Method and system for determining virtual output of a multi-energy x-ray imaging device

Abstract

The present disclosure relates to a method and apparatus for determining a virtual output of a multi-energy x-ray device. Based on the application for which the x-ray device is being used, a general algorithm can be determined or selected. Inputs received from the x-ray device can be substituted into the general algorithm to generate a virtual output algorithm for the x-ray device. The virtual output algorithm can then be used to calculate a virtual output.

Description

[0001] Cross-references to other applications

[0002] This application claims priority to U.S. Provisional Application No. 62 / 682,540, filed June 8, 2018, the contents of which are incorporated herein by reference. Technical Field

[0003] This disclosure relates generally to x-ray imaging, and more specifically to a method and system for determining the virtual output of a multi-energy x-ray imaging apparatus. Background Technology

[0004] The quality of medical images and their value as a tool depends solely on the extent to which they can convey the anatomy of the patient being imaged to the observer, such as a physician. The better the anatomy is understood, the more precise the information the physician has to make decisions.

[0005] In X-ray imaging, a major source of noise that frequently degrades image quality is anatomical noise. It is caused by the superposition of normal anatomy from a two-dimensional (2D) projection of a three-dimensional (3D) patient. This noise can blur the tissue being imaged or may be misread as anatomical abnormalities. A simple example of this is a chest X-ray taken for the purpose of assessing lung anatomy, in which the ribs are inevitably obscured in the resulting image. In this case, the ribs are the primary source of anatomical noise because they are not the anatomy of interest.

[0006] One proposed technique for reducing anatomical noise is dual-energy (DE) imaging. This technique utilizes a fundamental property of X-ray-matter interaction: different tissue types not only have different mass attenuation coefficients (μ / ρ(E)) across the diagnostic energy range, but also different rates of change of these coefficients.

[0007] One challenge in DE imaging stems from the need to obtain two separate low-energy and high-energy images. To achieve this, the X-ray spectrum absorbed at the detector should be reweighted at the low end of the diagnostic range for low-energy (LE) images and at the high end for high-energy (HE) images. DE imaging is capable of decomposing the patient's projection into images of soft tissue only and hard tissue only. Several mathematical methods exist for obtaining these DE images from LE and HE inputs, most notably logarithmic subtraction and basic decomposition.

[0008] In reality, complete removal of a specific tissue type is often impossible. Several factors contribute to the formation of non-ideal scenarios that cannot be captured by mathematical techniques. These include: a broad spectrum of X-ray flux that would result in each image being formed, contrary to the idealized source used in mathematical analysis; the inhomogeneity of the density or mass attenuation coefficient of the tissue being removed, making it impossible to determine the exact values ​​that should be used when calculating the weighting factors; and X-ray scattering from both the object being imaged and the detector that is not explained by Beer-Lambert's law. These non-idealities also mean that theoretical values ​​of the weighting factors cannot provide the best possible removal, requiring observers to calculate their ideal values ​​experimentally or qualitatively.

[0009] In practice, this spectral separation is achieved in two fundamentally different ways: the source spectra are different for the two images (referred to here as multiple-shot DE imaging), or the detector selectively absorbs different portions of a broader spectrum to form each image (referred to here as single-shot DE imaging). Regardless of the method used, a large separation of the two spectra is necessary to obtain high-quality tissue-selective images.

[0010] One way to acquire images at different energies is by acquiring them sequentially over time, without altering parts of the imaging system, but by changing the spectrum generated by the X-ray tube. This is the concept behind multiple-shot imaging (sometimes called kVp switching), where the first image is taken using a low-kVp X-ray tube, and immediately following, a second image is acquired using a high-kVp tube. Since the low-kVp and high-kVp beams will have different effective energies, the two resulting images will primarily contain information acquired at the low and high ends of the X-ray diagnostic spectrum, respectively. Alternatively, instead of modifying the source kVp between exposures, the source filter can be altered by rapidly moving the spectral filter into and out of the beam path. Considering the selective nature of the energy spectrum of the source filter, this would have the effect of presenting two different spectra to the detector.

[0011] This approach can also be extended to multi-energy images by obtaining several consecutive images at different kVp values ​​or source filters, which then allows for more spectral information to generate enhanced images algorithmically.

[0012] Unfortunately, the inherent time intervals in this technique cause motion artifacts to appear in the final images, which can pose a significant challenge to radiologists or observers interpreting them. These artifacts are noticeable distortions in the images caused by slight misalignments of the anatomy in consecutive images, and often originate from patient or object movement that occurs during and between image acquisitions.

[0013] Ideally, the source tube voltage could be changed instantaneously, allowing the next exposure to begin as soon as one is completed. However, currently commercially available sources require an interval of at least 150 ms to 200 ms between consecutive exposures. This is not only due to the varying voltage but also because the tube current needs to be adjusted to achieve the ideal relative intensity of the image. While this interval is short enough for most patients to avoid large movements, movements of the heart, breathing, and small muscles are confined to the entire interval. Motion artifacts will appear due to these movements, which can be specific obstacles in imaging the heart and lungs due to the large area of ​​the heart. Furthermore, this problem will be exacerbated by adding more image acquisitions to multi-energy imaging to allow for more patient movement, as the total acquisition time will increase.

