X-ray phase contrast detector

Abstract

The present invention is entitled "X-ray Phase Contrast Detector". The present disclosure relates to the manufacture and use of phase contrast imaging detectors that include sub-pixel resolution electrodes or photodiodes that are spaced to correspond to a phase contrast interference pattern. Systems using such detectors can employ fewer gratings than are typically used in phase contrast imaging systems, where certain functionality typically provided by detector-side analyzer gratings is performed by the sub-pixel resolution structures (e.g., electrodes or photodiodes) of the detector. Measurements acquired using the detector can be used to determine the offset, amplitude, and phase of the phase contrast interference pattern without requiring multiple acquisitions at different phase steps.

Description

Background Art

[0001] The subject matter disclosed herein relates to X-ray imaging, and in particular to phase contrast imaging techniques.

[0002] Non-invasive imaging techniques allow images of a patient's internal structures or features to be obtained without performing invasive procedures on the patient. Specifically, such non-invasive imaging techniques rely on various physical principles (such as differential transmission of X-rays through a target volume, reflection of acoustic waves within the volume, paramagnetic properties of different tissues and materials within the volume, decomposition of targeted radionuclides in the body, etc.) to collect data and construct images or otherwise represent observed internal features of the patient.

[0003] By way of example, in the context of X-ray absorption imaging techniques, X-rays are generated by an X-ray source and pass through an object or patient to reach a radiation detector. The X-rays are differentially absorbed based on the composition of the material in the beam path as they pass through the object or patient. This differential X-ray absorption allows the generation of an image using the detected X-rays, which provides information about the internal structure of the object or patient. Such images, which can be considered conventional X-ray images, can be used to convey information about materials (such as bones) that strongly absorb X-rays (i.e., high atomic weight or high-Z materials) that exhibit high contrast with less absorbing materials. However, such X-ray absorption-based imaging methods are less useful in generating information that is independent of the location of such high-Z materials. To address this deficiency, other imaging techniques can be used. One such technique that also uses X-rays is phase contrast imaging, which can produce information that is not available with conventional absorption X-ray imaging. In phase contrast imaging, a series of gratings or coherent sources used in conjunction with fewer gratings can be used to impart phase and interference properties to the X-ray beam. The resulting X-ray beam, when detected, yields additional information, particularly about low-atomic-weight structures and microstructures in the path of the X-ray beam. However, one consequence of using a grating is that a portion of the emitted X-rays may be absorbed by the grating, which may require a higher X-ray dose to obtain a suitable image. Summary of the invention

[0004] The following shows an overview of certain embodiments disclosed herein. It should be understood that these aspects are provided only to provide the reader with a brief overview of these specific embodiments, and these aspects are not intended to limit the scope of the present disclosure. In fact, the present disclosure may cover various aspects that may not be shown below.

[0005] In one embodiment, a phase contrast imaging detector is provided. According to the embodiment, the phase contrast imaging detector includes a plurality of pixels. Each pixel includes: a detection material that generates a measurable parameter in response to an X-ray photon; and a plurality of sub-pixel resolution readout structures. The sub-pixel resolution readout structures are in an alternating pattern corresponding to a phase contrast interference pattern. The phase contrast imaging detector also includes a readout circuit that is configured to read out a signal from the plurality of sub-pixel resolution readout structures.

[0006] In another embodiment, a phase contrast imaging system is provided. According to this embodiment, the phase contrast imaging system includes: an X-ray source, the X-ray source being configured to emit X-rays through an imaging volume during operation; and at least one grating position, the at least one grating position being located in the X-ray path through the imaging volume. The X-rays generate a spatial interference pattern when passing through the at least one grating. The phase contrast imaging system also includes a phase-sensitive detector, the phase-sensitive detector being configured to generate a signal in response to the X-rays emitted by the X-ray source and passing through the at least one grating. The phase-sensitive detector includes a plurality of pixels. Some or all of the pixels each include a plurality of sub-pixel resolution readout structures. The sub-pixel resolution readout structures are in an alternating pattern corresponding to the phase contrast interference pattern. The phase contrast imaging system also includes one or more processing components, the one or more processing components being configured to process the signals generated by the plurality of sub-pixel resolution readout structures to generate a clinical image.

