JP2009018154A - Photon counting x-ray detector with overrange logic control - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a photon counting X-ray detector with overrange logic control. <P>SOLUTION: A CT detector includes a first detector 252 configured to convert radiographic energy to electrical signals representative of energy sensitive radiographic data and a second detector 254 configured to convert radiographic energy to electrical signals representative of energy sensitive radiographic data and positioned to receive X-rays that pass through the first detector 252. A logic controller 262 is electrically connected to the first detector 252 and the second detector 254 and is configured to receive a logic output signal from the second detector 254 indicative of an amount of a saturation level of the first detector 252, compare the logic output signal to a threshold value, and output, based on the comparison, electrical signals from the first detector 252, the second detector 254, or a combination thereof to an image chain. <P>COPYRIGHT: (C)2009,JPO&INPIT

Description

  The present invention relates generally to diagnostic imaging, and more particularly to a stacked detector with logic control for selecting data usage for a pair of detectors.

  In general, in X-ray imaging systems such as X-ray and computed tomography (CT), an X-ray source emits X-rays toward a subject or object such as a patient or luggage. Hereafter, the terms “subject” and “object” can be used interchangeably to denote anything that can be imaged. After the beam is attenuated by the subject, it strikes the array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array generally depends on the x-ray attenuation. Each detector element of the detector array generates a separate electrical signal that represents the attenuated beam received at each detector element. The electrical signal is sent to a data processing system for analysis, which ultimately generates an image.

  In other typical X-ray imaging systems, positron emission tomography (PET), or single photon emission computed tomography (SPECT), the radiation source in the imaged object is X Emitting rays, X-rays are captured by a photon counting energy sensitive X-ray detector. The CT system is combined with a PET or SPECT system to create a fusion system (CT / SPECT or CT / PET) that provides images representing both anatomical structures and physiologically important (ie functional) information be able to. Such a combination system includes a source that emits X-rays towards an X-ray detector and a separate SPECT or PET detector that measures the X-rays emitted from a radiation source within the object.

  In some CT imaging systems, for example, the x-ray source and detector array are rotated in the imaging plane around the subject in the gantry. X-ray sources for such CT imaging systems typically include an X-ray tube that emits X-rays as a fan beam emanating from the focal point. Such an X-ray detector for a CT imaging system is generally configured within an arc centered on the focal point. In addition, such a detector includes a collimator for collimating the X-ray beam received by the detector having a focal point at the focal point. In addition, such a detector includes a scintillator for converting X-rays to light energy adjacent to the collimator and receives light energy from the adjacent scintillator and generates an electrical signal from the light energy. Including diodes. In general, each scintillator of the scintillator array converts X-rays into light energy. Each photodiode detects light energy and generates a corresponding electrical signal as a function of the light emitted by the corresponding photodiode. The output of the photodiode is then sent to a data processing system for image reconstruction.

  In some SPECT and PET systems, for example, one or more flat detector arrays are rotated in the imaging plane around the subject in the gantry. An X-ray radiation source in the imaged object emits photons in random directions. X-ray detectors typically include a collimator for collimating the X-ray beam received by the focused detector so as to be parallel rays contained within the imaging plane and orthogonal to the detector plane. In addition, such a detector includes a scintillator for converting X-rays to light energy adjacent to the collimator and receives light energy from the adjacent scintillator and generates an electrical signal from the light energy. Including the multiplier, the electrical signal is then sent to a data processing system for image reconstruction.

  Conventional CT imaging systems utilize a detector that converts the radiographic energy into a current signal that is integrated over a period of time and then measured and finally digitized. However, the difficulty with such detectors is that they cannot provide data or feedback regarding the number and / or energy of the detected photons. That is, a conventional CT detector has a scintillator component and a photodiode component. The scintillator component emits light when receiving X-ray imaging energy, and the photodiode detects the light emission of the scintillator component, and the intensity of the light emission. Providing an electrical signal as a function of Although it is generally recognized that CT imaging will not be a practical diagnostic imaging tool without the progress achieved using conventional CT detector designs, the difficulty with these detectors is that they are energy discriminating data. Or alternatively, counting the number of photons actually received by a given detector element or pixel and / or being unable to measure its energy. That is, the light emitted by the scintillator is a function of the number of X-rays struck and the energy level of the X-rays. Under the charge integration mode of operation, the photodiode is not capable of discriminating energy levels or photon numbers from scintillation. For example, two scintillators can emit light with equal intensity, thus providing equal output to each photodiode. Nevertheless, the number of x-rays and x-ray intensity received by each scintillator can be different, but result in equal light output.

  A typical PET or SPECT system uses a photon counting energy discrimination detector comprised of a scintillator and a photomultiplier tube. Such detectors have large detector elements and are therefore not easily adapted to CT applications that require high resolution images to obtain anatomical details in the imaged object. Thus, recent detector developments include the design of energy discriminating direct conversion detectors that can provide photon counting and / or energy discriminating feedback with high spatial resolution. In this regard, the detector can be operated in an X-ray counting mode, an energy measurement mode for each X-ray event, or both.

  An energy discriminating direct conversion detector can provide not only x-ray counting but also a measurement of the energy level of each detected x-ray. As a result, such a detector could potentially be used for SPECT or PET imaging. Although many materials can be used in the direct conversion energy discrimination detector configuration, semiconductors such as cadmium zinc telluride (CZT) and cadmium telluride (CdTe) have been shown to be preferred materials. .

  Generally, charge integration or charge pulse counting electronics are utilized to generate an image using a direct conversion detector. In the case of charge integration, the charge is integrated over the entire pixel area, and the integrated charge is used to generate a digital signal that represents the total number of x-rays and the energy level of each detected x-ray. In the case of pulse counting, the current pulse amplitude is compared to a threshold level and stored in a bin counter to record the photon energy.

  However, the drawbacks of direct conversion semiconductor detectors are that these types of detectors are commonly encountered in conventional CT systems at the x-ray photon flux rate, eg, 5-100 million per second per square millimeter. That is, it is impossible to count above the count (Mcps). Very high x-ray photon flux rates above 5-100 Mcps cause pile-up and polarization, which ultimately leads to detector saturation. That is, these detectors typically saturate at relatively low x-ray flux level thresholds. Above these thresholds, the detector response is not predictable or has a worse dose utilization. For SPECT and PET, the imaging flux level is below 5-100 Mcps, and such saturation in semiconductor detectors for SPECT and PET is not a practical problem. However, in the case of CT, saturation can occur at a detector location where a thin subject is placed between the detector and a radiographic energy source or x-ray tube. These saturation regions have been shown to correspond to thin subject paths near or outside the width of the subject projected onto the detector fan arc. In many cases, the subject is more or less circular or elliptical in the effect on attenuation of the X-ray flux and subsequent incident intensity on the detector. In this case, the saturation region corresponds to two separate regions at the end of the sector arc. As another less common, but rare case, saturation occurs at other locations, three or more separate areas of the detector. For elliptical subjects, saturation at the end of the fan arc is reduced by placing a bowtie filter between the subject and the x-ray source. In general, the filter is configured to the shape of the subject such that the total attenuation of the filter and the subject is equalized across the sector arc. In that case, the flux incident on the detector is relatively uniform across the sector arc and does not cause saturation. However, a problem may be that the bowtie filter may not be optimal if the subject population is not quite uniform and the shape is not exactly elliptical. In such cases, one or more isolated saturation regions can occur, or conversely, the x-ray flux can be over-filtered to produce very low flux regions. The low x-ray flux in the projection ultimately contributes to noise in the reconstructed image of the subject.

  Detector saturation causes loss of imaging information and results in artifacts in X-ray projections and CT images. In addition, hysteresis and other non-linear effects occur at flux levels near and above detector saturation. Direct conversion detectors are susceptible to a phenomenon called “polarization”, and when polarization occurs, charge trapping within the material changes the internal electric field and counts the detector in an unpredictable way. And alters the energy response, resulting in hysteresis where the response is altered by previous exposure histories. In particular, the photon counting direct conversion detector saturates due to the unique charge collection time (ie, dead time) associated with each x-ray photon event. Saturation occurs due to pulse pileup when the x-ray photon absorption rate for each pixel is on the order of the reciprocal of the charge collection time. The charge collection time is approximately proportional to the thickness of the direct conversion layer when the electric field and anode contact size are fixed, and therefore, the saturation rate can be increased when the direct conversion layer is thinner. However, a sufficient thickness is required to stop almost all X-rays. Incomplete acquisition of X-rays results in degraded image quality, i.e. noisy images, and poor utilization of the dose applied to the object being imaged.

  A further factor in charge collection time is the voltage applied in the thickness direction of the layer. A larger electric field (voltage / thickness) results in a charge collection time that is inversely smaller and a saturation rate that is proportionally larger. However, there is a reliability problem for the transfer of high voltage signals. Higher reliability can be obtained when a lower voltage is applied to a smaller thickness of the direct conversion layer. However, again, a sufficient thickness of the layer is required to sufficiently stop most of the x-rays.

  In addition to direct conversion detectors, other types of detectors saturate. A common example is a scintillator-photodiode configuration connected to an integrating preamplifier. The charge generated from each photon is transferred to the preamplifier. As the x-ray flux increases, the current flowing into the preamplifier or the total charge integrated over the integration period increases. The readout electronics have a limiting current or charge capability before saturating the amplifier. Amplifier saturation is associated with a non-linear response and loss of signal charge. This again results in poor dose utilization and image artifacts.

  Another detector configuration is a scintillator over photodiode connected to photon counting readout electronics. Similar configurations utilize a scintillator over avalanche-photodiode on an avalanche photodiode or a scintillator over photo-multiplier tube on a photomultiplier tube. The saturation of the X-ray flux rate in these photon counting cases is also related to the dead time for erasing the charge before the arrival of the next X-ray photon.

For photon counting direct conversion detectors, no practical solution to X-ray flux rate saturation is known in an imaging system that uses an X-ray source operating above the saturation point. For these systems, the total thickness of the X-ray absorption layer can be greater than 1.0 mm, for example. The higher the X-ray energy, the greater the thickness required to sufficiently stop most of the X-ray flux. A typical target value is to stop 90-99% of incident x-rays. In the case of two possible direct conversion materials used for X-ray spectroscopy, cadmium zinc telluride (CZT) or cadmium zinc (CdTe), most of the X-rays generated from the source at 100-200 kVp are stopped The thickness for diagnostic X-ray fluoroscopy and CT imaging is approximately 3.0 to 5.0 mm. For CZT or CdTe, for a pixel size on the order of 1.0 mm and a thickness on the order of 3.0-5.0 mm, a saturation limit of approximately 5-100 × 10 ⁶ X-rays / second / mm ² is generally used. Found. This limit is directly related to the charge collection time for CZT. Higher flux rates are theoretically possible by using smaller pixels. Each pixel has a size-independent count rate limit set by the charge collection time. The saturation flux rate is determined by dividing the count rate limit by the pixel area. Thus, the saturation flux rate increases as the pixel size decreases. Smaller pixels are also desirable because they make available higher spatial resolution information that can result in high resolution images. However, the small pixel size results in higher costs and there are more channels per unit area that need to be connected to readout electronics.

