Method and system for determining a virtual output of a multi-energy radiography apparatus

Abstract

A method and system for determining a virtual energy output of a multi-energy x-ray device.SOLUTION: A method comprising: receiving a plurality of outputs generated by a plurality of different x-ray absorption spectra from a multi-energy x-ray radiography apparatus; determining a general equation based on an application of the multi-energy x-ray radiography apparatus, a physical characteristic of the multi-energy x-ray radiography apparatus, or an exposure setting of an x-ray source; and dynamically generating at least one virtual energy output. The at least one virtual energy output represents an energy output of a virtual layer in the multi-energy radiography apparatus, and the virtual output equation is used to determine the at least one virtual output for the virtual layer.SELECTED DRAWING: Figure 8

Description

[Technical Field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application claims priority from U.S. Provisional Application No. 62 / 682,540, filed June 8, 2018, the contents of which are incorporated herein by reference.

[0002] FIELD OF THE DISCLOSURE This disclosure relates generally to radiography, and more particularly to methods and systems for determining the virtual output of a multi-energy radiography device. [Background technology]

[0003] The quality of a medical image, and therefore its value as a tool, depends solely on how well it conveys the anatomy of the patient being imaged to an observer, such as a physician. The better the anatomy is understood, the more accurate the information a physician has to make decisions.

[0004] In radiography, a significant source of noise that often degrades image quality is anatomical noise. This is caused by the superimposition of normal anatomical structures resulting from two-dimensional (2D) projections of a three-dimensional (3D) patient. This noise can obscure the tissue being imaged or be mistaken for an anatomical abnormality. A simple example of this is a chest radiograph obtained for the purpose of assessing pulmonary anatomy, where the ribs inevitably obstruct the resulting image. In this case, the ribs are not the anatomical structure of interest and are therefore the primary source of anatomical noise.

[0005] A proposed approach to reduce anatomical noise is dual-energy (DE) imaging, which exploits fundamental properties of x-ray-matter interaction, but different tissue types will not only have different mass attenuation coefficients (μ / ρ(E)) across the diagnostic energy range, but will also have different rates of change of these coefficients.

[0006] One challenge with DE imaging is the need to obtain two separate low-energy and high-energy images. To achieve this, the x-ray spectrum absorbed by the detector must be heavily weighted at the low end of the diagnostic range for the low-energy (LE) image and at the high end for the high-energy (HE) image. DE imaging can decompose the patient's projection into a soft-tissue-only image and a hard-tissue-only image. Several mathematical methods exist for obtaining these DE images from the LE and HE inputs, most notably logarithmic subtraction and basis decomposition.

[0007] In practice, perfect cancellation of a particular tissue type is generally impossible. Several factors contribute to the formation of non-ideal scenarios that cannot be captured by mathematical techniques. These include the broad spectrum of x-ray fluences that lead to the formation of each image, as opposed to the ideal source used in mathematical analysis; heterogeneity in the density or mass attenuation coefficients of the tissues being canceled, which makes it impossible to determine the exact values ​​that should be used when calculating weighting factors; and x-ray scattering from both the object being imaged and the detector, which is not accounted for by the Beer-Lambert law. These non-idealities also mean that theoretical values ​​for weighting factors may not provide the best cancellation possible, requiring the observer to calculate these ideal values ​​experimentally or qualitatively.

[0008] In practice, there are two fundamentally different ways to achieve this spectral separation: either the source spectra are different for the two images (hereafter referred to as multi-shot DE imaging), or the detector selectively absorbs different parts of a broader spectrum to form each image (hereafter referred to as single-shot DE imaging). Regardless of the method used, a large separation between the two spectra is essential to obtain high-quality, tissue-selective images.

[0009] One approach to obtaining images at different energies is to vary the spectrum produced by the x-ray tube, rather than as part of the imaging system, and acquire them sequentially in time. This is the concept behind multi-shot imaging (sometimes called kVp switching), in which one image is acquired using a low x-ray tube kVp, followed immediately by a second image at a higher kVp. Because the effective energies of the low-kVp and high-kVp beams are different, the two resulting images contain information primarily obtained at the low and high ends of the x-ray diagnostic spectrum, respectively. Alternatively, instead of changing the source kVp between exposures, source filtration can be changed by quickly moving a spectral filter in and out of the beam path. This has the effect of presenting two different spectra at the detector, provided the source filtration is selective across the energy spectrum.

[0010] This approach can also be extended to multi-energy images by acquiring multiple sequential images at different kVp values ​​or source filters, which allows for more spectral information to be obtained and improved images can then be generated algorithmically.

[0011] Unfortunately, the temporal separation inherent in this technique can introduce motion artifacts into the final images, presenting significant challenges to the radiologist or observer interpreting them. These artifacts are noticeable distortions of the image caused by slight misalignments of anatomical structures in successive images, typically resulting from patient or object motion during and between image acquisitions.