[0014] Alternative methods exist for obtaining multi-energy images, often referred to as single-shot imaging. This method takes the opposite approach to multi-shot imaging, achieving spectral separation at the detector rather than the source. This is achieved by vertically stacking two sensor layers to form a bilayer detector known as a sandwich structure. One layer (such as the top layer) primarily absorbs LE x-rays, while the second or bottom layer absorbs HE x-rays. Therefore, using this technique, only a single exposure is required, performed at higher kVp to allow coverage of a large spectrum of both LE and HE x-rays. This method can then be extended to multilayer detectors, which can obtain multiple images with increased effective energies at subsequent stacked layers.

[0015] The practical problem arising from the single-shot approach is that, to achieve ideal and efficient energy separation between these layers, the mass load (or equivalent, their thickness) of the sensitive material (which is a scintillator or direct conversion material) must be tailored to specific tissue types and patient anatomy. Since it is commercially feasible to build only a few specific configurations, this leaves a compromise solution that optimally fits all target applications and patient types for a single practical application.

[0016] Therefore, a novel method and apparatus are provided for mitigating or overcoming at least one of the disadvantages of the above-described imaging methods and apparatus. Summary of the Invention

[0017] In one aspect of this disclosure, a method is provided for determining at least one virtual output of a multi-energy X-ray imaging apparatus, the method comprising: receiving from the multi-energy imaging apparatus a plurality of outputs generated by different X-ray spectra; determining a general algorithm based on an application of the X-ray imaging apparatus, physical properties of the X-ray imaging apparatus, or X-ray source exposure settings; substituting the plurality of outputs as inputs into the general algorithm to determine parameters, and generating a virtual output algorithm for the multi-energy X-ray imaging apparatus and the determined application; and using the virtual output algorithm to generate at least one virtual output.

[0018] On the other hand, the multiple outputs received from the multi-energy X-ray imaging device are obtained from some or all layers of the multi-energy X-ray imaging device; the multi-energy X-ray imaging device is a single-shot multi-layer X-ray imaging device. In a further aspect, the multiple outputs received from the multi-energy X-ray imaging device are obtained from two or more X-ray exposures performed with different X-ray source exposure settings; the multi-energy X-ray imaging device is a multi-shot X-ray imaging device. In another aspect, the X-ray source exposure settings include source voltage, source current, or source filtering. In yet another aspect, determining the general algorithm includes: determining the X-ray application for which the multi-energy X-ray imaging device is being used; and selecting the general algorithm based on the determined application.

[0019] On the other hand, the selection of general algorithms includes: selection for multi-slice X-ray imaging devices. As this general algorithm (where a, b, and c are parameters, S...) i For the signal at each layer and l i (For the defined number of layers). On the other hand, the selection of a general algorithm includes: selecting for multi-slice X-ray imaging devices. As this general algorithm (where b and c are parameters), The thickness of the scintillator pre-filter for each layer and t i (where the thickness is the scintillator thickness of the layer). On the other hand, determining the general algorithm includes selecting a minimization algorithm as the general algorithm.

[0020] In another aspect, utilizing the virtual output algorithm includes obtaining a virtual output with a smaller noise component than the output obtained from a multi-energy X-ray imaging device. In one aspect, utilizing the virtual output algorithm includes obtaining a virtual output with a smaller object-scattered radiation component than the output obtained from a multi-energy X-ray imaging device. In one aspect, some or all of the at least one virtual output generated by the virtual output algorithm is used to correct faulty array pixels, faulty rows, or faulty regions in one or more sensor layers of a multi-layer X-ray imaging device. In yet another aspect, some or all of the at least one virtual output generated by the virtual output algorithm is used to obtain a measurement of bone mineral density or bone mineral areal density.

[0021] In another aspect of this disclosure, an x-ray imaging system is provided for determining at least one virtual output of an x-ray imaging system, the x-ray imaging system comprising: an x-ray source; a multi-energy x-ray imaging device including at least one sensor layer; a processor for receiving a plurality of inputs from the x-ray imaging device and for determining at least one virtual output of the x-ray imaging device, the processor further comprising a computer-readable medium having instructions stored therein, the instructions, if executed, causing the processor to: determine a general algorithm based on an x-ray imaging device application, physical properties of the x-ray imaging device, and / or exposure settings of the x-ray source; substitute the plurality of outputs of the multi-energy x-ray imaging device as inputs into the general algorithm to determine parameters for a virtual output algorithm for the x-ray imaging device and the determined application; and utilize the virtual output algorithm to generate at least one virtual output.

[0022] In another aspect, the multi-energy X-ray imaging device includes a set of sensor layers. In yet another aspect, the multi-energy X-ray imaging device includes at least two sensor layers. In yet another aspect, the multi-energy X-ray imaging device further includes at least one intermediate filter layer between at least two of the at least two sensor layers. In yet another aspect, the intermediate filter layer includes a metallic material filter, a photoconductor layer, or a scintillator layer. In yet another aspect, the multi-energy X-ray imaging device further includes at least one anti-grid layer between at least two of the at least two sensor layers.

[0023] In one aspect, each of at least one sensor layer includes a photoconductor layer or a scintillator layer. In another aspect, the photoconductor layers or scintillator layers of adjacent sensor layers are adjacent to each other. In yet another aspect, at least one of these sensor layers includes a glass substrate layer infused with a scintillator. In still another aspect, at least one of the sensor layers includes a flexible substrate layer and an X-ray absorber. Attached Figure Description

[0024] Embodiments of this disclosure will now be described by way of example only with reference to the accompanying drawings.