In addition, smaller pixels or detector elements have a larger perimeter to area ratio that results in greater crosstalk. The outer periphery is a region where charge is shared between two or more pixels (ie, crosstalk occurs). This charge sharing results in incomplete energy information and / or miscounting of X-ray photons, since the readout electronics are not configured to combine simultaneous signals at neighboring pixels. Very thin flux rates are possible with pixel size less than 0.1 mm and using thin photon counting direct conversion silicon layers, but these thin layers have sufficient stopping power to stop X-rays. not exist. In the case of an integral detector, the detector pixel size and preamplifier design are balanced to account for the expected x-ray flux rate during imaging. In the case of CT, the flux rate capability of a detector with integrating electronics is generally on the order of 10 ⁹ photons / second / mm ² . For an x-ray projection imager operating with a charge-preserving integral detector, the flux rate capability is simply on the same order. In the case of a photon counting detector using a scintillator and one of a photodiode / APD / photomultiplier tube, the dead time of the X-ray conversion layer is very fast and the dead time is usually an electron that can also be relatively high. Related to device readout bandwidth. The problems with these detectors vary. In the case of a photodiode, the electronics gain is not sufficient to overcome the electronics noise. In the case of APD, there is additional gain, but with associated gain-instability noise, temperature sensitivity, and reliability issues. In the case of photomultiplier tubes, these devices are too large and costly for high resolution detectors that cover a large area.

  Detector saturation can affect image quality by limiting the number of photons used to reconstruct an image and introducing image artifacts. A minimum image quality and therefore a minimum flux rate is required to use the image. In this regard, if the system is configured so that sufficient flux is received in one area of the detector, the other area of the detector is probably high enough to saturate the detector in this area. Likely to receive higher flux. Higher flux in these other regions is not necessary for image quality, but data loss due to detector saturation may need to be addressed by a correction algorithm to reduce image artifacts. In the case of CT imaging, reconstruction is not resistant to lost or erroneous data. For example, if the center of the detector is illuminated with the lowest flux for image quality purposes and the illuminated object is compact, the detector cell at the outer edge of the object's shadow and the detector cell outside the outer edge will be in these projection directions. It can be saturated due to the thickness of the thin object. Reconstruction of data set using these uncorrected saturation values causes severe artifacts in the image.

  Many imaging techniques have been developed to deal with saturation of any part of the detector. These techniques include maintaining low x-ray flux across the width of the detector array, for example by using low tube currents or currents that are modulated per view. However, this solution results in increased scan time. That is, there is a trade-off that the acquisition time for an image increases in proportion to the nominal flux required to acquire a certain number of x-rays that meet the image quality requirements.

  For combined CT and SPECT or CT and PET imaging, the availability of energy discriminating detectors with high flux rate capabilities provides an opportunity for shared detectors. The SPECT X-ray photon energy is similar to that in CT, so that the thickness of the semiconductor layer can be designed to meet both CT and SPECT requirements. However, for PET, the photon energy is 511 eV, about 5 times higher than that used for CT and SPECT.

  In typical imaging applications, x-rays are absorbed by the direct conversion material, which results in the generation of charge in the direct conversion material. In order to generate digital image information, the generated charge is collected at the divided anode, typically using either charge integration or charge pulse counting electronics.

  However, one difficulty with direct conversion semiconductor detectors can be that X-rays absorbed by the direct conversion material near or near the gap in the anode can result in a charge shared by at least two neighboring pixel anodes generated there. That is. When using charge integration electronics, charge sharing can appear as crosstalk between neighboring pixels, thus making the electronics susceptible to electronic noise amplification and making the image spatially blurred. When using pulse counting electronics, charge sharing can result in charge splitting between at least two anodes, with the amplitude of the charge pulses collected at the at least one anode being below the discrimination threshold. If it does, it results in a loss count. In addition, in the case of pulse counting, high energy x-rays can result in loss of detection quantum efficiency (DQE) by generating two or more counts collected at two or more neighboring anodes, Thus, the event is miscounted, eg, binning a single high energy event as two or more low energy events.

  Another difficulty with direct conversion semiconductor detectors is that the responses at the edges and corners of the direct conversion crystal may not be reproducible. Such locations in the direct conversion crystal generally have a charge trapping center that causes a change in the internal electric field when the incident X-ray flux changes. A changing internal electric field can cause a poor detector response that can lead to image quality problems.

  Many techniques have been developed to address charge sharing in direct conversion detectors. Energy discriminating detectors typically include a number of segmented anodes, typically 0.2-2.0 mm, that define a pixelated structure onto which the direct conversion material is electrically coupled. The anode defines the response area of the imaging pixel that divides the area of the detection surface. Smaller pixels are generally desirable because smaller pixels make higher spatial resolution information available, which can result in higher resolution images, and flux rate capability is generally improved with smaller pixels. However, smaller pixels can result in higher costs because there are more channels per unit area that need to be connected to readout electronics. In addition, smaller pixels or detector elements have a larger perimeter to area ratio that results in a larger percentage of charge sharing regions per unit area of the detector.

  Since the perimeter is an area where charge can be shared between two or more pixels, the readout electronics are not configured to combine nearly simultaneous signals in neighboring pixels, so that incomplete energy information and / or x-ray photons A false count occurs. The readout electronics can include a time match circuit configured to identify events that occur within a defined time window, and when the time window identifies the event, the detected event receives a bin count. To prevent. However, such electronic devices can be costly and difficult to implement.

In order to solve the problem regarding the reproducibility of the response at the edges and corners of the direct conversion crystal, a guard ring is generally placed on the anode face of the device or on the side wall of the crystal wall. However, the guard ring does not prevent charge trapping in the semiconductor and does not prevent a changing electric field from developing in the semiconductor.
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  Therefore, it is desirable to design a direct conversion energy discriminating CT detector that can produce images above the x-ray photon flux rate that typically causes saturation in the direct conversion detector. In addition, it is desirable to design an x-ray management system that accommodates changes in x-ray flux across the CT detector assembly, compensates for overrange or saturated detectors, and eliminates charge sharing within the direct conversion detector. Such a detector and flux management system allows the same detector to be used for both CT and SPECT imaging.

  The present invention relates to a multilayer CT detector operating at a very high counting rate that overcomes the above-mentioned difficulties.

  A CT detector is disclosed that enables energy discrimination, energy integration, and direct conversion. Also disclosed is a dual modality detector that enables both CT and single photon emission computed tomography (SPECT) detection. The detector includes multiple layers of semiconductor material. In this regard, the detector is configured to be split in the X-ray penetration direction to optimize count rate performance and avoid saturation. In addition, CT detectors can be fabricated with multiple detector elements or subpixels per contact area. In this context, a dynamic and flexible combination of the outputs of the individual detector elements can be implemented in order to prohibit the use of data from the saturation detector elements.

  CT detectors support not only X-ray photon counting, but also energy measurement or tagging as well. As a result, the present invention supports the acquisition of both anatomical details and tissue feature information. In this regard, energy discrimination information or data can be used to reduce effects such as beam hardening. In addition, these detectors support the acquisition of tissue discrimination data and thus provide diagnostic information representative of disease or other lesions. For example, it is possible to detect calcium in plaques in the view. This detector can also be used to detect, measure and characterize substances that can be injected into a subject, such as contrast agents, and other specialized substances such as targeted drugs. The contrast material can include, for example, iodine that is injected into the bloodstream for better visualization.

  Thus, according to one aspect of the present invention, a CT detector includes a first detector configured to convert X-ray imaging energy into an electrical signal representing energy-sensitive X-ray imaging data; And a second detector arranged to receive X-rays transmitted through the first detector. A logic controller is electrically connected to the first detector and the second detector, receives a logic output signal from the second detector representing the amount of saturation level of the first detector, and outputs a logic output signal Is configured to output an electrical signal from the first detector, the second detector, or a combination thereof to the image chain based on the comparison.

  According to another aspect, an x-ray imaging system includes a radiation source configured to project x-rays toward a scanned subject, a first detector layer, and a second detector layer. A detector assembly, wherein the second layer is arranged to receive x-rays from a radiation source transmitted through the first detector layer, and each of the first and second layers is from the radiation source. A detector assembly configured to receive projected x-rays and convert x-ray imaging energy into an electrical signal representative of energy sensitive x-ray imaging data; and at least one of the first and second detector layers Receiving data representative of the flux rate, determining whether an electrical signal has been output at the first detector layer for image reconstruction based on the received data, and determining image reconstruction based on the received data For In the second detector layer and a device configured to determine whether the output electric signal.

  According to another aspect, a method of making an imaging system includes providing an X-ray source, positioning a first detector to receive X-rays emitted from the X-ray source, and X Positioning a second detector to receive x-rays emitted from the source and transmitted through the first detector; and electrically coupling the logic device to the first and second detectors And wherein the logic device is configured to represent a level of x-ray flux at one of the first and second detectors.

  Various other features and advantages of the present invention will be made apparent from the following detailed description and the drawings.

  The drawings illustrate various embodiments presently contemplated for carrying out the invention.

  The operating environment of the present invention is described with respect to a 64-slice computed tomography (CT) system. However, one skilled in the art will appreciate that the present invention is equally applicable for use in other multi-slice configurations. Furthermore, the present invention will be described with respect to X-ray detection and conversion. However, those skilled in the art will further appreciate that the present invention is equally applicable to the detection and conversion of other high frequency electromagnetic energy. Although the present invention is described with respect to a “third generation” CT scanner, it is equally applicable to other CT systems.

  Referring to FIG. 1, a computed tomography (CT) imaging system 10 is shown as including a gantry 12 representing a “third generation” CT scanner. The gantry 12 has an x-ray source 14 that projects an x-ray beam 16 toward a detector assembly or collimator 18 on the opposite side of the gantry 12. Referring now to FIG. 2, the detector assembly 18 is formed by a plurality of detectors 20 and a data acquisition system (DAS) 32. The plurality of detectors 20 sense the projected x-rays that pass through the patient 22, and the DAS 32 converts the data into digital signals for later processing. Each detector 20 generates an analog electrical signal that represents the intensity of the impinging x-ray beam and thus represents the attenuated beam as it passes through the patient 22. During a scan to acquire x-ray projection data, the gantry 12 and its attached components rotate around the center of rotation 24.

  The rotation of the gantry 12 and the operation of the X-ray source 14 are managed by the control mechanism 26 of the CT system 10. The control mechanism 26 includes an X-ray controller 28 that provides power and timing signals to the X-ray source 14 and a gantry motor controller 30 that controls the rotational speed and position of the gantry 12. The image reconstructor 34 samples and digitizes the X-ray data from the DAS 32 and performs high speed reconstruction. The reconstructed image is provided as input to the computer 36, which stores the image in the mass storage device 38.