[0007] In an additional embodiment, a method for phase contrast imaging is provided. According to the method, an X-ray source of an imaging system is operated to emit X-rays through an imaging volume. After the X-rays pass through at least one grating that imparts a spatial interference pattern to the X-rays, the X-rays are received at a phase-sensitive detector. The phase-sensitive detector includes a plurality of pixels. Some or all of the pixels each include a plurality of sub-pixel resolution readout structures. The sub-pixel resolution readout structures are in an alternating pattern corresponding to the phase contrast interference pattern. The signals generated by the plurality of sub-pixel resolution readout structures are processed to determine the amplitude, offset, and phase of the phase contrast interference pattern. At least one or more of the phase, the amplitude, and the offset are used to generate a phase image, an absorption image, or a dark field image. BRIEF DESCRIPTION OF THE DRAWINGS

[0008] 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 refer to like parts throughout the several views, and in which:

[0009] Figure 1 is a schematic diagram of an imaging system for producing an image according to aspects of the present disclosure;

[0010] Figure 2 An example of a phase contrast imaging system employing three gratings is depicted;

[0011] Figure 3 depicts generalized waveforms showing parameters of a spatial interference pattern according to aspects of the present disclosure;

[0012] Figure 4 An example of a phase contrast imaging system employing two gratings and a phase sensitive detector according to aspects of the present disclosure is depicted;

[0013] Figure 5 Examples of detector modules and pixels are depicted;

[0014] Figure 6 depicts an example of a phase-sensitive detector module according to aspects of the present disclosure;

[0015] Figure 7 Another example of a phase-sensitive detector module according to aspects of the present disclosure is depicted;

[0016] Figure 8 depicts side and top views of a scintillator-based phase-sensitive detector pixel according to aspects of the present disclosure;

[0017] Fig. 9 depicts an example of a pixel including an interlaced comb structure according to aspects of the present disclosure; and

[0018] Fig.10 An example of a specific implementation of coincidence logic according to aspects of the present disclosure is depicted. DETAILED DESCRIPTION

[0019] One or more specific embodiments will be described below. In order to provide a concise description of these embodiments, all features of an actual implementation may not 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. In addition, it should be understood that such development efforts may be complex and time-consuming, but are still routine tasks for design, fabrication, and manufacturing for ordinary technicians who benefit from this disclosure.

[0020] 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 in addition to the listed elements. Furthermore, any numerical examples in the following discussion are intended to be non-limiting, and thus the additional numerical values, ranges, and percentages are within the scope of the disclosed embodiments.

[0021] Practicing physicians often use X-ray imaging to non-invasively gather information about structures within a patient's body. Conventional X-ray imaging relies on the differential absorption of X-rays by different constituent structures within the body to provide useful data. Specifically, such differentially absorbed X-rays can be used to generate images that exhibit good contrast for high atomic weight (i.e., high Z) materials within the body, but provide limited information about features that do not absorb X-rays well.

[0022] Phase contrast imaging is also an X-ray based imaging technique that typically utilizes a series of gratings (e.g., absorption gratings or phase gratings) to impart phase and spatial interference properties to an X-ray beam. The detected X-rays can then be used to reconstruct not only conventional X-ray absorption images, but also phase images (which have good contrast for low-Z materials (such as tissues)) and dark field images (which allow detection of sub-pixel microstructures (e.g., alveoli) that are detected despite not being resolved). In particular, phase images may be useful in contexts where there are low-Z objects that clinicians wish to view or detect in an environment that also includes high-Z features. One such example is a lesion within the liver parenchyma, where the tissue properties between the liver and the lesion are slightly different. To generate such a phase image, the Compton absorption effect is measured in the absence of photoelectric absorption. However, the gratings used for phase contrast imaging to generate such images typically absorb a significant portion of the X-rays emitted for imaging. Therefore, phase contrast imaging is typically performed with a higher X-ray dose than that used for conventional absorption imaging.

[0023] The techniques disclosed herein help to address these problems. Specifically, as disclosed herein, certain embodiments of the present invention eliminate one of the gratings typically used for phase contrast imaging (i.e., the detector-side grating, which is denoted as the analyzer grating), and employ a detector that integrates the functionality of the grating into the detector itself. In this way, dose losses attributable to the grating are avoided. This is particularly important in medical imaging, where X-rays absorbed by the detector-side grating have already passed through the patient, thereby creating additional X-ray dose to the patient that does not reach the detector. In addition, incorporating this functionality into the detector also eliminates the need for multiple acquisitions at different phase steps using the analyzer grating, as discussed in more detail herein.

[0024] Taking the foregoing into account, Figure 1 An example of an X-ray imaging system 10 suitable for acquiring data for reconstruction as discussed herein is provided in . It should be understood that the X-ray based imaging system 10 can be any suitable X-ray imaging modality, such as a computed tomography (CT) imaging system, a C-arm type imaging system, a tomosynthesis imaging system, a conventional radiography imaging system, a mammography system, a fluoroscopic imaging system, etc.