[0025] Figure 1 This is a schematic diagram of a three-layer X-ray imaging device;

[0026] Figure 2a This is a schematic diagram of a multi-slice X-ray imaging device. Figure 2a This refers to an X-ray imaging device with two or more layers.

[0027] Figure 2b This is a schematic diagram of a multiplex X-ray imaging device. Figure 2bThis refers to an X-ray imaging system that achieves two or more exposures under different source voltages, currents, and / or filters.

[0028] Figure 3a This is a flowchart outlining a method for determining the virtual image output of a multi-energy X-ray imaging device;

[0029] Figure 3b This is a flowchart outlining a method for determining the virtual layer output of a multi-layer X-ray imaging device;

[0030] Figure 3c This is a flowchart outlining a method for determining the virtual energy output of a multiplex X-ray imaging device;

[0031] Figure 4a It is a graph summarizing the total signal relative to the scintillator filter in the example;

[0032] Figure 4b It is a curve of the instance equation fitting for the sample output of the three-layer detector;

[0033] Figure 5 This is a schematic diagram of an indirect n-layer X-ray imaging device and a direct n-layer X-ray imaging device;

[0034] Figure 6a and Figure 6b These are schematic diagrams of different embodiments of an indirect two-layer X-ray imaging device and a direct two-layer X-ray imaging device;

[0035] Figure 7a and Figure 7b These are schematic diagrams of different embodiments of an indirect 3-layer X-ray imaging device and a direct 3-layer X-ray imaging device;

[0036] Figure 8 A simplified overall diagram of the radiographic imaging environment is shown;

[0037] Figure 9 The structure of a two-dimensional active matrix imaging array is shown;

[0038] Figure 10a This is a schematic diagram of an indirect n-layer X-ray imaging device with intermediate filters between layers, and a direct n-layer X-ray imaging device with intermediate filters between layers.

[0039] Figure 10b and 10c These are schematic diagrams of different embodiments of an indirect 3-layer X-ray imaging device with intermediate filters between some layers and a direct 3-layer X-ray imaging device with intermediate filters between some layers.

[0040] Figure 11aThese are schematic diagrams of an indirect n-layer X-ray imaging device with anti-scattering grids between layers, and a direct n-layer X-ray imaging device with anti-scattering grids between layers; and

[0041] Figure 11b and 11c These are schematic diagrams of different embodiments of an indirect 3-layer X-ray imaging device having an anti-scattering grid between some layers and a direct 3-layer X-ray imaging device having an anti-scattering grid between some layers. Detailed Implementation

[0042] This disclosure relates to a method and apparatus for determining the virtual output of a multi-energy X-ray imaging apparatus. In one embodiment, the method receives actual outputs from layers of a multi-layer X-ray imaging apparatus and then processes those outputs to determine the outputs of other non-existent layers within the multi-layer X-ray imaging apparatus as if they were actual physical layers within the X-ray imaging apparatus. In another embodiment, the method receives actual outputs from different spectral / energy exposures obtained from a multi-energy imaging apparatus and then processes those outputs to determine the outputs of other unobtained spectral / energy exposures.

[0043] Figure 8 A general schematic diagram of a radiographic imaging environment is shown. As illustrated, an X-ray source 10 generates an X-ray beam, or X-ray 11, directed toward an object 12 (e.g., a patient's hand) for imaging via a radiography detector system (RDS) 14. The results of the X-ray exposure can be observed on a computer or processor 16. In this current embodiment, which can be considered an indirect imaging system, the radiography detector system 14 includes a scintillator 15. In a direct imaging system, the X-ray 11 generates a charge within the radiography detector system 14, and a scintillator 15 is not required.

[0044] For some radiographic detector systems 14, synchronization hardware 18 is necessary to obtain correct timing between the x-ray source 10 and the radiographic detector system 14 that is sampling the incident x-ray beam 11. In this disclosure, the radiographic detector system 14 includes a large-area flat panel detector based on active matrix technology to image the object 12.

[0045] Typically, the object 12 to be imaged is placed between the radiation source 10 and the radiographic detector system 14. X-rays 11 passing through the object 12 interact with the radiographic detector system 14. In indirect imaging, the X-rays 11 pass through a fluorescent screen or scintillator 15 (such as structured cesium iodide (Csl), gadolinium oxysulfide (GOS), or calcium tungsten oxide (CaWO₃)). 4Photons are generated during the process. These indirectly generated photons further generate charges within the radiographic detector system 14.

[0046] Figure 9 This is a schematic diagram of a radiographic detector system 14. The RDS 14 includes an active matrix pixel array 20 having a two-dimensional matrix of pixel elements in which charges generated directly or indirectly by incident X-rays are sensed and stored. To access the charge stored at each pixel, gate lines 21 are typically driven sequentially by row switching control 22, such that all pixels in a row output their stored charge to data lines 23, which are coupled to charge amplifiers 24 at the end of each column of the active matrix pixel array 20. Charge amplifiers 24 send the pixel charge data to an analog-to-digital converter (A / D) 26, where the analog signal is converted into a digital representation. The digital representation is then stored in memory 28, awaiting transmission to a computer 16 at a time determined by control logic 29. In addition to their amplification function, the charge amplifiers can also perform multiplexing functions.