  Computer 36 receives commands and scanning parameters from an operator via console 40, which has some form of operator interface, such as a keyboard, mouse, voice activated controller, or any other suitable input device. . An associated display 42 allows the operator to view the reconstructed image and other data from the computer 36. 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, the computer 36 also operates a table motor controller 44 that controls the motorized table 46 to position the patient 22 and the gantry 12. Specifically, the table 46 moves the patient 22 so that the whole body or a part thereof passes through the gantry opening 48 of FIG.

  As shown in FIG. 3, the detector assembly 18 includes rails 17 that have a collimating blade or plate 19 therebetween. The plate 19 is arranged to collimate the x-ray 16 before the x-ray beam impinges on the detector 20 of FIG. 4 located on the detector assembly 18 for example. In one embodiment, detector assembly 18 includes 57 detectors 20, each detector 20 having pixel elements 50 with an array size of 64 × 16. As a result, detector assembly 18 has 64 rows and 912 columns (16 × 57 detectors), so that each time the gantry 12 rotates, 64 simultaneous slice data are collected. Enable.

  Referring to FIG. 4, the detector 20 includes a DAS 32, and each detector 20 includes a number of detector elements 50 arranged in a pack 51. The detector 20 includes pins 52 that are positioned within the pack 51 relative to the detector element 50. The pack 51 is disposed on a backlit diode array 53 having a plurality of diodes 59. The backlit diode array 53 is disposed on the multilayer substrate 54. Spacers 55 are disposed on the multilayer substrate 54. The detector element 50 is optically coupled to a backlit diode array 53 that is electrically coupled to a multilayer substrate 54. A flex circuit 56 is attached to the surface 57 of the multilayer substrate 54 and the DAS 32. Detector 20 is positioned within detector assembly 18 using pins 52.

  In operation of one embodiment, X-rays that impinge in the detector element 50 generate photons that traverse the pack 51, thereby generating an analog signal, the analog signal being a backlit diode array. Detected on diode in 53. The generated analog signal is transmitted to the DAS 32 through the multilayer substrate 54 and the flex circuit 56, where the analog signal is converted into a digital signal.

  As explained above, each detector 20 is designed to convert radiographic energy directly into an electrical signal containing energy discrimination data or photon count data. The present invention contemplates many configurations for these detectors, their components, and the manner in which data is read. Despite the differences between each of these embodiments, each detector shares two common features. One of these features is a multi-layer configuration of semiconductor films or layers. In one embodiment, each semiconductor film is fabricated from cadmium zinc telluride (CZT). However, one skilled in the art will readily recognize that other materials that allow direct conversion of radiographic energy can also be used. Another feature common between the various embodiments is the use of intrusion or intervening metallized films or layers that separate the semiconductor layers. As will be described, these metallization layers are used to apply a voltage to the semiconductor layer and to collect electrical signals from the semiconductor layer. As will also be described, detectors having such a design have improved saturation characteristics and photon count fidelity.

It is generally well known that the charge collection time of a semiconductor layer is inversely related to the maximum period count rate saturation threshold (MPR) of the layer. Thinner layers have faster charge collection and higher MPR. However, the thinner layer stops only a fraction of the incident x-rays. The charge collection time is approximately proportional to the dimension d, which is the smaller of the detector thickness and the pixel contact size, but the radiographic energy deposition efficiency increases exponentially with thickness. To increase. The count rate performance of CZT can be defined by MPR = μ _e E / d.

From this definition, assuming that contact size and thickness are equal, d = 0.3 cm, electric field E is 1000 V / cm, and μ _e is about 1000 cm ² / (V seconds), 3.0 megacounts / second A maximum period counting rate can be achieved. Since the arrival of X-rays is random rather than periodic, significant saturation effects occur at a rate 10 times lower than average. In other words, the count rate of a 3.0 mm thick CZT semiconductor layer can have a count rate capability in the range of 0.3 to 3.0 megacounts / second. However, as will be described, the configuration of a direct conversion semiconductor detector using multiple layers having a cumulative thickness corresponding to a single thicker layer can improve count rate performance.

  In addition, the thinner conversion layer not only improves count rate performance by reducing charge collection time, but also provides improved charge collection efficiency, thereby reducing polarization and detector count and energy response fidelity It also provides improvements. A thinner conversion layer also reduces charge sharing between pixel elements, thereby improving energy discrimination performance and spatial resolution.

  The improvement in flux rate performance by dividing the detector into multiple thin layers can be attributed to many factors. First, having multiple layers distributes the total flux rate between the different layers. Each layer experiences only a portion of the total flux. For example, incomplete x-ray attenuation of a thin first layer (compared to the x-ray attenuation depth), the saturation of this layer is at a higher count rate than the saturation of a thick layer that stops all x-rays Guarantee that.

  The second factor is that even if one layer is saturated, another layer is not saturated and the layer thickness can be configured to ensure that it provides valid data about that view. is there. For example, if one of the layers is configured to stop only 5% of the x-rays, the x-rays saturate a thick layer designed to stop all x-rays at 20 times the flux rate. The third factor is that the charge collection time decreases as the layer thickness and pixel size decrease. Charge collection time is approximately proportional to the smaller of the thickness or pixel contact size divided by the mobility in the layer and the electric field. A smaller thickness and / or pixel size provides a higher flux rate limit for the layer.

A fourth factor is that thinner layers also provide reduced crosstalk. The effect of pixel size on crosstalk is approximately given by the effective perimeter of the total pixel area. That is, crosstalk is increased or decreased 4W × aT / ^{W 2} as a coefficient, wherein, W is the pixel pitch, aT is the charge diffusion length, which is proportional to the thickness of the layer (charge spreading length). Thus, crosstalk is reduced as the layer thickness decreases. The competition between the flux rate saturation effect and the crosstalk effect can be traded off by examining their impact on the detection quantum efficiency (DQE (f)), an important figure of excellence in X-ray imaging detectors. it can. DQE decreases as a function of counting rate and does not decrease much for thinner layers. A design method for optimizing the number and thickness of layers is predicted when the maximum count rate is obtained after DQE (f) drops below any point on some threshold curve.

  The fifth factor is the reduction of polarization due to more efficient collection of electrons and holes. In thinner layers, electrons and holes can be collected by traveling a smaller distance, and thus electrons and holes are less susceptible to trapping.

  In addition to these five factors for improved count rate limits when using thin layers, flux rate limits (ie, count rates per unit area) are smaller pixels preferred in thin layers due to reduced crosstalk. Improved by using size.

  Referring now to FIG. 5, a portion of a two-layer CZT or direct conversion detector 20a according to one embodiment of the invention is shown in perspective view. The detector 20 a is defined by the first semiconductor layer 62 and the second semiconductor layer 64. During the fabrication process, each semiconductor layer 62, 64 is configured to have a number of electronic pixelated structures or pixels to define a number of sensing elements or contacts 65. This electronic pixelation is achieved by providing a 2D array 67, 69 of electrical contacts 65 on the layers 62, 64 of direct conversion material. Further, in one embodiment, this pixelation is defined two-dimensionally in the width direction and length direction of each semiconductor layer 62, 64.

  The detector 20a includes high voltage electrodes 66 and 68 adjacent to the semiconductor layers 62 and 64, respectively. Each high voltage electrode 66, 68 is connected to a power source (not shown) and is designed to power the respective semiconductor layer during the x-ray or gamma ray detection process. One skilled in the art understands that each high voltage connection layer should be relatively thin to reduce the x-ray absorption characteristics of each connection layer, and in one embodiment is a few hundred angstroms thick. Let's be done. As described in more detail below, these high voltage electrodes can be secured to the semiconductor layer by a metallization process.

  Referring now to FIG. 6, a cross-sectional view along line 6-6 of FIG. 5 illustrates the relative thickness of each semiconductor layer 62, 64 in one embodiment. In addition, for this embodiment, the pixel pitch and contact size are similar for both layers and are approximately equal to the thinner thickness. As with the high voltage electrodes 66, 68, the 2D contact arrays 67, 69 should absorb a minimum of radiographic energy. Each array or signal collection layer is designed to provide a mechanism for outputting an electrical signal generated by the semiconductor layer to a data acquisition system or other system electronics. A mechanism is provided for flexibly combining signals for pixels in different layers, as will be described in more detail later. One skilled in the art will appreciate that many (possibly hundreds) of interconnects (not shown) are used to connect all contacts 65 to the CT system electronics. Furthermore, the cumulative thickness of 3.0 mm of the semiconductor conversion layer is a thickness at which 99% of the X-rays are absorbed. Absorption is calculated using a physics model for a 120.0 kvp spectrum with 20.0 cm filtering, which is a typical X-ray spectrum for medical CT acquisition.

  In addition, as shown in FIG. 6, the thicknesses of the semiconductor layers 62 and 64 are different from each other, but the pixel pitch and contact size are the same. The layers are arranged in a specific order with respect to the X-ray direction in order to take advantage of the exponential absorption characteristics. In the case of FIG. 6, when X-rays enter from the upper side toward the lower common contact portion array, more X-rays are given in the semiconductor layer 62 than in the semiconductor layer 64. For example, assuming that the semiconductor layer 62 has a thickness of 1 millimeter and the semiconductor layer 64 has a thickness of 2.0 mm, the semiconductor layer 62 is expected to absorb about 92% of the x-rays, The second semiconductor layer 64 is expected to absorb about 7% of the X-rays. The combined total absorption of the two layers reproduces 99% efficiency of the 3.0 mm layer. One important advantage of this configuration when compared to a single layer of 3 mm thickness is the reduction of the polarization effect for these two thinner layers. This advantage itself allows operation with a 10-fold increase in flux rate in most practical applications.

  In addition, by combining the counting responses from the two layers using a specific self-correction algorithm, the split detector, detector 20a, is compared to a single layer of semiconductor material with a thickness of 3.0 mm. Can be configured to provide increased count rate performance. For example, a CT detector as described herein can be configured to have a first layer that absorbs 92% of the incident x-ray flux and a second layer that absorbs 7%. I want to think about. As a result, the second layer saturates at a flux rate 14 times higher than a 3.0 mm thick layer. As the incident flux rate increases, the second layer saturates or overranges with an x-ray flux that is greater than the x-ray flux required to saturate the 3.0 mm thick layer. This variability in the saturation characteristics of multiple layers of a single CT detector allows the overrange layer output to be estimated by the effective signal in the non-overrange or non-saturation layer. In this connection, the saturation state of a given detector layer is detected and, as a result, the signal for the saturated layer or an equivalent 3.0 mm thick layer is experienced from the output of the unsaturated layer in the detector. Estimated.

  An example of this self-correctability algorithm is that at high count rates above which the first layer is saturated, only the count response from the second layer weighted by fractional absorption is Assigned to that projection for each pixel. At low count rates, a weighted sum of responses from both layers is assigned to the projection for each pixel. A more sophisticated algorithm is that if the statistical error in a layer value increases as the count rate increases, the signal for the two layers is added to the combined sum so that the value is reduced in weight. Can be combined with a weighted inverse of DQE.