[0025] exist Figure 1 In the illustrated embodiment, the imaging system 10 includes an X-ray radiation source 12 and a detector 14. The X-ray source 12 may be an X-ray tube, or include any other X-ray radiation source suitable for acquiring medical or other images. The X-rays 16 generated by the source 12 enter an imaging region where an object to be imaged (e.g., a portion undergoing non-destructive evaluation or testing), a tissue sample to be imaged, a bag or package undergoing security screening, a patient undergoing an imaging regimen, etc. may be located. In the illustrated example, a patient 18 undergoing imaging during surgery is positioned within the imaging volume. In the depicted example, the X-rays 16 are collimated into a fan-shaped (planar) beam or a cone (volumetric) beam that passes through the imaged volume.

[0026] A portion of the X-ray radiation 20 passes through or around the patient 18 (or other subject of interest) and strikes a detector array, generally represented as detector 14. The detector elements (e.g., pixels) of the detector 14 generate electrical signals that are representative of the intensity of the incident X-rays 20 on the detector elements of the detector 14. As discussed herein, these signals are acquired and processed to reconstruct an image of features within the patient 18.

[0027] With respect to the detector 14 as used herein, various types of detectors are contemplated that may be suitable for phase contrast imaging according to the present invention. Generally speaking, the detector 14 as used herein will include an array of pixels. Each pixel is associated with some medium that interacts with the X-rays and with a sub-pixel resolution electrode that measures some property or signal generated in response to the interaction of the X-rays with the medium. In certain embodiments as discussed herein, the electrodes associated with a given pixel are in an interlaced or other discontinuous configuration (i.e., not a 1-D or 2-D array of adjacent electrodes), which, in implementation, provides the ability to distinguish signals at sub-pixel resolution within a given pixel, which in certain specific implementations can be used to provide functionality comparable to that achievable by providing an analyzer grating at the surface of the detector 14.

[0028] In the context of scintillator-based embodiments (e.g., in embodiments in which a scintillator is used to convert high energy X-ray photons into lower energy optical photons that are detected), the electrodes associated with each pixel may take the form of an interlaced or otherwise discontinuous photodiode electrode structure suitable for reading out signals generated in response to the optical photons. In the context of direct conversion detectors (i.e., detectors in which the X-ray photons themselves are detected and a response signal is generated (rather than secondary photons generated by the scintillator), such as via interaction with a semiconductor material, the electrodes associated with each pixel may similarly be provided as an interlaced or otherwise discontinuous set of electrodes. Such direct conversion detectors may be configured to provide certain additional functionality, such as energy integrating and / or photon counting detectors. In the photon counting context, detector 14 may also be used to collect coincidence detection and / or spectral information.

[0029] In the depicted example, gratings 22, 24 (e.g., absorption gratings or phase gratings) are located in the path of the X-ray beam. If the source 12 is sufficiently small and / or coherent, the grating 22 may be the only grating provided. However, in implementation, additional source-side gratings may be present near the source 12 to effectively make the larger and / or incoherent source 12 perceived as multiple spatially coherent line sources in the context of phase contrast. Such gratings may be present to facilitate phase contrast imaging. In the depicted example, an analyzer grating 24 is also provided relative to the source 12 and opposite to the patient 18 (i.e., on the detector side).

[0030] In this example, the source 12 and detector 14 (as well as any gratings, filters, collimators, etc.) may be part of an imager subsystem 30. In certain imaging modalities (e.g., computed tomography (CT), C-arm angiography, tomosynthesis), during a scanning process of acquiring projection data, the source 12 and detector 14 of the imager 30 may be moved along one or more axes relative to the patient or the object being imaged. For example, the imager 30 may be moved around a first axis of rotation, a second axis of rotation, or a third axis of rotation, or any combination thereof. In one embodiment, the translation and rotation of the imager 30 may be determined or coordinated according to a specified protocol. Alternatively, such as in non-destructive testing applications, the imager 30 may remain constant while the object is repositioned.

[0031] Movement of the imager 30, if any, may be initiated and / or controlled by one or more linear / rotational subsystems 46. The linear / rotational subsystems 46 may include support structures, motors, gears, bearings, etc. that enable relative rotational and / or translational movement of the imager 30. In one embodiment, the linear / rotational subsystems 46 may include a structural device (e.g., a C-arm, a rotating gantry, a turntable, etc.) that supports the source 12 and detector 14 or alternatively supports the subject or patient 18.