[0047] Turn Figure 1 This diagram illustrates a multilayer x-ray imaging detector element or device. In the current embodiment, detector element 14 includes three distinct sensor layers, referred to as a top layer 102, an intermediate or middle layer 104, and a bottom layer 106. As will be understood, in a preferred embodiment, each of the top layer 102, intermediate layer 104, and bottom layer 106 is identical to each other. Each sensor layer can be considered a separate layer of the multilayer x-ray detector element or imaging detector. In one embodiment, each layer may be an amorphous silicon (a-Si) flat panel sensor layer coupled to a scintillator layer. Alternatively, any type of indirect or direct conversion x-ray detection layer may be used for the individual layers. In other embodiments, such as Figure 2a As shown, the detector may include any number of stacked sensor layers (all labeled 102a to 102n, where n can be any number), each stacked sensor layer having its indirect or direct conversion material. During operation, each layer will produce an output that can be used by the methods of this disclosure to obtain further virtual outputs.

[0048] Alternatively, the X-ray imaging device can be part of a multiplex imaging system. In this case, the detector consists of only one sensor layer, but multiple images are obtained by changing the X-ray source properties (such as, but not limited to, kVp and / or filtering) and re-exposure. Each of these images can be considered as an output from the detector, which can then be used by the presented method to obtain further virtual outputs representing other source properties. Figure 2b The diagram shows an X-ray imaging detector used in a multiple imaging system.

[0049] Go to Figure 3a The diagram illustrates a flowchart outlining the basic steps of the method disclosed herein and how it can be used with a multi-energy X-ray imaging device or system to generate at least one virtual output. Figure 3b This is a flowchart outlining a method for determining the output of at least one virtual layer. In this embodiment, the method can be used in an X-ray detector element or X-ray imaging apparatus having two or more sensor layers. In one embodiment, the method and apparatus of this disclosure overcome the challenges of using X-ray detector imaging apparatuses with different X-ray absorber thicknesses. In one embodiment, the method can allow for simpler multilayer detector designs with more versatile and improved multi-energy imaging capabilities.

[0050] First, the X-ray imaging device is exposed to an X-ray source, such that the output from each layer is read out by readout electronics (such as, but not limited to, readout arrays) to a processor. In other words, the system receives input from a multi-energy imaging device (considered layer output), which can be classified as being generated by different X-ray absorption spectra (200).

[0051] Based on the application in which the x-ray imaging device is used, the processor can then input or substitute these inputs into a preferably predetermined or pre-selected general algorithm or equation to determine the virtual output algorithm (204) of the x-ray device. This means calculating or determining parameters for the general algorithm. The general algorithm can be selected based on any of the following: the application of the x-ray imaging device; the physical characteristics of the x-ray imaging device or system; and / or the specific x-ray source setup used in one or more exposures. Once these parameters are calculated, they can be input into or used in the general algorithm to determine or generate the virtual output algorithm. The virtual output algorithm can then be used to calculate the expected (or virtual) output of other virtual layers of the x-ray imaging device, such as images (204).

[0052] To aid in understanding the method, exemplary embodiments of the method are provided. An overview of the amount of signal remaining in the X-ray beam after it has passed through an object and been absorbed by a single infinitely thick scintillator is provided. The amount of signal remaining at any point in the beam path is defined as:

[0053]

[0054] Where Φ(E) is the spectrum of the residual beam. The average scintillator gain function typically takes the form of a common inorganic scintillator.

[0055] As in Figure 4aAs can be seen, the signal decays exponentially as it travels through the absorber. Considering an embodiment of a multilayer detector with layers of the same scintillator material and thickness, in each layer (S... i The signal obtained at each layer can be used to generate an equation that will describe its trend. It is expected that the signal at each layer decreases exponentially, and the rate of this exponential decrease will change as the amount of signal in the bundle decreases. This is because the signal at each layer will be... Figure 4a The difference between the values ​​of two points on the curve shown. Therefore, the equation chosen in this example is:

[0056]

[0057] Wherein, the value l in this equation i This is referred to as the number of layers. Mathematically, the number of layers corresponds to the total scintillator thickness of each layer. However, considering that parameters a, b, and c are fitted, for simplicity, l i Normalized to layer thickness, where l i =1, 2, 3. As will be understood, this is only for simplicity and is not necessary for this method. In fact, l can be modified... i The value is used to account for X-ray losses in detector elements other than scintillators and other non-ideals. By substituting the received output into the general equation or algorithm shown above, the parameters of the virtual output equation can be determined to provide a virtual output equation that can be used to generate any virtual layer of the detector. Figure 4b An example is shown on how, once the parameters of the fitted equation are found, it can be used to approximate the values ​​of the virtual layer.

[0058] Once fitted, the found parameters for each pixel can be used to generate an image of a virtual detector layer with any selected thickness and any selected pre-filtering amount. In this way, a virtual output algorithm for the X-ray imaging device and the application the X-ray imaging device is used for can be found, and then the virtual output algorithm and the application can be used to calculate the values ​​of the virtual layer. For example, an infinitely thick layer can be used... To calculate, or a top layer of half thickness can be used To calculate. Note that even if the virtual output equation directly gives the signal of a layer with the same thickness as those layers used to build the detector, by using the equation intelligently, the value of any desired layer thickness can also be obtained indirectly.