  It is contemplated that the CT detector assembly may be configured such that each CT detector is configured with such overrange correctability. However, it is also contemplated that only detectors in detector assemblies that are generally associated with overranging are configured with this overranging self-correctability. For example, detector detector peripheral detectors typically encounter higher flux conditions than more centrally located detectors. In this regard, the peripheral detector can be configured with overranging self-correctability, but the more centrally located detector is not so configured. In addition, layers using other detection mechanisms and detector materials that have high count rate capability but have poor count rate and / or energy response characteristics can be detected to estimate the count rate and energy response of the saturated layer. Can be used in certain parts of the vessel.

  In addition, it is contemplated that a given CT detector can have more than two semiconductor layers. In this regard, the effective signal output of two or more unsaturated layers can be used to estimate the output of the saturated layer. For example, the detector may include a first layer having a 35 times effective response, a second layer having a 10 times effective response, and a third layer having an effective response equivalent to the effective response of the first layer. And can be configured to comprise: In this connection, the first and third layers saturate with a higher x-ray flux layer than the second layer. Thus, if the second layer is overranged or saturated, the outputs of the first and third layers are used to compensate or effectively determine the overranged second layer output. Can do.

  Referring now to FIG. 7, another design contemplated for CZT or direct conversion detector is shown. In this embodiment, detector 20b also includes a pair of semiconductor layers 74,76. In contrast to the previously described embodiments, detector 20b includes a single common signal collection layer or 2D contact array 78. This single but common array 78 is designed to collect electrical signals from both semiconductor layers 74, 76 and output them to DAS or other system electronics. In addition, detector 20b also includes a pair of high voltage electrodes 80,82. Each high voltage electrode effectively operates as a cathode, and the contact portion of the 2D array 78 operates as an anode. In this regard, the voltage applied through the high voltage connections 80, 82 creates a circuit that goes through each semiconductor layer to the signal collection contact array 78.

  Yet another contemplated embodiment is shown in FIG. As shown in this embodiment, detector 20c includes four semiconductor layers 84, 86, 88, 90. The detector 20c further includes two conductive lines or paths 92, 94 that are electrically connected to the high voltage electrodes 87, 89, 91 and the collection contact arrays 93, 95. The conductive path 92 receives and moves electrical signals from the contact array 93, 95. In this regard, a single data output is provided to the CT system electronics. Similar to a single signal acquisition lead, a single high voltage connection 94 is used to power the four semiconductor layers 84-90 via electrodes 87, 89, 91. The detector 20c only requires a single high voltage connection.

  Referring to FIG. 9, a monolithic embodiment of the present invention is shown. Similar to the embodiment of FIG. 8, detector 20d includes four semiconductor layers 96-102. Each semiconductor layer 96-102 is connected to a pair of conductive layers. In this connection, one conductive layer is used for applying a voltage and the other conductive layer is used for collecting electrical signals generated by the respective semiconductor layers. In order to minimize the number of conductive layers, the detector 20d utilizes an alternating conductive layer structure. That is, every other conductive layer is used for high voltage connections and the other conductive layers are used for signal collection. In this connection, the conductive layers 104, 106, 108 are used for the application of relatively high voltages, and the layers 110, 112 include contacts for signal collection. As such, the high voltage connection layers 104, 108 are used to apply voltages to the semiconductor layers 96, 102, respectively. The high voltage connection layer 106 is used to apply a voltage to the semiconductor layers 98 and 100.

  As explained above, in one embodiment, each semiconductor layer is composed of a CZT material. One skilled in the art will appreciate that there are many techniques that can be used to construct such semiconductors. For example, molecular beam epitaxy (MBE) is one method that can be used to produce each thin layer of CZT material. Screen printing of CZT particles on a polymeric adhesive is a potentially low cost, low temperature method of forming a layer on a flexible wiring board. One skilled in the art will appreciate that many techniques can be used to metallize the semiconductor layer to provide the conductive connections described above.

  Furthermore, metallization can also be used to provide signal feedthrough to the collection contact, as shown in FIG. As shown, a single layer of semiconductor material 114 is sandwiched with an array 116 of collection contacts and a high voltage electrode layer 118. Prior to metallization of the semiconductor layer 114 to form the collection contact array 116 and the high voltage electrode layer 118, holes 120 are etched into the semiconductor 114 or otherwise formed. Thereafter, the holes 120 can be metallized to provide a signal feed path 122 from each collection contact 124. A signal feedthrough or conduction path 122 is configured in the hole 120 so as not to contact the adjacent adjacent high voltage electrode layer 118. In this regard, the signal path is directed to a bus (not shown) designed to move the electrical signals emitted by the individual collection contacts 124 to the CT system electronics throughout the detector. Extending in the vertical or X-ray receiving direction. As a result, a laminated structure is formed that consists of a series of thin layers laminated in the X-ray direction.

  So far, the present invention has been described with respect to a multilayer CT detector designed with different layer thicknesses but similar pixel sizes.

  For multi-layer CT detectors that include direct conversion semiconductor layers of varying thickness to reduce the possibility of saturation or overrange of energy sensitive CT detectors at the X flux rate commonly encountered in conventional CT scans. The invention has been described. However, as described below, the present invention is an energy sensitive overranging tolerant CT detector that utilizes variability in the electrical contacts of a multilayer CT detector to improve the saturation characteristics of the CT detector. Also related.

  Referring now to FIG. 11, an exploded side view of a portion of a CT detector that includes the present invention is shown. As shown, the detector 126 is formed on a substrate 128 that is secured to a detector frame (not shown) by fasteners (not shown). The substrate 128 supports a pair of detector layers 130, 132. Each detector layer 130, 132 is comprised of a radiation conversion component and a signal acquisition component. In the illustrated embodiment, layers 130, 132 having different thicknesses are stacked in the x-ray direction 134 and separated from each other by a flex layer 136. The CT detector 126 also includes a high voltage bias line 138 connected to the detector layer 130 to bias the detector assembly.

  As mentioned above, the detector 126 includes a pair of detector layers 130, 132. The detector layer can include a scintillator and a photodiode that are matched to a conventional CT detector or made from a direct conversion semiconductor material such as CZT, which can be in many detector elements or pixels. Combined. As shown in FIG. 11, the detector layer 130 differs from the detector layer 132 in the thickness of each conversion component 140, 142 and the number and size of each detector element array 144, 146. As shown, detector element array 144 is half the pixel pitch and four times the number of detector elements 148 than detector element array 146 for equal detector area. In addition, the contact area of detector element 148 is a quarter of the contact area of detector element 150. As will be explained, this variation of the detector element array within a single CT detector greatly enhances the saturation characteristics of the detector.

  By changing the size of the detector elements within a given detector 126, the charge collection time associated with each layer of the detector is changed. That is, those skilled in the art will readily appreciate that the charge collection time decreases as the conversion layer thickness decreases and the detector element size decreases. That is, the charge collection time of the detector layer is approximately proportional to the smaller of the conversion layer thickness or the detector element size divided by the mobility and electric field in the detector layer. The count rate saturation threshold increases with decreasing pixel size. Furthermore, for a given count rate saturation threshold proportional to area reduction, a smaller pixel area implies a higher flux rate saturation threshold. As such, as the conversion layer thickness and / or detector element size decreases, the flux rate limit for the corresponding detector layer increases, thereby improving the saturation characteristics for that layer of the CT detector. To do. This improvement and variability in saturation characteristics allows the detector to be configured so that some layers can withstand higher x-ray flux levels and provide input to a self-correctability algorithm.

  In the embodiment shown in FIG. 11, voltage bias lines or leads 138 extend directly from the substrate 128 to the conversion layer 130. However, as shown in FIG. 12, it is contemplated that the high voltage bias line 138 is disposed in contact with the flex layer 136. The flex layer 136 constitutes the routing layer and is used to connect the individual detector elements 148 to the readout electronics of the CT scanner 10, namely the DAS and the image reconstructor. In the embodiment shown in FIG. 12, the voltage bias line 138 can be metallized on the face of the flex layer 136 and the detector element 148 can be metallized on the opposite face.

  FIG. 13 shows the detector layers 130, 132 and their associated components from different directions. As shown, the thickness difference between the direct conversion components 140, 142 can be varied from the thickness difference shown in FIG. 11 to achieve different absorption and flux rate characteristics. For example, assuming that the detector element 148 has a pixel size of 0.25 mm and the detector element 150 has a pixel size of 1.0 mm, the detector layer 130 assumed to have a thickness of 0.4 mm. Stops one-half of the X-rays 134 impinging there, and the conversion layer 142 assumed to have a thickness of 4.6 mm causes the other half of the X-rays 134 not absorbed by the layer 140 to be absorbed. Stop. Absorption is calculated using a physics model for a 140 kvp spectrum with 3.0 cm filtering, which is a typical X-ray spectrum of a medical CT system. Compared to a single layer with a thickness of 5.0 mm, both detector layers 130, 132 each have a two-fold improvement in flux rate capability due to incomplete absorption in each of the layers. Further, layer 130 has an area that is 1/16 of a 1.0 mm pixel, resulting in a flux rate saturation threshold that is 16 times higher for 1.0 mm pixels. Further, the combination of reduced layer thickness and reduced detector element size was reduced by a factor of 4 for layer 130 compared to a single 5.0 mm detector layer with a 1.0 mm pixel size. Provides charge collection time. The flux rate improvement by each of these mechanisms is multiplicative. As a result of this combination of incomplete absorption (2 ×), smaller area (16 ×), and reduced charge collection time (4 ×), the total flux rate saturation threshold for detector layer 130 is 1. It can be 128 times higher compared to a 5.0 mm thick layer with a detector element pitch of 0 mm. Furthermore, layer 130 has less polarization due to improved charge collection time in the thin layer.

  Referring now to FIG. 14, the order of the layered detectors is reversed compared to the embodiment of FIG. This reversal results in 99% x-ray absorption at detector layer 132 with a detector element pitch of 1.0 mm. Therefore, only 1% of X-rays are left for absorption in detector layer 130 having a detector element pitch of 0.25 mm. Thus, the absorbed flux rate portion for detector layer 130 is 100 times smaller than that of a single layer detector having a thickness of 5.0 mm and a pitch of 0.25 mm. The detector layer 132 achieves a 6-fold increase in flux rate capability due to 4 times faster charge collection time and 1/16 pixel area. In total, the multilayer detector has a 6400 times improvement in flux rate performance compared to a single layer detector with a thickness of 5.0 mm and a detector element pitch of 1.0 mm.

  Referring now to FIG. 15, it is contemplated that the detector 126 may be configured with more than two detector layers. For example, the detector 126 can be designed to have three separate detector layers 130 (a), 130 (b), 132. For illustrative purposes, the detector of FIG. 15 is oriented in the same direction as the detector shown in FIG. 14 with the addition of another detector layer. Therefore, the detector shown in FIG. 15 includes a relatively thick conversion layer 140 and two relatively thinner conversion layers 142 (a), 142 (b). Further, the detector element pitch for detector layer 132 is four times that for detector layers 130 (a), 130 (b). A detector having the configuration shown in FIG. 15 is intended to perform different operations than the detectors described so far.

