Pixel design for radiation detectors

Abstract

The present method involves the use of reference pixels provided between primary pixels of the detector panel. Coincidence circuitry or logic can be employed so that the measurement signal resulting from the same X-ray event can be correctly aligned at the appropriate location on the detector panel. That is, the signals measured at both the reference and primary pixels can be combined to provide an accurate estimate of the measurement signal.

Description

[0001] Statement Regarding Federally Funded Research and Development

[0002] This invention was made with government support under Contract No. 1UO1EB017140-01A1 awarded by the National Institutes of Health. The government has certain rights in this invention. Background Art

[0003] The subject matter disclosed herein generally relates to the use of semiconductor-based radiation detectors.

[0004] Non-invasive imaging techniques allow for non-invasively obtaining images of internal structures or features of a subject (a patient, manufactured goods, luggage, packages, or passengers). Specifically, such non-invasive imaging techniques rely on various physical principles (such as the differential transmission of X-rays through a target volume or the reflection of sound waves) to acquire data and construct an image or otherwise represent the internal features of the subject.

[0005] For example, in X-ray-based imaging techniques, X-ray radiation passes across a subject of interest (such as a human patient), and a portion of the radiation impacts a detector that collects intensity data. In digital X-ray systems, the detector generates a signal representing the amount or intensity of the radiation that impacted discrete pixel areas on the detector surface. The signal can then be processed to generate an image that can be displayed for viewing.

[0006] In one such X-ray based technique, known as computed tomography (CT), a scanner may project a fan-shaped or cone-shaped beam of X-rays from an X-ray source at multiple view angles around an object being imaged, such as a patient. The X-ray beam is attenuated as it passes through the object and is detected by a set of detector elements that produce a signal representing the intensity of the incident X-ray intensity on the detectors. The signals are processed to produce data representing the line integral of the linear attenuation coefficient of the object along the X-ray path. These signals are often referred to as "projection data" or just "projections". By using reconstruction techniques such as filtered back projection, a volume or volume rendered image representing a region of interest of the patient or imaged object can be generated. In a medical context, pathological or other structures of interest can then be located or identified from the reconstructed image or rendered volume.

[0007] The radiation detectors used in these types of imaging techniques can operate in energy integrating mode (i.e., readout of the total integrated energy deposited during the acquisition interval) or in photon counting mode (detecting and counting each individual X-ray photon). Energy integration is the conventional mode of X-ray detectors in most clinical applications. However, photon counting detectors offer other benefits over energy integrating detectors, such as improved resolution, the ability to improve contrast-to-noise ratio by optimally weighting the detected photons, the ability to better delineate materials in the X-ray beam, etc. Photon counting detectors may also provide energy discrimination functionality, depending on their implementation, so that each detected photon can be characterized or "binned" based on its observed energy.

[0008] Radiation detectors typically operate based on two different physical principles. Some detectors employ a scintillation intermediate that, in response to an X-ray event, emits visible photons at the location of the X-ray event. The visible photons can then be detected and localized using known photodetection techniques. Alternatively, detectors can employ direct conversion of incident X-rays to electrical signals, such as detectors based on silicon strips or other semiconductor materials such as cadmium zinc telluride (CZT) or cadmium telluride (CdTe), which generate a measurable signal when the semiconductor substrate itself is exposed to X-ray photons. However, one problem that can arise in such direct conversion situations is that the charge cloud associated with the conversion event can span multiple detector pixels, which can lead to erroneous measurements. Summary of the Invention

[0009] In one embodiment, a radiation detector assembly is provided. According to this embodiment, the radiation detector assembly includes a primary pixel array and a reference pixel array. Each reference pixel is adjacent to at least one primary pixel. The radiation detector assembly also includes detection circuitry in communication with the primary pixel array and the reference pixel array. The detection circuitry is configured to detect signals in the primary pixel and the reference pixel that are temporally aligned.