[0059] Therefore, the advantage of the present disclosure lies in the ease of computing virtual multilayer detector elements with any number of layers of arbitrary thickness, even physically impossible detector configurations such as stacked layers or infinitely thick layers. This can be a combination of the benefits or advantages of both dual-energy techniques (where the virtual thickness can be customized to generate the best possible tissue subtraction image) and digital radiography techniques (where image quality can be improved by generating a single virtual layer of impractical thickness or by intelligently reducing noise through more sophisticated fitting methods).

[0060] Go to Figure 3b A flowchart outlining a method for determining the virtual layer output of a multi-layer X-ray imaging apparatus is shown. First, inputs (such as outputs from multi-layer X-ray detectors exposed to X-ray sources) are received from each layer of the multi-layer imaging apparatus (206). These inputs (or outputs) are then used, in place of each pixel, as inputs to a general algorithm to determine the parameters of a virtual output algorithm and generate the virtual output algorithm (208). The virtual output algorithm can then be used to generate a complete or partial image that will be generated by the virtual layers (210).

[0061] Go to Figure 3c A flowchart outlining a method for determining the virtual energy output of a multiplex X-ray imaging device is shown. First, the output of each exposure from the multiplex imaging device is received (212). Then, these outputs are used instead of each pixel as input to a general algorithm to determine the parameters of a virtual output algorithm and generate the virtual output algorithm (214). The virtual output algorithm can then be used to determine a complete or partial image of the virtual exposure (216).

[0062] Although about Figure 3a , Figure 3b or Figure 3c Some mathematical implementations or equations are disclosed to describe signal changes, but any number of equations or algorithms can be used as a general algorithm. These general equations or algorithms may require different numbers of fitting parameters and may have varying fitting quality. Some will fit the input signal precisely, while others can approximate the new signal curve by using the signal as a reference. However, they are all similar because they take the output or energy exposure of different layers as input or signal and the physical information of the detector and its operation, such as the layer scintillator thickness and material, or the different source voltages or filters used.

[0063] Furthermore, it should be noted that while the disclosed embodiments discuss the use of multilayer detectors with all equal absorbers to obtain the necessary fit, other configurations varying the sensor type and thickness are envisioned, and these configurations can improve fit accuracy and allow for more complex fitting algorithms. Even using only two layers, Figure 3bThe flowchart approach can also be useful. Similarly, Figure 3c The methods in the flowchart can be used as follows: Figure 2b The multi-shot switching detector system shown in the diagram allows any number of exposures at different source voltages, currents, and / or filtering states to be used as input to an algorithm that can then generate a virtual exposure image.

[0064] In multilayer detectors with fewer layers and therefore fewer outputs used by a general algorithm, the algorithm's fitting accuracy may be low. However, this can be improved, for example, by using known materials as intermediate filters to spectrally separate the beam spectrum between detector layers, thus allowing for a wider spectral coverage of the signal given to the algorithm. Given the physical configuration of the detector device, a general algorithm can be adapted to any configuration and generate an appropriate virtual output algorithm that allows for the computation of virtual layer signals. Similarly, given the exposure settings (such as voltage, current, and filtering) in a multi-shot imaging system, a general algorithm can be selected to adapt to chosen parameters and generate a virtual output algorithm that allows for the computation of virtual exposure signals.

[0065] The embodiments presented above are examples for illustrating the present technology. As mentioned, the implementation details of the methods disclosed herein can be modified to allow for better results in specific applications or for a given particular detector system. The simplest modification to the provided examples would be to change the general equation or algorithm to another exponentially decreasing equation, such as...

[0066]

[0067] Another example is using a multilayer detector with scintillators of the same or different thicknesses to fit the amount of signal in the beam, rather than the absorbed signal, thus approximating it with a fitting equation. Figure 4a The curve in the curve, and assuming that the signal at each layer is an explicit integral of the curve, can be represented by the general algorithm as follows:

[0068]

[0069] in, It is the thickness of the scintillator pre-filter for each layer, and t i It is the thickness of the scintillator in that layer.

[0070] Furthermore, the method disclosed herein can be modified for use with multilayer detectors containing scintillators of different materials and thicknesses. In this case, the input X-ray spectrum at each pixel can be fitted to a parameterized function. This is possible because it is known that the signal at each layer is proportional to the product of the residual spectrum at each layer and the absorption efficiency of that layer.

[0071] In another embodiment, a multilayer detector with two or more layers can be used, and the obtained signal is used to locate... The best fitting parameters.

[0072] In a further implementation, the dual-layer detector can be used with an intermediate filter made of the same scintillator material, and the signal can be fitted to the equation. However, l is used separately for the top signal and the bottom signal. i =1,5. This still effectively leaves S normalized to the double-layer thickness. i Again, note that l can actually be modified. i The value is taken into account for other detector elements. This implementation can be extended to the previously mentioned implementation that assumes the signal at each layer is an explicit integral of a curve with fitting parameters, but the intermediate scintillator material is taken into account in the choice of integration limit by adding the thickness of the intermediate scintillator material to the limit of the integral of those layers after the intermediate filter in the beam path. This can be achieved by utilizing S i We can further extend this by using different parameterized equations.

[0073] In another embodiment, a four-layer detector can be used and the signal can be fitted to any of the previously mentioned general equations or a new equation with four parameters (such as...). In yet another embodiment, more complex general algorithms, such as minimization algorithms in the form of Monte Carlo minimization algorithms, are also possible.

