of mobile electrons156and holes158in a small localized region of substrate104. As a result of the detector field, holes158drift toward cathode154and electrons156drift toward pixel electrodes108, thereby inducing charges on pixel electrodes108and cathode154. The induced charges on pixel electrodes108are detected and identify the time at which a photon was detected, how much energy the detected photon deposited in the substrate104and where in the substrate104the photon interaction occurred using readout electronics160, for example. Moreover, the induced charge on154, can also be utilized by readout electronics160to determine timing and energy information.

In an alternative embodiment, the imaging device includes a collimator (not shown) that includes a plurality of apertures defined through the collimator. During operation, photons, for example emission gammas and transmission x-rays, from source140are directed towards and/or through the collimator. The photons are collimated utilizing the collimator, such that a collimated flux of photons is directed towards surface112.

FIG. 4is a flow chart illustrating an exemplary method200to facilitate reducing polarization within detector100. As used herein, polarization is defined as a buildup of electric charge that occurs within detector100. Accordingly, method200facilitates reducing and/or eliminating the conditions i.e charge trapping, that causes polarization.

In the exemplary embodiment, method200includes coupling202at least one blocking contact to an image detecting device, and heating204the image detecting device to facilitate increasing an operational temperature of the imaging device.

FIG. 5is a cross-sectional view of exemplary detector100that includes at least one blocking contact300. As used herein, blocking contact (electrode) is one for which charge carriers are impeded from entering the semiconductor material from the contact even though an applied voltage would force them that way. For example, blocking contact300may be fabricated utilizing a Gold and/or platinum layer that is applied as a cathode to a surface of the semiconductor detector and provides a potential energy step to the motion of the carriers. Thus as in a Shottky barrier junction the carrier motion is facilitated in the direction that releases energy and is impeded in the direction that absorbs energy. Shottky junctions and blocking contacts may be prepared in other fashions such as by stacking layers of doped semiconductors or by incorporating intermediate layers of dielectric material such as oxides as well as by other means. Not all metal electrodes are blocking contacts. For examples an Indium cathode is not a blocking contact for electron carriers. A unique characteristic of blocking contacts as described herein is that a photocurrent increases with the decreasing temperature of the detector. The blocking layer that is useful in limiting the dark current will be on the cathode side for the case that the electrons have higher mobility, and on the anode side for the case that the holes or positive carriers have the higher mobility. In CZT the electrons have the higher mobility. Accordingly, blocking contacts300are applied to a surface of detector100.

In the exemplary embodiment, detector100also includes an external heat source310that is configured to increase an operational temperature of detector100. For example, during normal operation, a temperature within detector100will either increase and/or decrease based on the power supplied to electronics attached to the detector100and/or a temperature of the outside environment in which detector100is operated. For example, the ambient air temperature within the examination room may either increase and/or decrease, thus either increasing and/or decreasing the operational temperature of detector100. Therefore, as used herein, the term operating temperature of detector100is used to define the temperature of detector100operated under typical conditions within a known detector operating environment.

Accordingly, and in the exemplary embodiment, detector100also includes external heat source310that is configured to increase an operating temperature of detector100to a temperature that is greater than the typical operating temperature. For example, and in the exemplary embodiment, heat source310includes a heating device312that is positioned on top of blocking contacts300. In the exemplary embodiment, heating device312includes a first electrically insulating layer314, a second electrically insulating layer316, and an electrical heating element318that is positioned between first and second layers314and316, respectively, First and second layers314and316are fabricated from a material that is optimally configured to transmit heat generated by heating element318to detector100, thus increasing an operational temperature of detector100. In the exemplary embodiment, first and second layers314and316are fabricated from a relatively thin plastic polymer material, for example.

During operation, heating element318is activated such that a temperature of heating element318is increased. Increasing a temperature of heating element318facilitates conductively heating first and second layers314and316, thus heating an external surface of detector100. In the exemplary embodiment, a voltage is applied to heating element318to facilitate activating heating element318. Specifically, as the voltage applied to heating element318is increased, an operating temperature of detector100is also increased. Thus, the voltage may be varied and controlled automatically to facilitate operating detector100at a desired temperature.