[0010] In another embodiment, a method for detecting and correcting charge sharing events in a radiation detector is provided. According to this embodiment, the method includes reading out signals from a primary pixel array and a reference pixel array. Each reference pixel is adjacent to at least one primary pixel. The signal read out from the corresponding reference pixel is detected. Coincidence logic is performed to identify a corresponding primary pixel adjacent to the corresponding reference pixel from which the corresponding signal was read out. The signal from the corresponding reference pixel and the corresponding signal are combined. BRIEF DESCRIPTION OF THE DRAWINGS

[0011] These and other features, aspects, and advantages of the present invention will be better understood when the following detailed description is read with reference to the accompanying drawings, in which like characters represent like parts throughout, and in which:

[0012] Figure 1 is a schematic diagram of an embodiment of a computed tomography (CT) system configured to acquire CT images of a patient and process the images according to aspects of the present disclosure;

[0013] Figure 2 depicts a side view of a conventional pixel arrangement;

[0014] Figure 3 depicts a side view of a reference pixel and a primary pixel in a background of overlapping charge clouds according to aspects of the present disclosure;

[0015] Figure 4 depicts a side view of a reference pixel and a primary pixel in the context of a charge cloud centered about the reference pixel, in accordance with aspects of the present disclosure;

[0016] Figure 5 depicts a top view of a two-dimensional arrangement of reference pixels and primary pixels according to aspects of the present disclosure;

[0017] Figure 6 depicts a top view of another two-dimensional arrangement of reference pixels and primary pixels according to aspects of the present disclosure;

[0018] Figure 7 depicts a top view of additional two-dimensional arrangements of reference pixels and primary pixels according to aspects of the present disclosure;

[0019] Figure 8 depicts a side view of a reference pixel and a primary pixel with an anti-scatter collimator according to aspects of the present disclosure;

[0020] Figure 9 depicts a top view of a pixel of an anti-scatter collimator and a detector according to aspects of the present disclosure;

[0021] Figure 10 depicts a top view of an alternative arrangement of anti-scatter collimators and pixels of a detector according to aspects of the present disclosure;

[0022] Figure 11 depicts a top view of another arrangement of anti-scatter collimators and pixels of a detector according to aspects of the present disclosure;

[0023] Figure 12 An example of a conventional detection circuit is depicted;

[0024] Figure 13 depicts examples of circuits configured to detect and correct charge sharing in accordance with aspects of the present disclosure;

[0025] Figure 14 depicts a schematic diagram of a primary pixel and a reference pixel incorporating a charge cloud according to aspects of the present disclosure; and

[0026] Figure 15 Depicts the charge-to-volume ratio (V1 / V tot ) versus linear overlap length (h / R). DETAILED DESCRIPTION

[0027] One or more specific embodiments will be described below. In order to provide a concise description of these embodiments, not all features of an actual implementation may be described in the specification. It should be understood that in the development of any such actual implementation, as in any engineering or design project, many implementation-specific decisions must be made to achieve the developer's specific goals, such as complying with system-related and business-related constraints that may vary from implementation to implementation. Furthermore, it should be understood that such development efforts may be complex and time-consuming, but remain a routine task of design, fabrication, and manufacturing for those of ordinary skill having the benefit of this disclosure.

[0028] When introducing elements of various embodiments of the present invention, the articles "a," "an," "the," and "said" are intended to mean that there are one or more of the elements. The terms "comprising," "including," and "having" are intended to be inclusive and mean that there may be additional elements other than the listed elements. Furthermore, any numerical examples in the following discussion are intended to be non-limiting, and thus the appended numerical values, ranges, and percentages are within the scope of the disclosed embodiments.

[0029] While the following discussion is generally provided in the context of medical imaging, it should be understood that the present technology is not limited to such medical contexts. Indeed, the examples and explanations are provided in this medical context solely to facilitate explanation by providing examples of practical implementations and applications. However, the present method may also be used in other contexts, such as non-destructive inspection of manufactured parts or goods (i.e., quality control or quality review applications), and / or non-invasive inspection of packages, boxes, luggage, etc. (i.e., security or screening applications). In general, the present method may be desirable in any imaging or screening context where energy discrimination in a photon counting context is desired.

[0030] As discussed herein, the present method involves the use of reference or "shared" pixels provided between the primary (i.e., active or readout) pixels of a semiconductor (i.e., direct conversion) type detector panel (such as detectors based on CZT, CdTe, or silicon). Specifically, the size of the reference pixel can be designed to prevent the typical charge cloud from overlapping the two primary pixels. Coincidence circuitry or logic can also be employed so that the measurement signals generated by the same X-ray event can be correctly correlated, so that the signals measured at both the reference pixel and the primary pixel can be combined (e.g., summed) to provide an accurate measurement signal at the appropriate location on the detector panel. In this way, spectral correction for charge sharing events can be achieved.