[0074] These examples make it clear that different types of mathematical methods can be combined with any multi-layer X-ray detector to generate virtual layer signals. It will also be understood that the methods of this disclosure can be extended to any multi-energy detector system, including but not limited to multi-shot imaging systems, where individual image exposures are taken under different source voltages, currents, and / or filters. This method can be adapted to trends between different input spectra and thus allows extrapolation to other input source voltages and a better understanding of the material being imaged. It should be apparent that the approach taken by the methods of this disclosure is equally effective in other applications, such as multispectral 3D computed tomography imaging or real-time imaging.

[0075] Furthermore, the methods disclosed herein can be used to algorithmically transfer information between layers or exposures while maintaining local contrast. This allows for the correction of other problems typically encountered in X-ray imaging, including correcting faulty array pixels, faulty rows, or faulty regions, or reducing electronic or quantum noise. Array fault correction can allow for the relaxation of low or minimal defect density requirements on individual sensor layers. Similar improvements can be achieved for noise reduction, where data from multiple layers or multiple exposures can reduce uncertainties in the measurement of the true signal.

[0076] One way to correct faulty array pixels, faulty rows, or faulty regions in individual sensor layers of a multilayer X-ray detector apparatus using the method disclosed herein is as follows: First, identify individual faulty pixels or all pixels belonging to faulty rows or faulty regions in a sensor layer; obtain outputs corresponding to those pixels or regions from all other sensor layers in the multilayer detector apparatus, where if the value of one layer corresponds to a similar part of the object being imaged, the output from that layer corresponds to the output in another layer; fit these outputs to a general algorithm to generate a virtual output algorithm; use the virtual algorithm to obtain virtual outputs for all faulty pixels or faulty regions to match the physical properties of the original sensor layer; and use this virtual output to replace the values ​​of the faulty pixels in the original sensor layer. Clearly, this method can be reproduced for each sensor layer to remove all faulty pixel values ​​from some or all layers of the multilayer detector apparatus.

[0077] Noise reduction in sensor output data can also be achieved using the methods disclosed herein. This can be accomplished by selecting a general algorithm that requires fewer fitting parameters than the number of layers in a multi-layer imaging device or the exposure in a multi-shot imaging system, or by selecting an algorithm that does not weight all output data in the same way. Once a virtual output algorithm for this general algorithm is found, a virtual output layer or exposure with the same or similar physical properties as one of the device outputs can be generated. Depending on the nature of the selected general algorithm, this virtual output can have similar local contrast to the original device output, but with less noise components. It is also possible to replace only certain regions or spatial frequency components of the original output to achieve better results.

[0078] Another application of the method disclosed herein is for measuring bone mineral density using dual-energy X-ray absorptiometry. The parameters found for the virtual output algorithm, or the generated virtual layer or exposure image, can be used in conjunction with any additional information regarding the X-ray imaging equipment, the exposure settings used, or the configuration of the X-ray system to calculate the density or areal density in some or all of the imaged bone regions.

[0079] A further application of this method is object scattering correction. X-ray radiation typically scatters from the object being imaged, resulting in an overall loss of image quality. The method of this disclosure can utilize the differences in spectral properties in typical object-scattered radiation to separate and remove these differences from the final output image, thereby improving image quality.

[0080] exist Figure 5 (n layers) Figure 6a and Figure 6b (Two-layer approach) and Figure 7a and Figure 7b(Three-layer approach) schematically illustrates different multilayer detectors that can be used with the methods of this disclosure for both indirect scintillator-based X-ray detectors and direct photoconductor-based X-ray approaches. Given the properties of the materials used, a certain amount of scattered or fluorescent radiation (grouped here under the first item) is expected to occur from one layer to another when the detector is exposed, thereby altering the signal output from each layer, which may affect the method presented herein for determining the virtual output.

[0081] like Figure 5 As shown, detector 14 includes "n" sensor layers 500a, 500b, ..., 500n. As will be understood, "n" represents any number. For a direct multilayer X-ray detector, each sensor layer 500 includes a photoconductor layer 502 and a substrate layer 504. For an indirect multilayer X-ray detector, each sensor layer 500 includes a scintillator layer 506 and a substrate layer 508.

[0082] like Figure 6a As shown, the detector includes a first sensor layer 500a, an intermediate filter layer 510, and a second sensor layer 500b. For a direct multilayer X-ray detector, each sensor layer 500 includes a photoconductor layer 502 and a substrate layer 504. In the current embodiment, the intermediate filter layer 510 may be another photoconductor layer 512. For an indirect multilayer X-ray detector, each sensor layer 500 includes a scintillator layer 506 and a substrate layer 508, wherein the intermediate filter layer 510 may be another scintillator layer 514.

[0083] Figure 6b The illustrated embodiments are similar to Figure 6a In one embodiment, the positions of the photoconductor layer 502 and the substrate layer 504 within the sensor layer 500 are swapped (directly), and the positions of the scintillator layer 506 and the substrate layer 508 within the sensor layer 500 are swapped (indirectly).

[0084] like Figure 7a As shown, the detector includes a first sensor layer 500a, a second sensor layer 500b, and a third sensor layer 500c. For a direct multilayer X-ray detector, each sensor layer 500 includes a photoconductor layer 502 and a substrate layer 504. For an indirect multilayer X-ray detector, each sensor layer 500 includes a scintillator layer 506 and a substrate layer 508.