  Specifically, for the detector 126 of FIG. 15, at lower x-ray flux levels, no detector layer saturates and a smaller detector element 148 (a) to provide a single signal. 148 (b) are combined. That is, assuming that detector elements 148 (a), 148 (b) are a quarter pitch of detector element 150, the count data for detector elements 148 (a), 148 (b) is It is binned in a 4 × 4 manner so as to be equal to the pitch of the detector elements 150 of the detector layer 132. For medium flux levels, the detector layer 132 is saturated and only count data from the detector layers 130 (a), 130 (b) is used. The third layer can be configured to have a 1000 times saturation threshold compared to a 5.0 mm thick single layer detector assembly with a detector element pitch of 1.0 mm.

  Referring now to FIG. 16, it is contemplated that a vertical arrangement of CT detector components can also be used to achieve improved count rate performance. The detector 126 includes three direct conversion layers that are identical in thickness and size. The direct conversion layers 152 are separated from each other by a flex layer 154 and an array of detector elements 156. When using this configuration, the thickness of the conversion layer defines the detector element pitch in one direction, and the spacing between the detector elements defines the pitch in the other direction. Further, when using this arrangement, the response of the detector element is a function of the detector element “height” over a particular conversion layer. For example, a detector element in row 158 of elements having a size of 0.7 mm collects charge from approximately 1% of the x-rays impinged on the detector, and a detector element in row 159 having a size of 4.3 mm Absorbs almost 99% of the X-rays impinged on the detector. Thus, the flux rate saturation for element row 158 is 100 times greater than that of a single detector having a detector element pitch of 1.0 mm and a thickness of 5.0 mm.

  As mentioned above, the present invention relates to achieving improved saturation characteristics of CT detectors and assemblies using multiple direct conversion layers. The present invention also relates to achieving improved saturation characteristics of CT detectors through a reduction in detector element size. Each detector element of a CT detector is commonly referred to as a “pixel” and, therefore, in one embodiment, the present invention relates to “sub-pixelation” of a pixel region.

  Referring now to FIG. 17, a single pixel region 160 (shown in dashed lines) is pixelated into four equally sized subpixels 162. In the example shown, each subpixel 162 is connected to a dedicated read lead 164. Since the flux at a pixel is proportional to its area, the combined flux rate saturation threshold of four separate subpixels 162 is four times that achieved by a single pixel 160 covering the area of four subpixels 162. is there. In addition, each sub-pixel 162 has a faster charge collection time due to its size reduction with respect to the layer thickness. Faster charge collection time represents a greater saturation flux rate limit that simply exceeds the improvement in count rate performance achieved by a reduction in detector element size. Note that since each sub-pixel 162 has a similar size, the sub-pixel saturates at approximately the same x-ray flux level.

  On the other hand, referring to FIG. 18, the region achieved by a single pixel 160 can be pixelated into sub-pixels having different flux rate characteristics. For example, as shown in FIG. 18, subpixel 162 (a) is significantly larger than subpixel 162 (b). This asymmetry in subpixel size results in a composite pixel region having a different saturation threshold within the composite pixel region. Specifically, assuming that subpixel 162 (a) is 20 times larger than subpixel 162 (b), subpixel 162 (a) is saturated at an x-ray flux threshold that is 20 times that of subpixel 162 (b). To do.

  It is contemplated that any number of placements may be performed to place subpixel 162 (a) relative to subpixel 162 (b). In the arrangement shown in FIG. 19, the central arrangement of subpixels 162 (b) is believed to improve crosstalk characteristics between subpixels. That is, in the arrangement shown, sub-pixel 162 (b) is less likely to be affected by adjacent sub-pixels, so that crossing from sub-pixel 162 (a) when compared to the arrangement of FIG. You can better escape from talk.

  It will be appreciated that the flux rate is not the same everywhere in the CT detector. In this context, the present invention also includes an x-ray flux management system that detects and / or predicts saturation of a given portion of the CT detector so that appropriate correction measures can be taken. For example, it is well known that the end of the CT detector assembly often receives more x-ray flux than the central portion of the CT detector assembly due to the subject profile and the pre-subject filter attenuation profile. As such, it is contemplated that post-acquisition logic can be used to use only the output of the non-saturated channel for image reconstruction. In another embodiment, saturation of a given portion of the CT detector assembly is expected, and as a result, the portion of the CT detector that is expected not to be saturated is electrically connected to the scanner's DAS and expected to be saturated. The selection scheme (binning scheme) is established so that the parts are not connected. In yet another embodiment, the connectivity of detector elements to the system DAS is determined for each view during data acquisition. That is, the previous view data and other prior information is used to connect the detector elements to the DAS. This scheme provides dynamic but flexible selection of detector elements during data acquisition. In another embodiment, the detector element connectivity to the DAS is controlled in real time. In this regard, connectivity can be changed during data acquisition for a given view so that the connection is opened if a high photon rate is detected.

  A schematic diagram of a loop of detector elements or subpixels 164 for a given layer or array of detector elements, with each subpixel 164 electrically routed to a respective data system input 166 is shown in FIG. ing. In the example shown, the group is composed of four subpixels 164, so that four outputs 168 are provided at four data system inputs 166. Four outputs of the 4 × 4 DAS channel are input to the truth table circuit 170. The output 172 of truth table circuit 170 is a linear combination of the four inputs 174 to truth table circuit 170 and depends on whether any of the inputs are saturated. Each of the sub-pixels shown in FIG. 20 is designed to saturate with the same x-ray flux, but one sub-pixel of the group is adjacent by giving the contour of the subject to be imaged and pre-subject filtering. It is also possible to saturate without subpixel saturation.

  The following table shows a truth table for combining the outputs of the four subpixels. In the truth table, the value “1” represents non-saturation and the value “0” represents saturation. Therefore, if no subpixel is saturated, the value “1” is input to truth table circuit 170 for each subpixel. The truth table indicates that in such a situation, the outputs from all subpixels are considered acceptable and combined to provide a single output for that group of subpixels. On the other hand, if the subpixel associated with channel “A” or more precisely, channel “A” is saturated, but the other subpixels are not saturated, the sum of the unsaturated channels is output by the truth table circuit. The data associated with the saturation channel is ignored. For example, assuming a pixel region comprised of subpixels where one subpixel has a higher x-ray flux saturation threshold than another subpixel in the pixel region, if the x-ray flux is low, both subpixels Provide a valid output that is summed by the truth table circuit to provide a single output composed of count data from both sub-pixels. When the X-ray flux reaches a level at which only one of the sub-pixels is saturated, only the data from the non-saturated sub-pixel becomes the data output by the truth table circuit.

与えられたサブピクセルの飽和を決定するために、多くの技法が使用され得ることが認識されよう。例えば、与えられたサブピクセルについての計数レートデータは、閾値と比較されることができ、サブピクセルによって決定された計数レートが閾値を超える場合、そのサブピクセルについて、「０」の飽和値が真理表回路に入力される。例えば、検出器システムが、飽和閾値を毎秒１００万カウントとする直接変換検出器を使用して光子をカウントするように設計される場合、この閾値が、各サブピクセルに課される閾値レベルとなり、またはその何パーセントが、飽和閾値を引き下げるマージンを提供する。

It will be appreciated that many techniques can be used to determine the saturation of a given subpixel. For example, the count rate data for a given subpixel can be compared to a threshold, and if the count rate determined by the subpixel exceeds the threshold, a saturation value of “0” is true for that subpixel. Input to the table circuit. For example, if the detector system is designed to count photons using a direct conversion detector with a saturation threshold of 1 million counts per second, this threshold is the threshold level imposed on each sub-pixel, Or a percentage of that provides a margin to lower the saturation threshold.

  The flexible selection of subpixels within a given pixel area as described above allows the detector to output photon count data despite the saturation of some parts of the detector. It is also contemplated that it can be achieved for further enhancement. Described with respect to FIG. 20 was a signal management scheme that utilizes a single DAS channel for each subpixel. However, given many subpixels within a single CT detector, one DAS channel per subpixel may not be feasible. Thus, the present invention also contemplates a signal control scheme that utilizes one DAS channel for a group of subpixels. In this regard, the number of DAS channels required can be equal to the number required for CT detectors that do not include sub-pixelation.

  Referring now to FIG. 21, a switched network based signal management system is shown in which two or more subpixels are dynamically controlled and connected to a DAS input channel. In this regard, the output of each subpixel 164 in a given group of subpixels is input to a switch network 176. The switch network is designed to reconnect the subpixel outputs based on the subpixel saturation. The switch network can utilize a truth table to dynamically control the connectivity of the subpixel output. For saturated subpixels, the switch network discards them so that only unsaturated data is included in the output 180.

  For example, at low x-ray flux, none of the subpixels are saturated, so the output 168 from all subpixels 164 is combined into a single output 180 that is input to DAS 182. The DAS 182 includes a signal shaper 184 that is configured to extract a single photon event from the output of the switch network. It will be appreciated that a low noise / fast charge amplifier (not shown) can be connected to receive the output of the switch network. Thereafter, the output of the amplifier is input to the signal shaper 184. Signal shaper 184 provides an input to energy level discriminator 186. The energy level discriminator 186 is connected to the signal shaper 184 and is designed to filter the photons based on the energy level of the photons and compared to one or more thresholds. To achieve this goal, photons with energy levels outside the desired range are excluded from the processing for counting and image reconstruction. At a minimum, the discriminator 186 is designed to eliminate photons having energy levels that correspond to noise in the system. It is contemplated that multiple thresholds can be used to define the energy level range. The counting register 188 is configured to receive photons that have not been rejected by filtering at the energy level discriminator 186, count the number of photons received at the detector, and provide a corresponding output.

  DAS 182 counts the number of photons for a given pixel area 165 composed of given subpixels 164. Since the switch network does not connect the output of a given subpixel if a given subpixel is saturated, DAS determines the number of photons from only unsaturated subpixels. Although only four subpixels are shown, it is contemplated that a given pixel region can be subpixelized into fewer or more than four subpixels.

  Two switch states are shown in FIGS. As shown in FIG. 22, if the subpixel identified by input “A” is not saturated and all other subpixels are saturated, then only the output of subpixel “A” is Used for photon counting. As shown in FIG. 23, the switch network allows any combination of subpixel outputs, such as subpixels “A” and “C” where only subpixels “A” and “C” are unsaturated subpixels. Can be made.

  In one embodiment of the invention, the first detector is configured to operate in a photon counting mode. The second detector is configured to receive X-rays transmitted through the first detector and directly convert the X-ray imaging energy into an electrical signal including energy discrimination or photon count data, wherein the photon counting mode or energy Can be configured to operate in an integrated mode. The present invention contemplates many configurations for these detectors, their components, and the manner in which data is read. In one embodiment, the energy detector semiconductor of the first detector is fabricated from cadmium zinc telluride (CZT). One skilled in the art will readily recognize that other materials capable of direct conversion of radiographic energy can be used. The second detector can include a conventional scintillator array on a photodiode array or preferably an energy discriminating semiconductor configured to operate in an energy integration mode.