In another exemplary embodiment, heat source310includes a heating device320that is positioned adjacent detector100. In the exemplary embodiment, heating device320includes an electric fan assembly322, and a heat generating device324. In one embodiment, heat generating device324is a radiator, for example, that has a relatively warm heating fluid flowing therethrough. In another embodiment, heat generating device324is an electrical heating element that includes a plurality of openings extending therethrough.

During operation, fan assembly322is activated such that at least a portion of the air flow generated by fan assembly322is channeled through heat generating device324to facilitate increasing a temperature of the airflow. The heated air is then directed to an exterior surface of detector100, thus heating an external surface of detector100. In the exemplary embodiment, a temperature within detector100is maintained between approximately 10° C. and approximately 100° C. depending on other material factors.

FIG. 6is a perspective view of detector100that includes an insulating layer400that at least partially surrounds detector device100.FIG. 7is a top view of radiation detector100shown inFIG. 6. More specifically, detector100includes a bottom surface402, a first side404, a second side406, a third side408that is opposite first side404, and a fourth side410that is opposite second side406. Detector100also includes an upper surface412. In the exemplary embodiment, detector upper surface412is at least partially covered by blocking contact300as described previously herein. In the exemplary embodiment, heating device312is positioned on top of blocking contacts300to facilitate increasing a temperature within detector100. Moreover, in the exemplary embodiment, detector100also includes insulating layer400to facilitate retaining heat that is generated utilizing heating device312with detector100. More specifically, at least one of bottom surface402, first side404, second side406, third side408, and fourth side410may be, or may not be, at least partially covered with insulation layer400to facilitate retaining heat within detector100. In the exemplary embodiment, bottom surface402, first side404, second side406, third side408, and/or fourth side410are each substantially covered with insulation layer400to facilitate retaining heat within detector100. In another embodiment, bottom surface402, first side404, second side406, third side408, fourth side410, and heating device312, i.e. the upper surface of detector100, are covered with insulation layer400to facilitate retaining heat within detector100, i.e. detector100and/or multiple detector assemblies are completely encased together or separately within insulation layer400. In the exemplary embodiment, if insulating layer400is utilized to at least partially cover the surface of the incident radiation, for example, the cathode side of detector100, then the insulating layer400is fabricated from a material that is substantially transparent to the radiation, for example, insulating layer is between approximately 2 millimeters and approximately 5 millimeters in thickness, and fabricated from a foam rubber, for example.

In one embodiment, insulation layer400is a foam insulation that is selectively sized to at least partially encapsulate detector100. In another embodiment, insulation layer400is an insulating blanket formed using a fibrous insulation for example, that is wrapped around at least a portion of the exterior surfaces of detector100.

Utilizing insulation layer400as described herein facilitates maintaining the increased operational temperature within detector100. For example, during operation heating device312is utilized to increase the operating temperature of detector100to a predetermined operational temperature. When a temperature of the detector has reached the predetermined temperature, insulation layer400facilitates maintaining the detector at the predetermined temperature, thus the power consumed by heating device100is reduced, since heating device312is not utilized to compensate for heat losses to the surrounding environment, thus insulating layer400facilitates reducing detector heat loss, and thus reduces the total power consumed by detector100including heating device312. The insulation would also make the temperature easier to control by sensors and control circuits.

FIG. 8is a graph of measured data of a detector100with blocking contact300that is heated utilizing heat source310(shown inFIG. 5). The x-axis is the magnitude of the incident radiation flux as measured by the x-ray generator current control parameter. The y-axis is the measured count rate in the detector. The response curves of several pixels102,108and160are shown. At a relatively low temperature, approximately 26° C., the measured count rate (kcps) falls to approximately 0 at relatively high flux, whereas at a relatively high temperature (approximately 60° C.) the measured count rate continues to increase as a function of the incident flux (Ix).