[0031] Furthermore, as discussed herein, in certain embodiments, the measurement signal of a shared or reference pixel can be used to localize the measurement signal to a location smaller than a full primary pixel, i.e., to sub-pixel resolution. For example, based on a detection event at a particular reference pixel, an X-ray event at an adjacent primary pixel can be localized to a spatial subregion of the primary pixel. In this way, sub-pixel spatial resolution can be achieved.

[0032] In certain embodiments employing an anti-scatter collimator, this approach may be useful because reference pixels can be positioned generally in the region beneath the anti-scatter collimator, which is typically shielded from X-ray photon events. In this manner, the reference pixels can be used to obtain additional spatial and energy information attributable to charge-sharing events. Thus, the region beneath the anti-scatter collimator can be used to obtain useful signals, rather than dead zones on the detector panel.

[0033] Considering the previous discussion, Figure 1 An embodiment of an imaging system 10 for acquiring and processing image data using reference (i.e., "shared") pixels positioned between primary (i.e., readout or active) pixels according to the structures and methods discussed herein is shown. In the illustrated embodiment, the system 10 is a computed tomography (CT) system designed to acquire X-ray projection data and reconstruct the projection data into volumetric reconstructions for display and analysis. The CT imaging system 10 includes one or more X-ray sources 12, such as one or more X-ray tubes or solid-state emission structures, that allow X-rays to be generated at one or more energy spectra during an imaging session.

[0034] In some embodiments, the source 12 can be positioned near a filter assembly or beam shaper 22, which can be used to steer the X-ray beam 20, define the shape and / or extent of a high-intensity region of the X-ray beam 20, control or define the energy distribution of the X-ray beam 20, and / or otherwise limit X-ray exposure to those portions of the patient 24 that are not within the region of interest. In practice, the filter assembly or beam shaper 22 can be incorporated within the gantry, between the source 12 and the imaging volume.

[0035] An X-ray beam 20 enters an area in which a subject (e.g., a patient 24) or an object of interest (e.g., a manufactured part, luggage, a package, etc.) is located. The subject attenuates at least a portion of the X-rays 20, resulting in attenuated X-rays 26 that impinge on a detector array 28 formed from a plurality of detector elements (e.g., pixels) as discussed herein. The detector 28 may be an energy integrating detector, a photon counting detector, an energy discriminating detector, or any other suitable radiation detector. As an example, the detector 28 may be an energy discriminating photon counting detector whose output conveys information about the number and energy of photons that impacted the detector at a measurement location and within a time interval corresponding to a scan or imaging phase. In one embodiment, the detector 28 may be a direct conversion detector (i.e., a detector that does not employ a scintillator intermediate), such as a detector based on CZT, CdTe, or silicon semiconductor materials, which generates a measurable signal when the semiconductor substrate itself is exposed to X-ray photons.

[0036] In the depicted example, detector 28 is adjacent to or otherwise close to anti-scatter collimator 18, which is typically made of a material that absorbs or otherwise blocks X-ray photons. Consequently, X-rays that strike the membrane of anti-scatter collimator 18, such as X-rays that are reflected, deflected, or otherwise moved at an angle relative to the membrane of anti-scatter collimator 18, are prevented from reaching detector 28. In contrast, X-ray photons traveling along a relatively straight path from X-ray source 12 are unimpeded by anti-scatter collimator 18 and reach detector 28. For the purposes of this discussion, detector 28 may in some cases be referred to as being below or in the "shadow" of anti-scatter collimator 18 from the perspective of X-ray source 12. This characterization indicates that collimator 18 is always between X-ray source 12 and detector 28, and that X-ray source 12 is considered "upward" in relation to the X-ray illumination source (regardless of its actual orientation or position). However, it will be understood that this characterization does not necessarily represent absolute position or orientation information.

[0037] With respect to the detector 28, as discussed herein, the detector 28 generally defines an array of detector elements, each of which generates an electrical signal when exposed to an X-ray photon. The electrical signals are acquired and processed to generate one or more projection data sets. In the depicted example, the detector 28 is coupled to a system controller 30, which commands the acquisition of the digital signals generated by the detector 28.

[0038] The system controller 30 commands the operation of the imaging system 10 to perform filtering, inspection, and / or calibration schemes and can process the acquired data. The system controller 30 provides power, focus position, control signals, etc. for the X-ray inspection sequence relative to the X-ray source 12. According to certain embodiments, the system controller 30 can control the operation of the filter assembly 22, the CT gantry (or other structural support to which the X-ray source 12 and detector 28 are attached), and / or the translation and / or tilting of the patient support during the inspection.