  FIG. 24 shows a perspective view of a detector module according to an embodiment of the present invention. The detector module 200 includes a first detector layer 202 that includes a direct conversion material 206 that is typically made from semiconductor materials such as cadmium zinc telluride (CZT) and cadmium telluride (CdTe). The first detector layer 202 also has a plurality of pixelated anode contacts 208 attached to the first surface 210 of the direct conversion material 206. Pixelated anode contact 208 is attached to and electrically coupled to first flexible interconnect 209. The direct conversion material 206 includes a high voltage electrode 212 attached to the second surface 213 of the direct conversion material 206. The high voltage electrode 212 effectively functions as a cathode. One skilled in the art will appreciate that the high voltage electrode 212 should be relatively thin to reduce x-ray absorption characteristics. Therefore, the high voltage electrode 212 is preferably a few hundred angstroms thick. The high voltage electrode 212 can be affixed directly to the conversion material 206 by a metallization process, such as evaporation, and typically includes a metal such as gold, platinum, silver, copper, or aluminum. A high voltage lead 215 is attached to the high voltage electrode 212. The first detector layer 202 includes a spacer 203 disposed on the substrate 234, the spacer 203 having a hole 205 formed to aid in grid alignment.

  The second detector layer 204 of the detector module 200 includes a scintillator 230 that is optically coupled to the photodiode array 232. The photodiode array 232 includes a two-dimensional array of diodes, similar to the plurality of diodes 59 shown in FIG. A photodiode array 232 is attached to and electrically coupled to the second flexible interconnect 236.

  The substrate 234 provides mounting structure and mechanical support for the first and second flexible interconnects 209, 236. In the embodiment shown in FIG. 24, the substrate 234 is preferably an electrically inert substrate that does not need to bridge electrical connections across or through it. However, those skilled in the art will appreciate that the substrate 234 can also be electrically active to bridge electrical connections from one site to another. Substrate 234 also provides attachment structure and mechanical support to a pair of spacers 203 to which X-ray attenuation grid assembly 214 can be attached. The grid assembly 214 blocks or attenuates X-rays 16 emitted toward the outer edge or periphery 224 of the direct conversion material 206 and blocks X-rays 16 emitted toward the charge sharing region 226 of the direct conversion material 206. Or arranged to attenuate.

  In one embodiment, the grid assembly is between the first detector and the x-ray source to attenuate x-rays directed toward the gap between anodes or charge sharing regions forming a direct conversion detector. Be placed. X-rays that impinge on the direct conversion material generate a charge within the volume of the direct conversion material, and the charge generally moves through the direct conversion material to an anode structure arranged to collect the charge. Are known. Detector operation in imaging applications requires charge measurement by collecting charge on a segmented anode. The divided anode defines the response area of the detector surface. If x-rays impinge directly on the conversion material in the region near the gap between the anodes, charge may be shared between at least two adjacent anodes. Such charge sharing can cause image artifacts or other image problems that can be mitigated by the use of a grid structure that prevents X-rays from colliding directly with the conversion material in the charge sharing region. Can do.

  The detector module 200 includes an x-ray attenuation grid assembly 214. The grid assembly 214 includes a plurality of bars or bars 216, 218 that are spaced between or along the outer perimeter frame 220 and form an opening 222. The outer frame 220 is disposed substantially along the outer periphery 224 of the direct conversion material 206. X-ray attenuation grid assembly 214 includes a hole 207 that substantially matches the position of hole 205 in spacer 203. Accordingly, the holes 207, 205 serve as a positioning mechanism for the detector module 200 for an imaging system such as the imaging system 10 of FIG. The perimeter frame 220 is arranged to block or attenuate x-rays emitted from an x-ray source, such as the x-ray source 14 of FIG. 1, sent towards the outer edge or perimeter 224 of the direct conversion material 206. Similarly, the bars 216, 218 are arranged to block or attenuate X-rays emitted from an X-ray source, such as the X-ray source 14 of FIG. 1, toward the charge sharing region 226 of the direct conversion material 206. .

  Still referring to FIG. 24, according to one embodiment of the present invention, the second detector layer 204 of the detector module 200 includes a scintillator 230 optically coupled to the photodiode array 232. The photodiode array 232 is attached to and electrically coupled to the second flexible circuit 236, and the second flexible circuit 236 is attached to the substrate 234.

  In operation, x-rays 16 emitted from an x-ray source, such as x-ray source 14 of FIGS. In general, the low energy x-rays 16 interact directly with the conversion material 206, which results in the generation of charge therein that is transferred to at least one of the pixelated anode contacts 208. A pixelated anode contact 208 that defines the response area of the sensing surface collects charge that travels directly through the conversion material 206, which has a DAS integrated circuit (IC) 241 where the charge is converted to a digital signal. It can be read by the first flexible circuit 209 in the DAS 240. In general, high energy X-rays 16 are transmitted through the first detector layer 202, the substrate 234, and the photodiode array 232 and enter the scintillator 230 of the second detector layer 204. The photons generated there travel through the scintillator 230 and collide with the photodiode array 232. The signal generated in the photodiode array 232 can be read by the second flexible circuit 236 to the DAS 240 and DAS IC 241 where the charge is converted into a digital signal, eg, via the connector 243. Sent to the logic circuit.

  During low flux operation, the x-rays 16 impinge on the direct conversion material 206 of the first detector layer 202. The first detector layer 202 generally does not saturate during low flux operation, and if a low count rate is obtained there, the first detector layer 202 is generally in photon counting mode to form an image. Can be used. At higher flux rates, the first detector layer 202 generally operates in the photon counting mode, the second detector layer 204 operates in the energy integration mode, and the first detector layer 202 has a higher counting rate. If acquired, the combination of x-rays detected in both layers is used to form the image. At very high flux rates, typically on the order of 5-100 Mcps, the first detector layer 202 is saturated and only the X-rays 16 detected on the second detector layer 204 form an image. Is generally used to operate in energy integration mode.

  FIG. 25 shows a flow diagram of a data acquisition pipeline 250 according to one embodiment of the invention. A pair of detector layers 252 and 254 convert the detected x-rays into a charge. Read electronics 256 reads the charge from the first detector layer 252 and read electronics 258 reads the charge from the second detector layer 254. The logic controller 262 receives the total signal level logic value 260 retrieved from the readout electronics 256, 258. The logic switch / controller 262 compares the total signal level logic value 260 to a threshold and reads data for image reconstruction and transfers it from only the electronic device 256 to the image chain 264 or transfers it only from the read electronic device 258. Or from both. It is contemplated that the logic switch / controller 262 can be composed of separate electronic components and can include implementation using FPGA, DSP, or ASIC circuitry. It is also contemplated that the functions of logic controller 262 can be performed by computer processor software, firmware, or other hardware, as is commonly understood in the art. The logic controller 262 transfers data to the image chain 264 from the read electronics 256 only, from the read electronics 258 only, or both based on the comparison of the total signal level logic value 260 and the threshold. The threshold is set for the flux rate at which the first detector layer 252 is saturated. Most commonly, logic signals are extracted from both layers 252, 254, but those skilled in the art can extract logic signals from only the first detector layer 252 or the second detector layer 254. Recognize that you can.

  If a low detection level occurs in the readout electronics 258 in a low flux state, the first detector layer 252 is generally not saturated and the image is energy discrimination information retrieved via the readout electronics 256. Can be formed completely and independently. In this case, the logic controller 262 preferably forwards information extracted from only the first detector layer 252 to the image chain 264. If a higher detection level occurs at readout electronics 258 at higher flux conditions, some saturation can occur at first detector layer 252 and the image is energy-discriminated first detector layer. 252 and information from both the second detector layer 254 operating in energy integration mode can be formed. In this case, the logic controller 262 preferably forwards information retrieved from both the first detector layer 252 and the second detector layer 254 to the image chain 264. In the highest flux state, when a very high detection level occurs in the readout electronics 258, the first detector layer 252 is generally fully saturated and the image is completely from the second detector layer 254. Can be formed. In this case, the logic controller 262 preferably forwards information extracted from only the second detector layer 254 to the image chain 264.

  26 is a cross-sectional view of the detector module of FIG. 24 taken along line 26-26. In this embodiment, as described above, detector module 200 includes a first detector layer 202 having a direct conversion material 206 to which a pixelated anode contact 208 is attached. A high voltage electrode 212 is attached to the direct conversion material 206 and a grid assembly 214 is attached to the high voltage electrode 212. The pixelated anode contact 208 is attached to and electrically coupled to the first flexible circuit 209, and the first flexible circuit 209 can be attached to the substrate 234. The second detector layer 204 includes a scintillator 230 attached to the photodiode array 232, which is attached to and electrically coupled to the second flexible circuit 236. A second flexible circuit 236 can also be attached to the substrate 234. The DAS 240 includes an IC 241 to read the first and second detector layers 202, 204, respectively.

  27 and 28 show an alternative detector module embodiment according to the present invention. The detector module 300 includes a first detector layer 301 having a direct conversion material 302 to which a pixelated anode contact 304 is attached. A high voltage electrode 306 is attached to the direct conversion material 302 and a grid assembly 308 is attached to the high voltage electrode 306. The pixelated anode contact 304 is electrically coupled to the first flexible circuit 310 and the IC 321 of the DAS 320. The detector module 300 also includes a second detector layer 303 having a scintillator 312 attached to the photodiode array 314. In addition to those shown in FIGS. 24 and 26, detector module 300 includes a multilayer substrate 316 that electrically couples photodiode array 314 to second flexible circuit 318 and IC 322 of DAS 320. The multilayer substrate 316 includes an electrical path through which the electrical signal extracted from the second detector layer 303 is transmitted to the second flexible circuit 318. The electrical signal generated by the first detector layer 301 is passed to the first flexible circuit 310, the DAS 320 to which the IC 321 is attached, and the connector 325.

  As shown in FIG. 27, the first and second circuits 310, 318 can extend from the opposite sides of the second detector layer 303 to the ICs 321, 322. Alternatively, as shown in FIG. 28, the first and second circuits 310, 318 can extend from the same side of the second detector layer 303 to the ICs 321, 322.

  FIG. 29 illustrates an alternative detector module embodiment according to the present invention. The detector module 350 includes a first detector layer 351 having a direct conversion material 352 to which a pixelated anode contact 354 is attached. A high voltage electrode 356 is attached to the direct conversion material 352 and a grid assembly 358 is attached to the high voltage electrode 356. Pixelated anode contact 354 is attached to and electrically coupled to flexible circuit 360 and IC 321 of DAS 320. The detector module 350 also includes a second detector layer 353 having a scintillator 362 attached to the photodiode array 364, which is attached to the multilayer substrate 366, and the multilayer substrate 366 includes the photodiode array 364. , Electrically coupled to flexible circuit 360 and DAS 370. The multilayer substrate 366 includes an electrical path through which the electrical signal extracted from the second detector layer 353 is transmitted to the flexible circuit 360. The electrical signal generated by the first detector layer 351 is passed to the flexible circuit 360 and the IC 371 of the DAS 370.