[0032] The system controller 48 may control the linear / rotational subsystem 46 that initiates and / or controls the movement of components of the imager 30. In implementation, the system controller 48 may be combined with one or more processing devices that include or communicate with a tangible non-transitory machine-readable medium that collectively stores instructions executable by one or more processors to facilitate the execution of imaging operations. The system controller 48 may also include features that control the timing of activation of the source 12, for example, to control the acquisition of X-ray data obtained during a particular imaging sequence. The system controller 48 may also perform various signal processing and filtering functions, such as initial adjustment for dynamic range, interleaving of digital projection data, and the like. Thus, in general, the system controller 48 may be viewed as commanding the operation of the imaging system 10 to execute an examination protocol. It should be noted that, for ease of discussion, reference is made below to the system controller 48 as a unit that controls acquisition, movement, and the like using the imager 30. However, the present disclosure also encompasses embodiments in which the system controller 48 functions in conjunction with other control devices (e.g., other control circuits local to the imager or remote from the system 10).

[0033] In the present context, the system controller 48 includes signal processing circuitry and various other circuitry that enables the system controller 48 to control the operation of the imager 30 and the linear / rotational subsystem 46. In the illustrated embodiment, the circuitry may include an X-ray controller 50 configured to operate the X-ray source 12. The circuitry of the system controller 48 may also include one or more motor controllers 52. The motor controller 52 may control the activation of the various components responsible for moving the source 12 and the detector 14. In other words, the motor controller may implement a particular acquisition trajectory or motion for the relative motion of the components of the imager 30.

[0034] The system controller 48 is also shown as including one or more data acquisition systems 54. In general, the detector 14 may be coupled to the system controller 48, and more specifically to the data acquisition system 54. The data acquisition system 54 may receive data collected by the readout electronics of the detector 14, and in certain embodiments may process the data (e.g., by converting analog signals to digital signals or performing other filtering, transformation, or the like).

[0035] It should be noted that tangible, non-transitory machine-readable media and processors configured to execute instructions stored on such media present in the system 10 may be shared between various components of the system controller 48 or other components of the system 10. For example, as shown, the X-ray controller 50, the motor controller 52, and the data acquisition system 54 may share one or more processing components 56, each of which is specifically configured to cooperate with one or more memory devices 58 storing instructions that, when executed by the processing components 56, perform image acquisition and reconstruction techniques. In addition, the processing components 56 and the memory components 58 may coordinate to perform various image reconstruction processes.

[0036] The system controller 48 and the various circuits it includes, as well as the processing component 56 and the memory component 58, may be accessed or otherwise controlled by an operator via an operator workstation 60. The operator workstation 60 may include any special or general purpose computer that may include one or more programs (e.g., one or more imaging programs) that enable operator input for the techniques described herein. The operator workstation 60 may include various input devices, such as a mouse, keyboard, trackball, or any other similar feature that enables an operator to interact with a computer. The operator workstation 60 may enable an operator to control various imaging parameters, for example, by adjusting certain instructions stored on the memory device 58.

[0037] The operator workstation 60 may be communicatively coupled to a printer 62 for printing images, patient data, and the like. The operator workstation 60 may also be in communication with a display 64 that enables the operator to view various parameters in real time, to view images generated from the acquired data, and the like. In certain embodiments, the operator workstation 60 may also be communicatively coupled to a picture archiving and communication system (PACS) 66. Such a system may enable storage of patient data, patient images, image acquisition parameters, and the like. Such stored information may be shared throughout the imaging facility and may also be shared by other facilities, such as remote clients 68. Remote clients 68 may include hospitals, physician offices, or any other similar clients.

[0038] about Figure 2 Various aspects of the present method can be further understood by describing a conventional phase contrast imaging technique. Such phase contrast imaging techniques typically determine the phase shift angle for each voxel by comparing the spatial interference pattern observed with an object in the beam path to the spatial interference pattern when the object is not present. In a conventional phase contrast computed tomography (CT) imaging system (aspects of which are described in Figure 2 In the context of an X-ray source 12 (shown in FIG. 1 ), gratings (typically three gratings) may be employed to generate interference patterns. In this example, a source-side grating G0 (indicated by reference numeral 80) is located near the X-ray source 12 to ensure spatial coherence and effectively create an array of individually coherent but mutually incoherent emissions in response to operation of an incoherent X-ray source. Phase objects in the beam path cause each coherent subset of X-rays to be slightly refracted. The angular deviations so introduced result in variations in the local transmission intensity through a pair of gratings G1 and G2 (indicated by reference numerals 82 and 84, respectively), which variations in the local transmission intensity can be detected by the X-ray detector 14 (e.g., Figure 1 82 is located between the source 12 and the detector 14 and imprints a periodic interference pattern onto the wavefront. The grating G2 84 is located near the detector (i.e., G2 is the detector-side grating) and resolves the sub-pixel resolution interference pattern modulation. In implementation, the gratings may be fabricated from silicon wafers using photolithography and electroplating where appropriate. It may be desirable for the G1 grating 82 to be close to the light source 12 in order to obtain the best possible sensitivity (which is Figure 2 function of the distance L2 shown).