[0039] In addition, via the motor controller 36, the system controller 30 can control the operation of the linear positioning subsystem 32 and / or the rotation subsystem 34 for moving the subject 24 and / or components of the imaging system 10, respectively. The system controller 30 may include signal processing circuitry and associated memory circuitry. In such embodiments, the memory circuitry may store programs, routines, and / or coded algorithms that are executed by the system controller 30 to operate the imaging system 10 (including the X-ray source 12 and / or the filter assembly 22) and process the digital measurements acquired by the detector 28 according to the steps and procedures discussed herein. In one embodiment, the system controller 30 may be implemented as all or part of a processor-based system.

[0040] The source 12 may be controlled by an X-ray controller 38 included within the system controller 30. The X-ray controller 38 may be configured to provide power, timing signals, and / or focal spot size and spot position to the source 12. Furthermore, in some embodiments, the X-ray controller 38 may be configured to selectively activate the source 12 so that tubes or emitters at different locations within the system 10 may operate synchronously with or independently of each other, or to switch the source between different energy distributions during an imaging session.

[0041] The system controller 30 may include a data acquisition system (DAS) 40. The DAS 40 receives data collected by the readout electronics of the detector 28, such as digital signals from the detector 28. The DAS 40 may then convert and / or process the data for subsequent processing by a processor-based system, such as a computer 42. In certain embodiments discussed herein, circuitry within the detector 28 may convert the detector's analog signals into digital signals before transmission to the data acquisition system 40. The computer 42 may include or be in communication with one or more non-transitory memory devices 46 that may store data processed by the computer 42, data to be processed by the computer 42, or instructions to be executed by the image processing circuitry 44 of the computer 42. For example, a processor of the computer 42 may execute one or more sets of instructions stored on a memory 46, which may be a memory of the computer 42, a memory of a processor, firmware, or the like.

[0042] The computer 42 may also be adapted to control features enabled by the system controller 30 (i.e., scanning operations and data acquisition), such as in response to commands and scanning parameters provided by an operator via an operator workstation 48. The system 10 may also include a display 50 coupled to the operator workstation 48 that allows the operator to view relevant system data, imaging parameters, raw imaging data, reconstruction data (e.g., soft tissue images, bone images, segmented vascular trees, etc.), material base images, and / or material decomposition results, etc. Additionally, the system 10 may include a printer 52 coupled to the operator workstation 48 and configured to print any desired measurements. The display 50 and printer 52 may also be connected to the computer 42 (e.g., directly or via the operator workstation 48). Figure 1 ). In addition, the operator workstation 48 may include or be coupled to a picture archiving and communication system (PACS) 54. The PACS 54 may be coupled to a remote system or client 56, a radiology information system (RIS), a hospital information system (HIS), or an internal or external network so that others at different locations can access the image data.

[0043] With reference to the preceding discussion of the entire imaging system 10, the present method involves the use of reference pixels, also referred to herein as shared pixels, located between the primary or readout pixels of the detector 28. To discuss the present method, it may be useful to initially describe the occurrence of charge sharing in a conventional solid-state detector. As an example, turning to Figure 2 , a conventional detector panel 28B is depicted having primary pixel electrodes 70 that collect charge signals generated in response to X-ray photons 72 (shown as traveling toward detector 28B along dashed arrows) absorbed in semiconductor sensor material 90. Each X-ray interaction with sensor material 90 may be associated with a charge cloud 74 that originates from the corresponding X-ray photon 72 interaction.

[0044] Under the influence of the electric field, a charge cloud is transported from the photon absorption site in the bulk of sensor material 90 to electrodes 70 on the sensor's surface. These electrodes define the boundaries of each individual pixel. The proportion of charge ultimately collected by each electrode (pixel) depends on the amount of charge that overlaps with each electrode upon arrival. This allows the charge cloud to be split or shared among multiple pixels 70. In this example, the rightmost X-ray photon will interact with pixel 70B, far from the edge, so the resulting signal will be recorded only in pixel 70B. Conversely, the leftmost X-ray photon 72 will interact with pixel 70B along the edge, and the resulting charge cloud 74 will overlap between pixels 70A and 70B. This second scenario results in what is known as charge sharing. In pulse height analysis of the collected energy spectrum, charge sharing creates the appearance of two separate X-rays, the sum of whose energies is approximately equal to the actual X-ray energy. Consequently, a single event is mistakenly perceived as two separate events, each with a fraction of the energy of a true single event if correctly recorded.