  FIG. 30 illustrates an alternative detector module embodiment according to the present invention. The detector module 400 includes a first detector layer 401 having a direct conversion material 402 to which a pixelated anode contact 404 is attached. A high voltage electrode 406 is attached to the direct conversion material 402 and a grid assembly 408 is attached to the high voltage electrode 406. The pixelated anode contact 404 is attached to and electrically coupled to the multilayer substrate 410. Multilayer substrate 410 includes an electrical path through which electrical signals extracted from first detector layer 401 are transmitted. Therefore, the electrical signal generated by the first detector layer 401 is passed to the flexible circuit 410 and the DAS 420 via the multilayer substrate 410. The detector module 400 also includes a second detector layer 403 having a scintillator 412 attached to the photodiode array 414, which is electrically coupled to the flexible circuit 416. The electrical signal generated by the second detector layer 403 is passed to the flexible circuit 416, the DAS 420 with the IC 421 attached thereto, and the connector 425.

  FIG. 31 shows an assembly procedure for assembling a detector module according to an embodiment of the present invention. In one embodiment, the assembly procedure corresponds to the assembly of the detector module 300 of FIG. In this embodiment, the DAS 320 has a first flexible circuit 310 and a component 309 of the first detector layer 301 attached thereto. The DAS 320 further includes a second flexible circuit 318 and a component 311 of the second detector layer 303 attached thereto. Components 309 of the first detector layer 301 include, but are not limited to, a direct conversion material 302, a pixelated anode contact 304, a high voltage electrode 306, and a grid assembly 308. Components 311 of the second detector layer 303 include, but are not limited to, a multilayer substrate 316, a photodiode array 314, and a scintillator 312. During assembly of the detector module 300, the DAS 320 ICs 321, 322 are attached to the flexible circuits 310, 318, and the flexible circuits 310, 318 are each attached to the interconnect board 323. The component 309 of the first detector layer 301 is attached to the first flexible circuit 310 and the component 311 of the second detector layer 303 is attached to the second flexible circuit 318. In one embodiment, the second detector layer 303 is convolved at location 330, followed by the convolution of the first detector layer 301 to location 332. The first and second flexible circuits 310, 318 are preferably glued together so that the first and second detector layers 301, 303 remain substantially superimposed. One skilled in the art will recognize that the other detector assemblies described, for example, in FIGS. 26 and 28-30 can be assembled in a similar manner.

  FIG. 32 illustrates an alternative detector module embodiment according to another embodiment of the present invention. The detector module 500 includes a first detector layer 502 having a direct conversion material 503 to which a pixelated anode contact 504 is attached. A high voltage electrode 506 is attached to the direct conversion material 503 and a grid assembly 508 is attached to the high voltage electrode 506. The pixelated anode contact 504 is attached to and electrically coupled to the first flexible circuit 510. The detector module 500 also includes a second detector layer 512 having a scintillator 514 attached to and electrically coupled to the photodiode array 516, the photodiode array 516 being extracted from the second detector layer 512. The multi-layer substrate 518 has an electrical path through which electrical signals are electrically coupled and transmitted to the second flexible circuit 520. The first flexible circuit 510 and the second flexible circuit 520 are also electrically coupled to each other. The IC 522 connected to the first flexible circuit 510 is not attached to the interconnect board 525 but is attached adjacent to the scintillator 514. The electrical signal can be processed by IC 522 or IC 524, and thus some of IC 522 is located closer to the signal source. Thus, the overall circuit capacitance can be reduced due to the close proximity of the IC 522 and the second detector layer 512.

  FIG. 33 illustrates an alternative detector module embodiment according to another embodiment of the present invention. Similar to the embodiment of FIG. 32, some of the ICs 562 are located closer to the power supply and similarly reduce the overall capacitance of the electrical circuit. The detector module 550 includes a first detector layer 552 having a direct conversion material 554 to which a pixelated anode contact 556 is attached. A high voltage electrode 558 is attached to the direct conversion material 554 and a grid assembly 560 is attached to the high voltage electrode 558. Pixelated anode contact 556 is attached to and electrically coupled to chip package 560 having IC 562 therein. The electrical signal generated by the first detector layer 552 is passed to the chip package 560, which provides a signal that is passed to the flexible circuit 572. Thus, the overall circuit capacitance can be reduced. Detector module 550 includes a second detector layer 564 having a scintillator 566 attached to and electrically coupled to photodiode array 568. The photodiode array 568 is attached to the multilayer substrate 570, and the multilayer substrate 570 is attached to the flexible circuit 572. The electrical signal generated by the second detector layer 564 is transferred to the multilayer substrate 570, and is similarly transferred to the flexible circuit 572 and transmitted to the IC 576. DAS 574 includes IC 562 and IC 576.

  34-36 illustrate alternative detector module embodiments according to the present invention. In these illustrated embodiments, the multilayer detector includes two layers of direct conversion material. As such, the direct conversion material can be configured to operate in energy discrimination or energy integration modes.

  FIG. 34 illustrates a detector module according to one embodiment of the present invention. The detector module 600 includes a first detector layer 601 having a direct conversion material 602 with a pixelated anode contact 604 attached thereto. A high voltage electrode 606 is attached to the direct conversion material 602 and a grid assembly 608 is attached to the high voltage electrode 606. Pixelated anode contact 604 is attached to and electrically coupled to first flexible circuit 610. The first flexible circuit 610 is attached to the substrate 612. The detector module 600 also includes a second detector layer 603 having a direct conversion material 614 to which the pixelated anode contact 616 is attached. A high voltage electrode 618 is attached to the direct conversion material 614. The pixelated anode contact 616 is attached to the second flexible circuit 620 and the second flexible circuit 620 is attached to the substrate 622. The electrical signal generated in the first detector layer is transmitted to the IC 624 of the DAS 626, and the electrical signal generated in the second detector layer 603 is transmitted to the IC 628 via the second flexible circuit 620. Is done.

  FIG. 35 illustrates an alternative detector module embodiment according to the present invention. The detector module 650 includes a first detector layer 651 having a direct conversion material 652 to which the pixelated anode contact 654 is attached. A high voltage electrode 656 is attached to the direct conversion material 652 and a grid assembly 658 is attached to the high voltage electrode 656. The pixelated anode contact 654 is attached to and electrically coupled to the multilayer substrate 655, and the multilayer substrate 655 is attached to and electrically coupled to the first flexible circuit 656. The detector module 650 also includes a second detector layer 653 having a direct conversion material 660 to which the pixelated anode contact 662 is attached. A high voltage electrode 664 is attached to the direct conversion material 660. The pixelated anode contact 662 is attached to and electrically coupled to the multilayer substrate 666, and the multilayer substrate 666 is attached to and electrically coupled to the second flexible circuit 668. The electrical signal generated in the first detector layer is transmitted to the IC 670 of the DAS 672, and the electrical signal generated in the second detector layer 653 is transmitted to the IC 674 of the DAS 672 via the second flexible circuit 668. Is transmitted.

  FIG. 36 illustrates an alternative detector module embodiment according to the present invention. The detector module 700 includes a first detector layer 701 having a direct conversion material 702 to which a pixelated anode contact 704 is attached. A high voltage electrode 706 is attached to the direct conversion material 702 and a grid assembly 708 is attached to the high voltage electrode 706. A pixelated anode contact 704 is attached to and electrically coupled to the first flexible circuit 710. The first flexible circuit 710 is attached to the substrate 712. The detector module 700 also includes a second detector layer 703 having a direct conversion material 714 to which the pixelated anode contact 716 is attached. A high voltage electrode 718 is attached to the direct conversion material 714. The pixelated anode contact 716 is attached to the second flexible circuit 720, and the second flexible circuit 720 is attached to the substrate 712. The electrical signal generated in the first detector layer is transmitted to the IC 724 of the DAS 726, and the electrical signal generated in the second detector layer 703 is transmitted to the IC 728 via the second flexible circuit 720. Is done.

  Referring now to FIG. 37, a parcel / baggage inspection system 890 that includes a detector as described herein is shown. System 890 includes a rotatable gantry 892 having an opening 894 through which a parcel or baggage can pass. A rotatable gantry 892 houses a radiation source 896 and a detector assembly 898. A conveyor system 800 is also provided, which includes a conveyor belt 802 supported by a structure 804 to allow the parcel or baggage 806 to be scanned automatically and continuously through the opening 894. The object 806 is fed through the opening 894 by the conveyor belt 802, after which imaging data is acquired, and the conveyor belt 802 delivers the parcel 806 out of the opening 894 in a controlled and continuous manner. As a result, postal inspectors, baggage handlers, and other security personnel can non-invasively inspect the contents of parcel 806 for explosives, knives, pistols, embargoes, and the like.

  According to one embodiment of the present invention, the CT detector includes a first detector configured to convert X-ray imaging energy into an electrical signal representing energy-sensitive X-ray imaging data; And a second detector configured to receive X-rays transmitted through the first detector, wherein the second detector is configured to convert the electrical signal into energy-sensitive radiographic data. A logic controller is electrically connected to the first detector and the second detector, receives a logic output signal from the second detector representing the amount of saturation level of the first detector, and outputs a logic output signal Is configured to output an electrical signal from the first detector, the second detector, or a combination thereof to the image chain based on the comparison.

  According to another embodiment, an x-ray imaging system includes a radiation source configured to project x-rays toward a scanned subject, a first detector layer, and a second detector layer. Wherein the second layer is arranged to receive x-rays from a radiation source transmitted through the first detector layer, each of the first and second layers comprising a radiation source At least one of a detector assembly and first and second detector layers configured to receive the projected x-rays from and convert the radiographic energy into an electrical signal representative of energy sensitive radiographic data Receiving data representing the flux rate at the image, determining whether an electrical signal has been output at the first detector layer for image reconstruction based on the received data, and image reconstruction based on the received data of In the second detector layer in order and a device configured to determine whether the output electric signal.

  According to another embodiment, a method of fabricating an imaging system includes providing an x-ray source, positioning a first detector to receive x-rays emitted from the x-ray source, Positioning a second detector to receive X-rays emitted from the X-ray source and transmitted through the first detector, and electrically coupling the logic device to the first and second detectors And a logic device configured to represent a level of x-ray flux at one of the first and second detectors.

  Although the invention has been described in terms of various embodiments thereof, equivalents, alternatives, and modifications are possible, apart from those explicitly stated, and are within the scope of the appended claims. Will be recognized.