[0039] About the spatial interference patterns created in phase contrast imaging and go to Figure 3, a brief explanation may help understand the following discussion. Specifically, the interference pattern created in phase contrast imaging has three parameters that can be measured (the period is known by design): (1) offset; (2) phase; and (3) amplitude. As in conventional X-ray imaging, the offset as used herein is caused by overall attenuation. The phase is the spatial offset of the interference pattern relative to a reference, which corresponds to the gradient of the real part of the total refractive index obtained by X-rays passing through the object to be tested. The amplitude of the interference pattern is the difference between the peaks and valleys of the pattern, which can be reduced in the presence of microstructures. These three parameters are Figure 3 is shown in the context of a sample pattern or waveform.

[0040] Considering the foregoing, although Figure 2 The conventional arrangement shown is suitable for phase contrast imaging, but such an arrangement may require a high X-ray flux because some portion of the X-rays are absorbed by each of the three gratings, thereby reducing the flux reaching the detector 14. By way of example, compared to conventional CT (i.e., absorption-based imaging) in which all of the X-ray flux (without the portion absorbed by the patient 18) reaches the detector 14, in grating-based phase contrast imaging, only a limited portion (e.g., 25%) of the transmitted flux may reach the detector 14 due to accidental absorption by the various gratings. In a medical context, this may result in a higher flux being employed in order to stop patient motion and obtain useful images. Grating G2 284 absorbs a portion of the X-ray flux that passes through the patient, thereby reducing the dose efficiency of the imager subsystem 30.

[0041] To reduce the flux loss attributable to the grating in phase contrast imaging, the present invention eliminates the detector side grating G2 284 and integrates the functionality of the grating into the phase sensitive detector 14A, as shown in FIG. Figure 4 As shown. In addition to eliminating the loss of X-ray flux (and associated patient dose) due to X-ray absorption by grating G2, this technique also has the benefit of eliminating the need to perform multiple acquisitions at different phase steps, which is typical in phase contrast imaging schemes. Specifically, in conventional phase contrast imaging acquisitions, when the analyzer grating G2 84 is used in front of a standard CT detector, only one measurement can be made for each pixel at each time. To do this, the analyzer grating G2 84, which has a periodicity equal to the frequency of the interference pattern, is typically offset, and additional subsequent measurements are made (e.g., for a total of 3 measurements). By employing a phase-sensitive detector 14A (e.g., a detector in which multiple pixel electrodes provide sub-pixel resolution corresponding to the period of the interference pattern, as discussed in more detail below), all 3 independent measurements can be made simultaneously.

[0042] Accordingly, Figure 5 , Figure 6 and Figure 7 An example of a detector module 90 that can be used to form the detector 14A and an example of pixels 92 disposed in the detector module 90 are depicted. In one embodiment, the detector module 90 includes a plurality of pixels 92 formed using silicon, cadmium zinc telluride (CZT), cadmium telluride (CdTe), or other suitable detector materials that generate signals in response to x-rays without a scintillator intermediate component (i.e., a direct conversion detector material). However, in other embodiments, a scintillator may be present as part of the detector 14, such as with respect to Figure 8 and Fig. 9 Discussed in more detail. Figures 5 to 7 As shown, the length of the pixel 92 may increase in the direction of the X-ray beam path 20 to allow for similar responses at different depths within the detector. Alternatively, a single long pixel may be provided that is aligned with the direction of the X-ray beam. The readout circuitry 96 is depicted as being disposed on the side or bottom of the detector module 90 so that the detector modules 90 may be combined side by side or end to end to provide a large or configurable field of view.

[0043] With regard to these drawings, Figure 5 A detection or conversion material 94 (such as a semiconductor material) is depicted associated with a pixel 92 that interacts with the X-rays. Figure 5 In the example of , a single continuous electrode 98 is associated with the conversion material 94 of the pixel 92 within the corresponding pixel. Figure 6 and Figure 7 Depicted within respective pixels 92, a plurality of non-continuous electrodes 98A, 98B, and 98C are associated with the conversion material 94 of each respective pixel 92 to provide sub-pixel resolution measurements. In the depicted example, Figure 6 and Figure 7 The electrodes are provided in an interlaced comb pattern which, in one embodiment, may have a sub-pixel resolution distance and pitch corresponding to the frequency of the spatial interference pattern of a given phase contrast imaging system.