[0085] Figure 7b The embodiments shown are similar to Figure 7aIn one embodiment, an intermediate filter layer is added between the second sensor layer 500b and the third sensor layer 500c. As will be understood, the intermediate filter layer may also be placed between the first sensor layer 500a and the second sensor layer 500b. Alternatively, the intermediate filter layer 510 may be placed between the first and second sensor layers, and between the second and third sensor layers.

[0086] To overcome the challenge of reducing or minimizing radiation scattered by the X-ray absorption layer, various strategies can be employed. One strategy could be to select a material with a low-k edge (such as an amorphous selenium photoconductor), where the k-fluorescent X-rays have energies less than 12 keV and therefore do not travel very far; alternatively, a CsI scintillator with 33 keV fluorescence X-rays could be selected. Similarly, an intermediate filter made of the same material as the selected scintillator can be used to reduce the effect of scattered radiation. Furthermore, the orientation of the sensor layer can be as follows: Figure 6a , Figure 7a and Figure 7b The orientation of the sensor layer 500a is schematically shown to minimize the distance between (directly) photoconductor layers 502 or between (indirectly) scintillator layers 506, thereby reducing the scattering distance associated with X-ray k-fluorescence.

[0087] Other techniques can be used to reduce cross-scattering between layers. For example... Figure 11a , Figure 11b and Figure 11c As shown, this involves adding an antiscattering grid between the sensor layers in any of the previously mentioned configurations. This antiscattering grid will disproportionately absorb scattered radiation and thus reduce the signal value of the layer relative to the scattering (called the scatter-to-primary ratio).

[0088] Figure 11a This is a schematic diagram of a multilayer detector 500 comprising multiple sensor layers 500a, 500b, ..., 500n, where "n" can be any number. An antiscattering grating layer 516 is located between the sensor layers 500. As in previous embodiments, each direct sensor layer includes a photoconductor layer 502 and a substrate layer 504, and each indirect sensor layer includes a scintillator layer 506 and a substrate layer 508.

[0089] Figure 11b This is a schematic diagram of a multilayer detector 500 comprising three (3) sensor layers 500a, 500b and 500c and a single anti-scattering grid layer 516 between the first sensor layer and the second sensor layer. Figure 11cThis is a schematic diagram of a multilayer detector 500 comprising three (3) sensor layers 500a, 500b and 500c and a single anti-scattering grid layer 516 between the second and third sensor layers.

[0090] Similarly, such as Figure 10a , Figure 10b and Figure 10c As shown, an intermediate filter can be added between sensor layers. This intermediate filter will disproportionately absorb scattered photons because these photons are primarily at the low end of the diagnostic X-ray spectrum. The specific material type of the intermediate filter can be selected to tune the absorption of scattered energy. In one embodiment, the material choice for one or more intermediate filters is metallic, such as copper, aluminum, or silver.

[0091] Figure 10a This is a schematic diagram of a multilayer detector 500 comprising multiple sensor layers 500a, 500b, ..., 500n, where "n" can be any number. An intermediate filter layer 518 is located between the sensor layers 500. As in previous embodiments, each direct sensor layer includes a photoconductor layer 502 and a substrate layer 504, and each indirect sensor layer includes a scintillator layer 506 and a substrate layer 508.

[0092] Figure 10b It is a schematic diagram of a multilayer detector 500 comprising three (3) sensor layers 500a, 500b and 500c and a single intermediate filter layer 518 between the first sensor layer and the second sensor layer. Figure 11c It is a schematic diagram of a multilayer detector 500 comprising three (3) sensor layers 500a, 500b and 500c and a single intermediate filter layer 518 between the second sensor layer and the third sensor layer.

[0093] Another technique involves reducing or minimizing the distance between X-ray absorber layers by utilizing the thinnest possible substrate. The thickness of the X-ray absorber layers can be significantly reduced by using a flexible substrate. Finally, this distance can be completely eliminated by combining the substrate and absorber layers in the form of a substrate incorporating a scintillator.

[0094] In the foregoing description, numerous details have been set forth for illustrative purposes in order to provide a comprehensive understanding of these embodiments. However, it will be apparent to those skilled in the art that these specific details may not be necessary. In other instances, well-known structures may be illustrated in block diagram form to avoid obscuring the understanding. For example, specific details are not provided regarding whether elements of the embodiments described herein are implemented as software routines, hardware circuits, firmware, or combinations thereof.

[0095] Embodiments of this disclosure or components thereof may be configured or represented as a computer program product stored in a machine-readable medium (also referred to as a computer-readable medium, processor-readable medium, or computer-usable medium having computer-readable program code embodied therein). A machine-readable medium may be any suitable tangible, non-transitory medium, including magnetic, optical, or electrical storage media (including magnetic disks, compact optical disc read-only memories (CD-ROMs), memory devices (volatile or non-volatile), or similar storage mechanisms). The machine-readable medium may contain different sets of instructions, code sequences, configuration information, or other data that, when executed, cause a processor or controller to perform the steps in the method according to embodiments of this disclosure. Those skilled in the art will appreciate that other instructions and operations necessary for carrying out the described embodiments may also be stored on a machine-readable medium. Instructions stored on a machine-readable medium may be executed by a processor, controller, or other suitable processing means and may be connected to a circuit interface to perform the described tasks.