[0045] Now go to Figure 3 and Figure 4 , in accordance with aspects of the present method, a reference pixel 80 is positioned between each pair of primary pixels 70. In this example, the reference pixel 80 has a relatively small area on the surface of the detector 28. According to one embodiment, the size of the reference pixel 80 is designed to approximate the size of the charge cloud 74. That is, in this example, the width of the reference pixel 80 may correspond to or slightly exceed a known or average diameter (e.g., approximately 50 μm) of the charge cloud 74 generated in a given imaging (e.g., X-ray) application. As a result, except in rare cases where an X-ray interacts with the center of the reference pixel 80 (e.g., Figure 4 As shown), charge sharing events in a given dimension will typically be (as Figure 3 shown) between the reference pixel 80 and the adjacent primary pixel 70, rather than between two adjacent primary pixels (e.g., pixels 70A and 70B).

[0046] Further explanation of the arrangement of reference pixels 80 and primary pixels 70 may benefit from a top view. As an example, and turning to Figure 5 , shows one possible arrangement of reference pixels 80 and primary pixels 70 in a two-dimensional array. In this example, each reference pixel 80 has exactly two primary pixels 70 as neighboring pixels.

[0047] Go to Figure 6 and Figure 7 , in these examples, the reference pixel 80 is shaped and / or sized to maximize the overall fill factor. As an example, Figure 6An example is depicted where the reference pixel 80 in one dimension is sized to extend into and substantially fill the intersection region between pixels. Figure 7 , an example is provided in which the reference pixels 80 are shaped such that each reference pixel extends partially into the intersection region such that, in combination, the reference pixels substantially fill the intersection region.

[0048] Go to Figure 8 , provides an example of an embodiment provided in conjunction with an anti-scatter collimator 18. In this configuration, because the collimator 18 blocks X-ray photons, thereby preventing primary X-ray events in the shielded area of ​​the detector, the area of ​​the detector 28 below the anti-scatter collimator 18 generally does not generate a useful signal. As a result, in conventional approaches, the area of ​​the detector 28 below the anti-scatter collimator may be ineffective as a dead zone on the detector panel 28.

[0049] However, in certain embodiments of the present method, the reference pixel 80 is typically located below the diaphragm 86 of the anti-scatter collimator 18. Depending on the embodiment, the width of the reference pixel 80 can be equal to or slightly larger than the corresponding width of the corresponding diaphragm 18. However, in general, the reference pixel 80 is sized to prevent charge sharing events from occurring between two adjacent primary pixels 70 and, therefore, typically has a width that corresponds to at least the diameter of an expected charge sharing event. That is, the reference pixel 80 is sized so that when an X-ray photon is detected, all or most of the charge sharing events are shared by the pair of primary and reference pixels 70, rather than the pair of primary pixels. As an example, in one embodiment, the width of the primary pixel 70 (as viewed from the side view depicted) can be approximately 550 μm, while the width of the reference pixel 80 (as viewed from the same view) can be approximately 160 μm. In another example, the reference pixel 80 can have a width that is less than or equal to half the width of the primary pixel, such that the count rate in the reference pixel is significantly less than the count rate in the primary pixel when exposed to the same X-ray flux.

[0050] In the depicted example, X-rays interact with semiconductor sensors 90 (e.g., silicon, CZT, or CdTe) of detector 28, generating charge clouds 74. Vertical lines are provided relative to each charge cloud 74 to better visualize the relationship of the respective charge cloud 74 to reference pixel 80 and primary pixel 70. As shown in this example, although reference pixel 80 is positioned below membrane 86 of anti-scatter collimator 18 and is therefore shielded from primary X-ray interaction events, reference pixel 80 is still exposed to charge sharing events 74.

[0051] By illustrating this concept, Figure 8A first charge sharing event 74A is shown generating signals only at primary pixel 70A and reference pixel 80. In contrast, a second sharing event 74B generates signals only at primary pixel 70B and reference pixel 80. Primary X-ray interaction is not possible at reference pixel 80 due to the presence of corresponding diaphragm 86, but charge sharing event 74 can be properly attributed to primary pixel 70 where the X-ray event occurred.

[0052] It will be appreciated that charge sharing correction as discussed herein relies not only on obtaining a measurement signal from each pixel where shared charge exists, but also on being able to identify that two such signals result from the same (i.e., single) X-ray interaction with the detector sensor 90. That is, separate measurement signals on adjacent reference pixel 80 and primary pixel 70 will only indicate a charge sharing event if they are determined to occur sufficiently close in time to one another (i.e., coincide in time). Coincidence between two reference pixels will generally not be detected.