[0044] For example, in Figure 6 In , electrodes 98A and 98B form a pattern comprising alternating arms of two interwoven combs. Figure 7, electrodes 98A, 98B, and 98C form a pattern comprising alternating arms of three interlaced combs, one of which is spanned by a conductive bridge 100 disposed outside the plane of the other conductive traces so as to allow individual electrical connections to be made to each electrode. It may be noted that embodiments are envisioned in which all pixels 92 of the detector module 90 have multiple sub-pixel resolution electrodes, as well as embodiments in which only some of the pixels 92 of the detector module 90 (such as some or all of the pixels 92 in the center of the detector module 90) have multiple sub-pixel resolution electrodes. As discussed herein, the staggered pattern of electrodes 98 may have a spacing corresponding to the period of the phase contrast interference pattern, thereby allowing all three parameters of interest (i.e., offset, phase, and amplitude) to be measured simultaneously, rather than requiring multiple discrete measurements in which the analyzer grating is offset, resulting in the need for increased acquisition time and patient dose. Although in Figure 6 and Figure 7 Patterns of two and three interleaved sub-pixel resolution electrodes are shown in , respectively, but alternative configurations are contemplated using additional interleaved sub-pixel resolution electrodes as needed to accurately estimate the phase contrast signal.

[0045] The direction in which the electrodes 98 (or other patterned features, such as photodiodes) extend may vary depending on the embodiment and may be in the direction of the X-ray beam 20 (as shown) or orthogonal to the X-ray beam 20. When extending parallel to the X-ray beam 20, the lines of the comb 98 form a line pattern in the direction of the X-ray beam 20, such as Figure 6 and Figure 7 Conversely, when the pattern features are orthogonal to the X-ray beam 20, any pattern having a pitch corresponding to the phase contrast interference pattern as described above may be used, including coded apertures.

[0046] Go to Figure 8 , another embodiment involving a photodiode-based phase contrast detector is shown. In this example, a side view (above) and a top view (below) are shown together to better depict various features. A single detector pixel 92 is shown, wherein the side view shows the stacking of components associated with the detector pixel 92, and the top view shows only the sub-pixel photodiode arrangement with associated separate readout paths (a1, a2, a3) corresponding to different phases 1, 2, and 3 of the spatial interference pattern.

[0047] Specifically, Figure 8The top view shown depicts a high-resolution scintillator 106 in optical communication with a plurality of sub-pixel photodiodes 110 via an optical coupling layer 108, which may not be present in certain embodiments. The scintillator 106 may be formed of a high-Z material and may be arranged in columns or other structures corresponding to the pixels 92 (i.e., the scintillator 106 may be pixelated), or may be monolithic with respect to the detector surface.

[0048] In the depicted example, the X-ray photons 20 interact with the scintillator 106 to generate low energy photons (e.g., optical photons) that can be detected by the sub-pixel photodiodes 110. As shown, the sub-pixel photodiodes 110 are grouped into an alternating, separated pattern to form three interlaced photodiodes 110A, 110B, and 110C, having a spacing distance corresponding to the frequency of the spatial interference pattern of the phase contrast imaging system. In this example, the pixel 92 is read out using three separate interlaced sub-pixel photodiodes 110A, 110B, and 110C, so that the readout channel associated with the pixel 92 is composed of multiple sub-pixels. In some specific implementations, the width of the sub-pixel photodiode 110 ranges from a few μm (e.g., 2 μm to 5 μm) to tens of μm (e.g., 30 μm to 80 μm). In one embodiment, the sub-pixel photodiode 110 can be manufactured as a complementary metal oxide semiconductor (CMOS) device, a lateral metal semiconductor metal (MSM) device, and the like. As in the previous examples and as further explained below, the signals of corresponding (i.e., same phase) sub-pixel photodiodes 110 are combined to form a "phase" comb in either the pre-readout (i.e., analog) domain or the post-readout (i.e., digital) domain. Figure 8 A pattern of three interleaved sub-pixel resolution electrodes is shown in , but alternative configurations are envisioned using 2 or more interleaved sub-pixel resolution photodiodes as needed to accurately estimate the phase contrast signal.