FIG. 9is a graph of the photo current for a known detector during normal operation. Photo current is combination of the signal current attributable to the flow of the charge generated from absorbed x-ray or gamma ray photons and the dark current. In detectors known as ‘ohmic’ in the trade the dark current increases with flux, a phenomena sometimes called photoconductive gain. This dark current contribution to the photocurrent causes the same problems of amplifier overload, signal dependent heating described above for the dark current by itself.FIG. 10is a measured graph of detector100showing a reduced photo current under conditions of high flux and sustained operation achieved using the methods and apparatus described herein. More specifically,FIG. 9illustrates along the Y-axis, the dark current per pixel,102,108, and160of CZT measured. As shown, the dark current (Id) increases with Flux and Temperature for a CZT sample fabricated using non blocking, or ohmic contacts, for example.FIG. 10illustrates that although the photo current (Id) still increases with flux using a blocking contact300, heating the detector utilizing heat source310substantially eliminates the increase commensurate with the flux. The calculation in the inset box show that the photo-current is equal to the signal current alone with no significant contribution from dark current caused by photoconductive gain. Thus, at low temperature the current increases with flux and the signal disappears with flux. This behavior is supposed to be caused by polarization charge causing increased current emanating from the cathode and decreased E field in the detector bulk. Accordingly, utilizing both blocking contact300and heat source310facilitates both reducing dark currents and reducing polarization within detector100, Reducing polarization in turn reduces the dark current emanating from the cathode while at the same time preserving the E field in the bulk and the thereby the detector operation even at high flux.

The above-described imaging device detectors provide a cost-effective and reliable means for reducing polarization within an imaging detector. More specifically, the imaging detector includes a blocking contact and a heat source that work together to reduce currents and polarization within detector100.

The detector components illustrated are not limited to the specific embodiment described herein, but rather, components of the detector may be utilized independently and separately, or repetitively from other components described herein. For example, the detector components described above may also be used in combination with different imaging systems. A technical effect of the embodiment of the systems and methods described herein includes improving the detector performance, i.e. high flux, dark current, and spectral improvement, within a detector fabricated using a CZT material, by coupling at least one blocking contact to the detector, and heating the detector to a temperature that is greater than the typical operating temperature of known detectors.

Moreover, the combination of heating the detector and utilizing a blocking contact allows mainly the unique combination of lowered dark current and higher charge mobility and lifetime. This in turns allows critical improvements of the energy spectrum and high rate operation for the applications of NM and CT. Additionally, the combination of heating the detector and utilizing a blocking contact facilitates preventing high flux desensitization, signal dependent instabilities, and energy response degradation, while increasing the useful flux limit. Moreover, reducing the dark currents allows the detector to be directly coupled to the readout electronics. Direct coupled electronics are significantly simpler to build and allow reduced input capacitance. Reduced input capacitance allows reduced noise and increased signal. Low dark current prevents signal dependent heating in the semiconductor which is a source of gain and count rate instability. Signal dependent instabilities are very important and cannot be calibrated away and thus may prevent the use of semiconductor detectors in the critical applications of medical imaging. Thus, Improving detector response at high and low fluxes utilizing the methods and apparatus described herein facilitates a manufacturer to utilize lower and/or relatively less expensive grades of CZT material to fabricate the detector. Accordingly, the detector described herein can achieve photon counting in transmission medical imaging, thereby providing the signal noise (sometime called Swank noise) reduction anticipated by photon counting techniques. This noise reduction can result in improved image quality or reduced patient dose. Photon counting in transmission medical imaging can also reduce the noise derived from dark current in current mode detectors that affects the parts of the image derived from the smallest detector signal amplitudes. Moreover, the single response, i.e. reduced spectral tail, semiconductor radiation detector described herein can be utilized in a nuclear medicine application such a simultaneous dual isotope imaging where the energies of the two sources are too close to be resolved by the known detectors. The successful operation of CZT for the NM application with a reduced tail spectral component also allows the sensitivity/resolution tradeoff currently made to be further optimized. Improved energy resolution in transmission imaging can allow energy resolution and therefore materials determination in a calcium scoring application, for example.