[0053] With this in mind, one aspect of the present method is to use coincidence logic to identify charge sharing events. Such logic can be implemented as part of an application specific integrated circuit (ASIC), or as a routine executable by a general purpose processor or circuit, where the routine or ASIC is provided in the data readout path of the detector 28. For example, such an ASIC or routine can be provided as part of the detector 28 (such as part of the readout circuitry provided on the detector itself), or provided downstream of the detector on a DAS or other detector control or readout component.

[0054] As an example, such coincidence logic (whether implemented as an ASIC or executable code) can be triggered when a signal is read out from a reference pixel 80. It will be appreciated that because the reference pixel 80 has a relatively small surface area compared to the primary pixel 70, and in some cases, because the reference pixel 80 is at least partially positioned below the membrane of the anti-scatter collimator 18, the reference pixel 80 will generate fewer counts than the larger, fully exposed primary pixel 70. Thus, when an event is detected at the reference pixel 80 (as indicated by the signal read out from the reference pixel), the coincidence logic can be triggered and, when executed, will detect which adjacent primary pixel 70 detected a corresponding event close enough in time to be attributed to the same X-ray interaction. The charge sharing event can then be attributed to that primary pixel 70, and the appropriate signal added.

[0055] exist Figures 9 to 11Three high-level examples of conforming resolution are shown in FIG. Each example depicts a top view of an anti-scatter collimator with X-ray blocking membranes 86 positioned above the detector. In these examples, reference pixels 80 are positioned below membranes 86 and are therefore not visible. In contrast, primary pixels 70, positioned between membranes 70, can be seen in the top view. Figure 9 A two-dimensional (2D) anti-scatter collimator (i.e., grid) is depicted; Figure 10 A one-dimensional (1D) anti-scatter collimator is depicted; and Figure 11 A 2D anti-scatter collimator is depicted in which the spaces between the diaphragms 86 are positioned at a 2x2 pixel array pitch, rather than individual pixels.

[0056] Go to Figure 9 In this example, the signal detected at a given reference pixel will be resolved to (e.g., attributed to) one of the two adjacent primary pixels 70A or 70B based on coincidence logic, depending on which of these primary pixels also detected the event at the same time (or sufficiently close in time) as the corresponding reference pixel, as shown by arrow 92. Figure 9 Only a single coincidence determination in one dimension is shown, but there are also reference pixels extending in the vertical dimension to resolve charge sharing events in the other dimension.

[0057] Similarly, go to Figure 10 In the 1D anti-scatter collimator example, a signal detected at a given reference pixel in one dimension will be resolved to one of the two adjacent primary pixels 70A or 70B based on coincidence logic, depending on which of these primary pixels also detects the event at the same time (or sufficiently close in time) as the corresponding reference pixel, as shown by arrow 92. However, unlike Figure 9 Unlike the example of , there is no diaphragm and reference pixel in the second dimension to prevent the charge sharing event from being resolved in the second dimension. That is, each primary pixel 70 is adjacent to the reference pixel 80 on only two sides. Figure 9 Compared to the illustrated embodiment, the accuracy of the correction is reduced because charge sharing events may still occur between primary pixels 70. However, there is still a benefit in being able to correct for those charge sharing events that occur near a reference pixel.

[0058] Go to Figure 11 In this example, the anti-scatter collimator membrane 86 is not positioned around each individual pixel, but rather around the 2×2 array of primary pixels 70. In this embodiment, as in Figure 10 In the 1D collimator example, each primary pixel 70 is adjacent to a reference pixel 80 on only two sides. Therefore, as in the previous example, Figure 9The accuracy of the correction is reduced compared to the illustrated embodiment because, in each 2x2 pixel array, charge sharing events may still occur between primary pixels 70. However, there is still a benefit in being able to correct for those charge sharing events that occur near a reference pixel.

[0059] With regard to the logical and electrical implementation of coincidence detection, it may be useful to initially understand the operation of conventional photon counting, energy discriminating detection circuits. Figure 12 An example of this circuit is shown in Figure 1. In this example, the charge from each pixel enters a charge-sensitive amplifier, a shaping circuit, and sometimes a sample-and-hold circuit, which are collectively shown as U1. The output of U1 is a voltage proportional to the amplitude of the collected charge. This voltage is routed to one or more threshold comparators (here shown as Th1 to Th3), each of which compares the input voltage to a reference voltage (V E1 、V E2 and V E3 ). If this voltage exceeds the reference voltage, the digital output is activated, which causes the corresponding register of the counter to be incremented by 1 count. In this way, the incoming charge pulses are counted according to their energy.