[0049] Go to Fig. 9 ,like Figure 7The patterned pixel shown having a pattern of three interlaced combs 98A, 98B, 98C is shown in conjunction with an X-ray photon 104 that interacts with a detection material (e.g., a direct conversion material or, alternatively, a scintillating material) to generate measurement signals a1, a2, and a3 that can be read out for the pixel 92. For each detected X-ray photon 104, a charge cloud is created that diffuses as it travels through the detection material or an intermediate substrate to generate a signal that is measured or detected by one or more of the sub-pixel resolution electrodes 98 (or, alternatively, photodiodes 110). In this way, the total energy deposited during the event (i.e., the total energy deposited by a given X-ray photon interaction) can be determined. In a detector system equipped with coincidence detection logic (discussed in more detail below), measurements at different electrodes 98 (or photodiodes 110, as the case may be) can be determined to coincide in time and, therefore, be attributed to the same X-ray photon. Furthermore, the center of the charge cloud (i.e., the position in the interference pattern) can be located at sub-pixel resolution, and the size of the charge cloud (which may depend on the depth within the detection material where the photon is absorbed and on the photon energy) can be determined more accurately. Accordingly, the measured values ​​a1, a2, and a3 associated with different corresponding electrodes 98 (or photodiodes 110) detecting an X-ray photon event can be used to determine the three parameters of interest for a single X-ray photon.

[0050] For example and reference Figure 3 , we can assume that the interference pattern has an intensity distribution I(x) of the following form:

[0051] (1) The period is p and the offset is A DC , amplitude is A and phase is Assuming that the pattern is sampled at three locations each spaced one-third (1 / 3) of the period and having measured amplitudes a1, a2, and a3, then a1 = I(0), a2 = I(p / 3), and a3 = I(2p / 3). It will be appreciated that the period p need not be measured, as it is determined by the geometry of the system or device. The parameters of the interference pattern can then be calculated from the measured amplitudes a1, a2, and a3 according to the following:

[0052] Offset(A DC ):

[0053] (2)

[0054] Amplitude (A):

[0055] (3)

[0056] Phase :

[0057] (4) If a2-a3≥0, then If a2-a3<0, then

[0058] In this example, three sub-pixel measurement channels (i.e., a1, a2, a3) are required to explicitly measure three parameters of interest. Alternative formulas can be used to account for the size of each sub-pixel resolution electrode or photodiode, the different number of sub-pixel resolution electrodes or photodiodes, improved robustness in the presence of noise, and the like. In some embodiments, the detection threshold of each sub-pixel channel can be reduced relative to the detection threshold for a single pixel readout channel to avoid signal loss, that is, considering that the signal associated with the X-ray photon event can be separated between multiple channels rather than being attributable to a single channel. That is, the threshold-based cutoff value can be adjusted to consider that a given signal can be separated between two or more electrode readout channels (rather than a single electrode readout channel) to prevent discarding valid measurements. Coincidence detection logic can be used to count only X-ray photon events where the sum of all channels is above a suitable total threshold. That is, the detection threshold of each sub-pixel electrode channel can be reduced relative to the threshold that may be used for a single electrode readout, while still requiring that the sum of all sub-pixel channels is at least equal to the threshold that would be used when only a single electrode readout channel is used.

[0059] With respect to the charge clouds associated with X-ray photon events and the use of coincidence detection logic described above, the importance of such logic (as can be determined from the previous discussion) relates in part to the ability to assign or otherwise associate all detected charge for an X-ray photon event (which may be spread across multiple sub-pixel resolution electrodes 98 or photodiodes 110) to a single sub-pixel resolution electrode 98 or photodiode 110. Fig.10 , shows an example of coincidence detection in the context of multiple sub-pixel resolution electrodes 98 (or photodiodes in an alternative context). In this example, X-ray photons interacting with the detection material of the corresponding pixel 92 of the detector 14 (e.g., a photon counting detector with built-in analog-to-digital conversion (ADC) for each channel) cause the generation of an initial charge cloud 120. A voltage applied between the top electrode 98 and the bottom electrode (not shown) causes the charge cloud 120 to move and spread toward the top electrode (i.e., the sub-pixel resolution electrode 98) (indicated by reference numeral 122), thereby causing a measurable signal (a1, a2, a3) to be generated at the multiple electrodes 98. In this example, the signal (e.g., the induced charge) at each electrode 98 is collected and processed by the corresponding charge sensitive amplifier (CSA) and shaper 126 of each channel. The signal of each channel is passed to the ADC circuit 128, which is composed of a comparator output based on a threshold voltage V th Trigger, where V this set close to but above the noise floor (e.g., 5keV to 10keV).

[0060] Each detected event generates a timestamp t and a digital output d for each associated electrode (or photodiode). The timestamp t is passed to the coincidence logic 140, which determines whether charge has been shared between the electrodes 98 based on the temporal proximity of the corresponding timestamps t. Based on the determination that charge from the X-ray photon event is shared between the electrodes 98, the "winner takes all" logic can be activated, and the correct (i.e., incident) counter is incremented for the electrode (or photodiode) determined to correspond to the position of the X-ray photon event (such as based on the maximum measured signal). In some specific implementations, the "winner takes all" logic may also require that the total energy measured for the X-ray event is greater than an energy threshold in order to increment the appropriate counter. In some embodiments, an additional "winner takes all" aggregation step may occur to account for potential charge sharing between pixels 92 (such as potential charge sharing caused by X-ray photons incident near the edges of two or more pixels 92).