[0012] Ideally, the source tube voltage could be changed instantaneously so that the next exposure could begin immediately after one was completed. However, currently available sources require an interval of at least 150–200 ms between successive exposures. This is because, to obtain the ideal relative intensity of the image, not only voltage changes but also changes in tube current are required. While this interval is short enough for most patients to refrain from significant movement, cardiac, respiratory, and minor muscle movements will still occur during the interval. These movements result in motion artifacts, which can be particularly problematic in cardiopulmonary imaging due to the prominent cardiac presence. Furthermore, this problem becomes more severe as more image acquisitions are added in multi-energy imaging, lengthening the total acquisition time and increasing patient movement.

[0013] An alternative method for obtaining multi-energy images exists, commonly referred to as single-shot imaging. This method takes the opposite approach to multi-shot imaging, achieving spectral separation within the detector rather than at the source. This is achieved by vertically stacking two sensor layers to form a two-layer detector, known as a sandwich configuration. One layer, such as the top layer, absorbs primarily LE X-rays, while the second or bottom layer absorbs HE X-rays. Therefore, to allow for a larger spectrum covering both LE and HE X-rays, this technique, performed at a higher kVp, requires only a single exposure. This method has since been extended to multi-layer detectors, allowing for multiple images with increasing effective energy in subsequent stacks.

[0014] A practical problem that arises with the single-shot approach is that the mass loading (or equivalently, their thickness) of the sensitive materials - be they scintillators or direct conversion materials - must be tailored to the specific tissue type and patient anatomy in order to obtain ideal and effective energy separation between the layers. At a commercial level, it is only feasible to build a few specific configurations, leaving as the only practical solution a compromise solution that best suits all of the target applications and patient types.

[0015] Accordingly, a novel method and apparatus is provided to mitigate or overcome at least one drawback of the above-described imaging methods and apparatus. Summary of the Invention

[0016] In one aspect of the present disclosure, there is provided a method for determining at least one virtual output for a multi-energy radiography device, including receiving a plurality of outputs from a multi-energy radiography device generated by different x-ray spectra; determining a generic algorithm based on an application of the radiography device, physical characteristics of the radiography device, or exposure settings of an x-ray source; determining parameters; substituting the plurality of outputs as inputs into the generic algorithm to generate a virtual output algorithm for the multi-energy radiography device and the determined application; and utilizing the virtual output algorithm to generate at least one virtual output.

[0017] In another aspect, the plurality of outputs received from the multi-energy radiography device are obtained from some or all layers of the multi-energy radiography device, and the multi-energy radiography device is a single-shot multi-layer radiography device. In a further aspect, the plurality of outputs received from the multi-energy radiography device are obtained from two or more x-ray exposures taken with different x-ray source exposure settings, and the multi-energy radiography device is a multi-shot radiography device. In another aspect, the x-ray source exposure settings include source voltage, source current, or source filtration. In yet another aspect, determining the general algorithm includes determining an x-ray application for which the multi-energy radiography device is being used and selecting the general algorithm based on the determined application.

[0018] In another aspect, selecting the general algorithm includes selecting a general algorithm for a multi-layer radiography device.

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[0019] In yet another aspect, utilizing the virtual output algorithm includes obtaining a virtual output having a smaller noise component than an output obtained from a multi-energy radiography device. In one aspect, utilizing the virtual output algorithm includes obtaining a virtual output having a smaller object scatter component than an output obtained from a multi-energy radiography device. In one aspect, some or all of the at least one virtual output generated by the virtual output algorithm is used to correct for defective array pixels, lines, or regions in one or more sensor layers of a multi-layer radiography 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 area density.

[0020] In another aspect of the present disclosure, there is provided an X-ray imaging system for determining at least one virtual output of an X-ray imaging system, the X-ray imaging system including: an X-ray source; a multi-energy X-ray imaging device including at least one sensor layer; and a processor for receiving a plurality of inputs from the X-ray imaging device and determining at least one virtual output of the X-ray imaging device, the processor further including a computer readable medium having stored thereon instructions that, when executed, cause the processor to determine a generic algorithm based on an application of the X-ray imaging device, physical characteristics of the X-ray imaging device, and / or exposure settings of the X-ray source, to substitute the plurality of outputs of the multi-energy X-ray imaging device as inputs into the generic algorithm to determine parameters of a virtual output algorithm for the X-ray imaging device and the determined application, and to generate the at least one virtual output using the virtual output algorithm.

[0021] In another aspect, the multi-energy X-ray imaging device includes a pair 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.

[0022] In one embodiment, each of the at least one sensor layer includes a photoconductor layer or a scintillator layer. In another embodiment, the photoconductor or scintillator layers of adjacent sensor layers are adjacent to each other. In a further embodiment, at least one of the sensor layers includes a glass substrate layer doped with a scintillator. In yet another embodiment, at least one of the sensor layers includes a flexible substrate layer and an x-ray absorber.