[0096] The above embodiments are intended to be illustrative only. Changes, modifications, and variations can be made to particular embodiments by those skilled in the art without departing from the scope defined solely by the appended claims.

Claims

1. A method for determining at least one virtual output of a multi-energy X-ray imaging device: Receive multiple outputs generated from multiple different absorbed X-ray spectra from the multi-energy X-ray imaging device; A general algorithm is determined based on the application of the x-ray imaging equipment, the physical properties of the x-ray imaging equipment, and / or the x-ray source exposure settings. The multiple outputs are substituted as inputs into the general algorithm to determine parameters, and a virtual output algorithm is generated for the multi-energy X-ray imaging device and the determined application. as well as The virtual output algorithm is used to generate the at least one virtual output; The general algorithm includes one of the following: For multi-slice X-ray imaging equipment Where a, b, and c are parameters. For the first i Signal at the layer, i For layers, and For the defined number of layers; or For multi-slice X-ray imaging equipment Where b and c are parameters, For the first i Signal at the layer, i For layers, and For the defined number of layers; or For multi-slice X-ray imaging equipment Where a, b, and c are parameters. For the first i Signal at the layer, For the first i The thickness of the scintillator pre-filter layer, i As a layer, For the first i The thickness of the scintillator layer; or Minimization algorithm.

2. The method according to claim 1, wherein, The plurality of outputs received from the multi-energy X-ray imaging device are obtained from some or all of the layers of the multi-energy X-ray imaging device; The multi-energy X-ray imaging device is a single-shot multi-layer X-ray imaging device.

3. The method according to claim 1, wherein, Determining the general algorithm includes: Determine the application of the x-ray imaging equipment, wherein the multi-energy x-ray imaging equipment is being used in the x-ray imaging equipment application; and The general algorithm is selected based on the determined application.

4. The method according to claim 1, wherein, The minimization algorithm is the Monte Carlo minimization algorithm.

5. The method according to claim 1, wherein, The virtual output algorithm includes: A virtual output with a smaller noise component than that obtained from the multi-energy X-ray imaging device is obtained.

6. The method according to claim 1, wherein, The virtual output algorithm includes: A virtual output is obtained that has a smaller component of object-scattered radiation than the output obtained from the multi-energy X-ray imaging device.

7. The method according to claim 2, wherein, Some or all of the at least one virtual output generated by the virtual output algorithm are used to correct faulty array pixels or faulty rows in one or more sensor layers of the multilayer X-ray imaging device.

8. The method according to claim 2, wherein, Some or all of the at least one virtual output generated by the virtual output algorithm are used to correct faulty regions in one or more sensor layers of the multilayer X-ray imaging device.

9. The method according to claim 1, wherein, Some or all of the at least one virtual output generated by the virtual output algorithm are used to obtain bone mineral density measurements.

10. The method according to claim 1, wherein, Some or all of the at least one virtual output generated by the virtual output algorithm are used to obtain bone mineral area density measurements.

11. An x-ray imaging system for determining at least one virtual output of an x-ray imaging system, comprising: X-ray source; A multi-energy X-ray imaging device including at least one sensor layer; A processor configured to receive multiple inputs generated from multiple different absorbed x-ray spectra from the x-ray imaging device, and to determine at least one virtual output of the x-ray imaging device, the processor further comprising a computer-readable medium storing instructions that, if executed, cause the processor to: A general algorithm is determined based on the application of the x-ray imaging equipment, the physical properties of the x-ray imaging equipment, and / or the exposure settings of the x-ray source; The multiple outputs of the multi-energy X-ray imaging device are used as inputs to the general algorithm to determine the parameters of the virtual output algorithm for the X-ray imaging device and the determined application. as well as The virtual output algorithm is used to generate the at least one virtual output; The general algorithm includes one of the following: For multi-slice X-ray imaging equipment Where a, b, and c are parameters. For the first i Signal at the layer, i For layers, and For the defined number of layers; or For multi-slice X-ray imaging equipment Where b and c are parameters, For the first i Signal at the layer, i For layers, and For the defined number of layers; or For multi-slice X-ray imaging equipment Where a, b, and c are parameters. For the first i Signal at the layer, For the first i The thickness of the scintillator pre-filter layer, i As a layer, For the first i The thickness of the scintillator layer; or Minimization algorithm.

12. The x-ray imaging system according to claim 11, wherein, The multi-energy X-ray imaging device includes: At least two sensor layers.

13. The x-ray imaging system according to claim 12, wherein, The multi-energy X-ray imaging device also includes: At least one intermediate filter layer between at least two of the at least two sensor layers.

14. The x-ray imaging system according to claim 13, wherein, The intermediate filter layer includes a metal material filter, a photoconductor layer, or a scintillator layer.

15. The x-ray imaging system according to claim 12, wherein, The multi-energy X-ray imaging device also includes: At least one anti-grid layer between at least two of the at least two sensor layers.

16. The x-ray imaging system according to claim 11, wherein, Each of the at least one sensor layer includes: Photoconductor layer or scintillator layer.

17. The x-ray imaging system according to claim 15, wherein, The photoconductor layers or scintillator layers of adjacent sensor layers are adjacent to each other.

18. The x-ray imaging system according to claim 12, wherein, At least one of the sensor layers includes a glass substrate layer infused with a scintillator.

19. The x-ray imaging system according to claim 12, wherein, At least one of the sensor layers includes a flexible substrate layer and an X-ray absorber.

Images