[0061] Technical effects of the present invention include a phase contrast imaging detector having a detector that includes a patterned pixel surface (e.g., staggered or interlaced photodiodes or electrodes). A system using such a detector can perform phase contrast imaging with only two gratings instead of three gratings, where the functionality of the third grating is performed by a discontinuous pattern of pixel electrodes or photodiodes. Measurements collected using the detector can be used to determine the offset, amplitude, and phase of the phase contrast interference pattern without requiring multiple acquisitions at different phase steps of the analyzer grating.

[0062] This written description uses examples to disclose the invention, including the best mode, and also to enable those 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 fall 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 that have minor differences from the literal language of the claims.

Claims

1. A phase contrast imaging detector, the phase contrast imaging detector include: Multiple pixels, each pixel includes: a detection material that generates a measurable parameter in response to the X-ray photons; a plurality of electrodes, wherein the plurality of electrodes are coupled to and interact with the detection material, and the plurality of electrodes are in an alternating pattern of interlaced combs, the spacing between the interlaced combs corresponding to the frequency of the phase contrast interference pattern; and a readout circuit coupled to the plurality of electrodes and configured to read out signals from the plurality of electrodes, wherein the alternating pattern of the interlaced comb is aligned with the direction in which the X-ray beam impinges on the pixel; The length direction of the fingers of the interlaced comb is aligned with the direction in which the X-ray beam is incident on the pixel; and The length of the pixels varies along the direction of the X-ray beam.

2. The phase contrast imaging detector of claim 1, wherein the plurality of electrodes comprises two, three or more electrodes.

3. A phase contrast imaging detector according to claim 2, wherein the detection material comprises a direct conversion material capable of generating a charge cloud in response to each X-ray photon, wherein the charge cloud is measured by the two, three or more electrodes.

4. The phase contrast imaging detector of claim 1, wherein readout signals from the plurality of electrodes are used to simultaneously determine the shift, amplitude and phase of the phase contrast interference pattern.

5. The phase contrast imaging detector according to claim 4, wherein the intensity distribution I(x) of the phase contrast interference pattern is: in, p is the period, A DC is the offset, A is the amplitude, and φ is the phase.

6. The phase contrast imaging detector of claim 1, wherein the plurality of electrodes are used as a detector-side analyzer grating in a phase contrast imaging system.

7. A phase contrast imaging detector according to claim 1, wherein the signals collected from the multiple electrodes can be used to simultaneously determine the amplitude, offset and phase of the phase contrast interference pattern without the need for multiple collections at different phase steps of the analyzer grating.

8. The phase contrast imaging detector of claim 1, wherein the readout circuit is configured to locate the center of a charge cloud generated in response to an incident X-ray photon to a corresponding electrode using coincidence logic.

9. A phase contrast imaging system, the phase contrast imaging system include: an X-ray source configured to emit X-rays through an imaging volume during operation; at least one grating located in a path of X-rays passing through the imaging volume, wherein the X-rays have a phase contrast interference pattern when passing through the at least one grating; a phase-sensitive detector configured to generate a signal in response to X-rays emitted by the X-ray source and passing through the at least one grating, wherein the phase-sensitive detector comprises: a plurality of pixels, wherein some or all of the pixels each include a plurality of electrodes, wherein the plurality of electrodes are in an alternating pattern of interlaced combs, the spacing between the interlaced combs corresponding to the frequency of the phase contrast interference pattern; and one or more processing components configured to process the signals generated by the plurality of electrodes to generate an image, wherein the alternating pattern of the interlaced comb is aligned with the direction in which the X-ray beam impinges on the pixel; The length direction of the fingers of the interlaced comb is aligned with the direction in which the X-ray beam is incident on the pixel; and The length of the pixels varies along the direction of the X-ray beam.

10. The phase contrast imaging system of claim 9, wherein the pixel further comprises a direct conversion material capable of generating a charge cloud in response to each X-ray photon, wherein the charge cloud is measured by the plurality of electrodes.

11. The phase contrast imaging system of claim 9, wherein the plurality of electrodes comprises two, three or more electrodes.

12. The phase contrast imaging system of claim 9, wherein the image comprises one or more of a phase image, an absorption image, or a dark field image.

Images