[0060] on the contrary, Figure 13 A circuit capable of detecting and correcting charge sharing using signals from a reference pixel 80 is depicted, as discussed herein. In the illustrated example, electronic channels (channels 1 to 3) are assigned to pixels 70A, 80B, and 70B, respectively. There are two types of channels, one for the main (i.e., regular) pixels 70 (here, channels 1 and 3), and a different type of channel for the reference pixel 80 (here, channel 2). In this example, all channels are equipped with input amplifiers / shapers (U1, U2, U3) and a set of energy thresholds (Th4 to Th12) and associated output counters.

[0061] In the depicted example, the channel associated with reference pixel 80 (i.e., channel 2) includes three trigger comparators connected to the reference pixel amplifier and its neighboring pixel amplifiers (Th1 to Th3). The reference voltage of the trigger threshold is set to a low value (V TRIG ), so that any signal above the noise floor is sufficient to trigger the comparator.

[0062] In the depicted embodiment, the trigger threshold logic outputs are connected to switches S1 and S2 via AND gates (A1 and A2). If both the primary pixel 70A and the reference pixel 80 trigger, then switch S1 is set to connect the output of delay D1 along with a delayed version of the reference pixel 80 signal to a summing circuit (SUM). In this case, the energy threshold for channel 1 is never triggered. X-ray events are reported only through channel 2. This switch is set to Figure 13 Shown in.

[0063] If channel 1 is triggered and channel 2 is not, then S1 connects D1 to the energy comparator and counter of pixel 1. The X-ray event is reported through channel 1. If both channels 2 and 3 are triggered, the corresponding process occurs.

[0064] Delay circuits D1 and D3 ensure that switches S1 or S2 are set to their desired states before the analog pulse arrives from U1 or U3. Delay D2 ensures that the pulse from U2 arrives at SUM at the same time.

[0065] While the previous examples show a one-dimensional pixel array, extension to a two-dimensional pixel array is straightforward.

[0066] In addition to charge sharing resolution, the use of reference pixels 80 as discussed herein can also allow for sub-pixel spatial resolution. Specifically, the detection of charge sharing events 74 discussed herein can be used to localize the event on a primary pixel. Consequently, this sub-pixel spatial information can be used by reconstruction algorithms to improve imaging resolution.

[0067] As an example, the ratio of the charges collected by the reference pixel 80 and the corresponding primary pixel 70 can be used to estimate the center position of the charge cloud 74. Figure 14 As shown, it would be useful to estimate the distance Δ×2 from the center of the primary pixel 70 (having a pitch p) to the center of the charge cloud 74 .

[0068] In this example, the total charge q is proportional to the energy of the X-ray. Let q1 and q2 be the charges collected in the reference pixel 80 and the main pixel 70, respectively. Then

[0069] (5)E tot ∝q=q1+q2。

[0070] The radius R of the charge cloud 74 is a function of the X-ray energy, or equivalently, the total charge:

[0071] (6)R=f(q).

[0072] Let V tot Let V1 be the volume of the charge cloud 74 and let V2 be the volume that overlaps with the reference pixel 80. Then the ratio of the overlap to the total volume will be

[0073] (7)

[0074] Where h is the linear length of the overlap. The ratio of the charge overlap with the reference pixel and the primary pixel, denoted as V1 and V2 respectively, can be written as:

[0075] (8)

[0076] Go to Figure 15 , showing the charge-to-volume ratio (V1 / V tot ) versus the linear overlap length (h / R). Since this ratio is single-valued, measurement of the ratio uniquely determines h. Since R and h are known, the distance from the center of the primary pixel 70 can be calculated as:

[0077] (9)

[0078] The technical effects of the present invention include the use of reference pixels provided between the primary pixels of a semiconductor (i.e. direct conversion) type detector panel. The size of the reference pixel can be designed to prevent a typical charge cloud from overlapping two primary pixels. Coincidence circuitry or logic can also be employed so that the measurement signals generated by the same X-ray event can be correctly correlated at appropriate locations on the detector panel (i.e. the signals measured at both the reference pixel and the primary pixel can be combined to provide an accurate estimate of the measurement signal). In this way, spectral correction of charge sharing events can be achieved. Furthermore, the measurement signal of the reference pixel can be used to localize the measurement signal to a location that is smaller than a full primary pixel, i.e. sub-pixel resolution.