It is a pictorial diagram of CT imaging system. FIG. 2 is a schematic block diagram of the system shown in FIG. 1. 1 is a perspective view of one embodiment of a CT system detector assembly. FIG. 1 is a perspective view of a CT detector that includes the present invention. FIG. FIG. 3 is a partial perspective view of a two-layer detector according to the present invention. FIG. 6 is a cross-sectional view of FIG. 5 taken along line 6-6 of FIG. FIG. 6 is a cross-sectional view of a direct conversion detector according to a further embodiment of the present invention. FIG. 6 is a cross-sectional view of a direct conversion detector according to a further embodiment of the present invention. FIG. 6 is a cross-sectional view of a direct conversion detector according to a further embodiment of the present invention. FIG. 6 is a cross-sectional view showing a signal feedthrough created in another embodiment of the present invention. 6 is a schematic cross-sectional view of a CT detector according to another embodiment of the present invention. FIG. FIG. 6 is a cross-sectional view of an alternative embodiment of a portion of a CT detector according to the present invention. FIG. 6 is a cross-sectional view of a portion of a CT detector according to yet another embodiment of the present invention. FIG. 6 is a cross-sectional view of yet another embodiment of a CT detector according to the present invention. FIG. 15 is a schematic cross-sectional view similar to FIGS. 13-14 showing an alternative embodiment of the present invention. FIG. 2 is a perspective view of a portion of a CT detector with components oriented in a vertical configuration. FIG. 6 is a schematic top view illustrating sub-pixelization of detector element regions according to the present invention. FIG. 6 is a top view of a single CT detector element region showing asymmetric subpixelation according to another embodiment of the present invention. FIG. 6 is a top view of alternative asymmetric subpixelization for a single CT detector element region according to an alternative embodiment of the present invention. FIG. 4 is a schematic block diagram illustrating a combination of each subpixel output of a given CT detector element region according to another embodiment of the present invention. FIG. 5 is a schematic circuit diagram illustrating flexible selection of sub-pixel outputs for a given CT detector element region according to another embodiment of the present invention. FIG. 6 is a schematic circuit diagram illustrating selection of a sub-pixel output for a given CT detector element region based on the saturation state of each sub-pixel according to a further embodiment of the present invention. FIG. 6 is a schematic circuit diagram illustrating selection of a sub-pixel output for a given CT detector element region based on the saturation state of each sub-pixel according to a further embodiment of the present invention. 1 is a perspective view of a two-layer detector module according to an embodiment of the present invention. FIG. FIG. 3 is a schematic block diagram of a data acquisition pipeline according to an embodiment of the present invention. FIG. 26 is a cross-sectional view of the detector module of FIG. 24 taken along line 26-26. FIG. 6 is a cross-sectional view of an alternative detector module embodiment according to an embodiment of the present invention. FIG. 6 is a cross-sectional view of an alternative detector module embodiment according to an embodiment of the present invention. FIG. 6 is a cross-sectional view of an alternative detector module embodiment according to an embodiment of the present invention. FIG. 6 is a cross-sectional view of an alternative detector module embodiment according to an embodiment of the present invention. It is the schematic which shows the assembly procedure of the detector module by one Embodiment of this invention. FIG. 6 is a cross-sectional view of an alternative detector module according to an embodiment of the invention. FIG. 6 is a cross-sectional view of an alternative detector module according to an embodiment of the present invention. FIG. 6 is a cross-sectional view of an alternative detector module according to an embodiment of the invention. FIG. 6 is a cross-sectional view of an alternative detector module according to an embodiment of the invention. FIG. 6 is a cross-sectional view of an alternative detector module according to an embodiment of the present invention. It is a pictorial diagram of a CT system used in a non-invasive package inspection system.

Explanation of symbols

DESCRIPTION OF SYMBOLS 10 Computer tomography (CT) imaging system 12 Gantry 14 X-ray source 16 X-ray beam 17 Rail 18 Detector assembly or collimator 19 Collimating blade or plate 20 Multiple detectors 20a Direct conversion detector 20b Detector 20c Detector 20d detector 22 patient 24 center of rotation 26 control mechanism 28 X-ray controller 30 gantry motor controller 32 data acquisition system 34 image reconstructor 36 computer 38 mass storage device 40 console 42 display 44 table motor controller 46 motorized table 48 gantry Opening 50 Detector element 51 Pack 52 Pin 53 Backlit diode array 54 Multilayer substrate 55 Spacer 56 Flex circuit 57 Surface 59 Multiple diodes 62 First semiconductor layer 64 Second semiconductor layer 65 Sensing element or contact 66 High voltage electrode 67 2D array 68 High voltage electrode 69 2D array 74 Semiconductor layer 76 Semiconductor layer 78 Single common signal collection layer or 2D contact Array 80 high voltage electrode 82 high voltage electrode 84 semiconductor layer 86 semiconductor layer 87 high voltage electrode 88 semiconductor layer 89 high voltage electrode 90 semiconductor layer 91 high voltage electrode 92 conductive line or path 93 collection contact array 94 conductive Line or path 95 Collection contact array 96 Semiconductor layer 98 Semiconductor layer 100 Semiconductor layer 102 Semiconductor layer 104 Conductive layer 106 Conductive layer 108 Conductive layer 110 Layer 112 layer 114 Semiconductor material 116 Collection contact array 118 High voltage electrode layer 120 Hole 122 Signal feed path 124 Collection contact 126 Detection 128 substrate 130 detector layer 132 detector layer 134 X-ray direction 136 flex layer 138 high voltage bias line 140 conversion component 142 conversion component 144 detector element array 146 detector element array 148 detector element 150 detector element 152 direct conversion layer 154 Flex Layer 156 Detector Element 158 Line 159 Line 160 Pixel Area 162 Subpixel 164 Dedicated Read Lead 165 Pixel Area 166 Data System Input 168 Output 170 Truth Table Circuit 172 Output 174 Input 176 Switch Network 180 Output 182 DAS
184 Signal shaper 186 Energy level discriminator 188 Counting register 200 Detector module 202 First detector layer 203 Spacer 204 Second detector layer 205 Hole 206 Direct conversion material 207 Hole 208 Pixelated anode contact 209 First Flexible interconnect 210 First surface 212 High voltage electrode 213 Second surface 214 X-ray attenuation grid assembly 215 High voltage lead 216 Bar or bar 218 Bar or bar 220 Peripheral frame 222 Opening 224 Outer edge or perimeter 226 Charge sharing region 230 Scintillator 232 Photodiode Array 234 Substrate 236 Second Flexible Interconnect 240 DAS
241 DAS integrated circuit 243 Connector 250 Data acquisition pipeline 252 Detector layer 254 Detector layer 256 Read electronics 258 Read electronics 260 Total signal level logic 262 Logic controller 264 Image chain 300 Detector module 301 First detector layer 302 Direct Conversion Material 303 Second Detector Layer 304 Pixelated Anode Contact 306 High Voltage Electrode 308 Grid Assembly 309 Component 310 First Flexible Circuit 311 Component 312 Scintillator 314 Photodiode Array 316 Multilayer Substrate 318 Second Flexible Circuit 320 DAS
321 IC
322 IC
323 interconnect board 325 connector 350 detector module 351 first detector layer 352 direct conversion material 353 second detector layer 354 pixelated anode contact 356 high voltage electrode 358 grid assembly 360 flexible circuit 362 scintillator 364 photodiode array 366 multilayer substrate 370 DAS
371 IC
400 detector module 401 first detector layer 402 direct conversion material 403 second detector layer 404 pixelated anode contact 406 high voltage electrode 408 grid assembly 410 multilayer substrate 412 scintillator 414 photodiode array 416 flexible circuit 420 DAS
421 IC
425 connector 500 detector module 502 first detector layer 503 direct conversion material 504 pixelated anode contact 506 high voltage electrode 508 grid assembly 510 first flexible circuit 512 second detector layer 514 scintillator 516 photodiode array 518 Multilayer substrate 520 Second flexible circuit 522 IC
524 IC
525 Interconnect Board 550 Detector Module 552 First Detector Layer 554 Direct Conversion Material 556 Pixelated Anode Contact 558 High Voltage Electrode 560 Grid Assembly 562 IC
564 Second detector layer 566 Scintillator 568 Photodiode array 570 Multilayer substrate 572 Flexible circuit 574 DAS
576 IC
600 detector module 601 first detector layer 602 direct conversion material 603 second detector layer 604 pixelated anode contact 606 high voltage electrode 608 grid assembly 610 first flexible circuit 612 substrate 614 direct conversion material 616 pixelated Anode contact portion 618 High voltage electrode 620 Second flexible circuit 622 Substrate 624 IC
626 DAS
628 IC
650 Detector module 651 First detector layer 652 Direct conversion material 653 Second detector layer 654 Pixelated anode contact 655 Multilayer substrate 656 High voltage electrode 658 Grid assembly 660 Direct conversion material 662 Pixelated anode contact 664 High Voltage electrode 666 Multi-layer substrate 668 Second flexible circuit 670 IC
672 DAS
674 IC
700 detector module 701 first detector layer 702 direct conversion material 703 second detector layer 704 pixelated anode contact 706 high voltage electrode 708 grid assembly 710 first flexible circuit 712 substrate 714 direct conversion material 716 pixelated Anode contact portion 718 High voltage electrode 720 Second flexible circuit 724 IC
726 DAS
728 IC
800 Conveyor System 802 Conveyor Belt 804 Structure 806 Baggage 890 Inspection System 892 Rotatable Gantry 894 Opening 896 Radiation Source 898 Detector Assembly

Claims (10)

A first detector (202, 252) configured to convert radiographic energy into an electrical signal representing energy sensitive radiographic data;
A second detection configured to convert radiographic energy into an electrical signal representative of energy sensitive radiographic data and arranged to receive X-rays transmitted through the first detector (202, 252). Vessels (204, 254),
A logic controller (262) electrically connected to the first detector (202, 252) and the second detector (204, 254);
Receiving a logic output signal (260) from the second detector (204, 254) representing the amount of saturation level of the first detector (202, 252);
Comparing the logic output signal (260) to a threshold;
Based on the comparison, configured to output an electrical signal from the first detector (202, 252), the second detector (204, 254), or a combination thereof to an image chain A computed tomography (CT) detector (200) comprising a logic controller (262). The CT detector (200) of claim 1, wherein the logic controller (262) is one of an FPGA circuit, a DSP circuit, an ASIC circuit, software implemented on a computer readable storage medium, and firmware. The CT detector (200) of claim 1, wherein the first detector (202, 252) comprises an energy discrimination detector having a direct conversion layer. At least one of a charge sharing region (226) and an end trapping region (224) formed in the direct conversion layer is further provided, toward the at least one of the charge sharing region (226) and the end trapping region (224). The CT detector (200) of claim 3, further comprising an x-ray attenuating material arranged to attenuate x-rays transmitted in the direction. The CT detector (200) of claim 3, wherein the energy discrimination detector is configured to operate in a photon counting mode. The CT detector (200) of claim 1, wherein the second detector comprises an energy discrimination detector configured to operate in one of a photon counting mode and an energy integration mode. The CT detector (200) of claim 1, wherein the second detector (204, 254) comprises a scintillating layer optically coupled to a photodiode array. If the logic output signal indicates that no saturation has occurred in the first detector (202, 252), the logic controller (262) is in the first detector (202, 252). The CT detector (200) of claim 1, further configured to output only a converted electrical signal. When the logic output signal indicates that the first detector (202, 252) is saturated, the logic controller (262) is only in the second detector (204, 254). The CT detector (200) of claim 1, further configured to output a converted electrical signal. The CT detector (200) of claim 1, wherein the logic output signal represents a rate of flux incident on the first detector.

Images