[0079] This written description uses examples to disclose the invention, including the best mode, and to enable any person skilled in the art to practice the invention, including making and using any devices or systems and performing any included methods. The patentable scope of the invention is defined by the claims and may include other examples that occur to those skilled in the art. Such other examples are intended to be within the scope of the claims if they have structural elements that do not differ from the literal language of the claims, or if they include equivalent structural elements with insubstantial differences from the literal language of the claims.

Claims

1. A radiation detector assembly comprising: main pixel array; a reference pixel array, wherein each reference pixel is adjacent to at least one primary pixel, the reference pixel having an area smaller than the primary pixel; and A detection circuit is in communication with the primary pixel array and the reference pixel array, wherein the detection circuit is configured to detect signals in the primary pixel and the reference pixel occurring simultaneously.

2. The radiation detector assembly of claim 1 , further comprising: An anti-scatter collimator includes X-ray attenuating diaphragms arranged in a one-dimensional or two-dimensional geometry, wherein the anti-scatter collimator is positioned above the reference pixel array.

3. The radiation detector assembly of claim 1 , wherein: The detection circuit is further configured to combine signals determined to be coincident and measured in the reference pixel and the primary pixel that are adjacent to each other.

4. The radiation detector assembly of claim 3, wherein: Combining the signals includes summing the signals. 5 . The radiation detector assembly of claim 1 , the detection circuit further configured to detect a concurrent signal in response to a signal detected in a reference pixel.

6. The radiation detector assembly of claim 5, wherein: The signal from the reference pixel exceeds a noise threshold to trigger detection of a coincident signal.

7. The radiation detector assembly of claim 1 , wherein: The width of the reference pixel is less than or equal to half the width of the primary pixel, so that the count rate in the reference pixel is significantly lower than the count rate in the primary pixel when exposed to the same flux of X-rays.

8. The radiation detector assembly of claim 1 , wherein: The width of the reference pixel corresponds to the diameter of a charge cloud induced by an X-ray interaction with a semiconductor sensor of the radiation detector assembly.

9. The radiation detector assembly of claim 1 , wherein: The detection circuit is configured to associate an output signal of the detection circuit with a corresponding reference pixel or a corresponding primary pixel depending on whether a coincidence signal is detected.

10. The radiation detector assembly of claim 1, wherein: The detection circuit includes a first type of channel for the main pixel and a second type of channel for the reference pixel.

11. The radiation detector assembly of claim 1 , wherein: The detection circuit includes a respective channel for each primary pixel and each reference pixel, and wherein each channel includes at least one threshold comparator and an associated digital counter, and wherein the detection circuit is configured to increment the digital counter when the X-ray sensing signal exceeds the threshold of the respective comparator.

12. The radiation detector assembly of claim 1, wherein: The detection circuit includes a corresponding channel for each primary pixel and each reference pixel, and wherein each channel includes a peak detection circuit and an analog-to-digital converter configured to output a digital signal representing the X-ray sensing signal.

13. The radiation detector assembly of claim 1 , wherein: The detection circuit is configured to generate an output representing the relative position of absorbed X-rays between the corresponding primary pixel and the adjacent reference pixel by combining signals from the corresponding primary pixel and the adjacent reference pixel, thereby generating sub-pixel spatial resolution.

14. The radiation detector assembly of claim 13, wherein: The relative positions of the absorbed X-rays are determined based on a ratio of signals induced in the corresponding primary pixel and adjacent reference pixels.

15. A method for detecting and correcting charge sharing events in a radiation detector, comprising: Reading out signals from a primary pixel array and a reference pixel array, wherein each reference pixel is adjacent to at least one primary pixel and the reference pixel has an area smaller than the primary pixel; detecting a signal read out from a corresponding reference pixel; executing coincidence logic to identify a corresponding primary pixel adjacent to the corresponding reference pixel and to read out a corresponding signal, the coincidence logic detecting the signal of the reference pixel and the corresponding signal occurring simultaneously; and The signal from the corresponding reference pixel and the corresponding signal are combined.

16. The method according to claim 15, wherein Combining the signal and the corresponding signal includes summing the signal and the corresponding signal.

17. The method according to claim 15, wherein: Detecting the signal read out from the corresponding reference pixel includes comparing the signal to a noise threshold.

18. The method according to claim 15, further comprising: A sub-pixel spatial resolution is determined based on a ratio of the signal to the corresponding signal.

Images