[0023] Embodiments of the present disclosure will now be described, by way of example only, with reference to the accompanying figures. [Brief explanation of the drawings]

[0024] [Figure 1] FIG. 1 is a schematic diagram of a three-layer X-ray imaging device. [Figure 2a] 1 is a schematic diagram of a multi-layer radiography device, depicting an radiography device having two or more layers. [Figure 2b] 1 is a schematic diagram of a multi-shot radiography device, representing an radiography system in which two or more exposures are obtained with different source voltages, currents, and / or filtering. [Figure 3a] 1 is a flowchart outlining a method for determining a virtual image output of a multi-energy radiography device. [Figure 3b] 1 is a flowchart outlining a method for determining virtual layer output of a multi-layer radiography device. [Figure 3c]1 is a flowchart outlining a method for determining the virtual energy output of a multi-shot radiography device. [Figure 4a] 1 is a graph outlining an example of total signal versus scintillator filtration. [Figure 4b] 10 is a graph of an example equation fit to the sample output of a three-layer detector. [Figure 5] 1 is a schematic diagram of an indirect n-layer X-ray imaging device and a direct n-layer X-ray imaging device. [Figure 6a] 1A-1C are schematic diagrams of different embodiments of an indirect bilevel radiography device and a direct bilevel radiography device. [Figure 6b] 1A-1C are schematic diagrams of different embodiments of an indirect bilevel radiography device and a direct bilevel radiography device. [Figure 7a] 1A-1C are schematic diagrams of different embodiments of an indirect triple-layer radiography device and a direct triple-layer radiography device. [Figure 7b] 1A-1C are schematic diagrams of different embodiments of an indirect triple-layer radiography device and a direct triple-layer radiography device. [Figure 8] A typical diagram of an X-ray environment is shown. [Figure 9] 1 shows a two-dimensional active matrix imaging array structure. [Figure 10a] 1 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. [Figure 10b] 1A and 1B are schematic diagrams of different embodiments of an indirect triple layer radiography device with intermediate filters between some layers and a direct triple layer radiography device with intermediate filters between some layers. [Figure 10c] 1A and 1B are schematic diagrams of different embodiments of an indirect triple layer radiography device with intermediate filters between some layers and a direct triple layer radiography device with intermediate filters between some layers. [Figure 11a] 1 is a schematic diagram of an indirect n-layer radiography device with anti-scatter grids between layers, and a direct n-layer radiography device with anti-scatter grids between layers. [Figure 11b]1A-1C are schematic diagrams of different embodiments of an indirect triple layer radiography device with anti-scatter grids between some layers and a direct triple layer radiography device with anti-scatter grids between some layers. [Figure 11c] 1A-1C are schematic diagrams of different embodiments of an indirect triple layer radiography device with anti-scatter grids between some layers and a direct triple layer radiography device with anti-scatter grids between some layers. DETAILED DESCRIPTION OF THE INVENTION

[0025] The present disclosure relates to a method and apparatus for determining the virtual output of a multi-energy radiography device. In one embodiment, the method receives actual output from layers of a multi-energy radiography device and then processes the output to determine the output of other, non-existent layers in the multi-energy radiography device as if they were actual physical layers in the radiography device. In another embodiment, the method receives actual output from different spectral / energy exposures obtained from a multi-shot radiography device and then processes the output to determine the output of other, non-existent spectral / energy exposures.

[0026] 8 shows a general diagram of an X-ray imaging environment. As shown, an X-ray source 10 generates X-ray beams, i.e., X-rays 11, that are transmitted toward an object 12, e.g., a patient's hand, for imaging by an X-ray detector system (RDS) 14. The results of the X-ray exposure are viewable on a computer or processor 16. In the current embodiment, which may be considered a fluoroscopic system, the X-ray detector system 14 includes a scintillator 15. In a direct imaging system, the X-rays 11 generate electrical charges within the X-ray detector system 14, eliminating the need for a scintillator 15.

[0027] Some X-ray detector systems 14 require synchronization hardware 18 to achieve precise timing between the X-ray source 10 and the X-ray detector system 14 sampling the impinging X-ray beam 11. In this disclosure, the X-ray detector system 14 includes a large area flat panel detector based on active matrix technology to achieve imaging of the object 12.

[0028] Typically, an object 12 to be imaged is positioned between an X-ray source 10 and an X-ray detector system 14. X-rays 11 passing through the object 12 interact with the X-ray detector system 14. In fluorography, the X-rays 11 generate photons as they pass through a phosphor screen or scintillator 15, such as structured cesium iodide (CsI), gadolinium oxysulfide (GOS), or calcium tungsten oxide (CaWO4). These indirectly generated photons then generate further charge within the X-ray detector system 14.

[0029] FIG. 9 is a schematic diagram of the X-ray detector system 14. The RDS 14 includes an active matrix pixel array 20 having a two-dimensional matrix of pixel elements, where charge generated directly or indirectly by incident X-rays is sensed and stored. To access the charge accumulated at each pixel, gate lines 21 are typically driven sequentially by row switching control 22, causing all pixels in a row to output their accumulated charge onto data lines 23 coupled to charge amplifiers 24 at the end of each column of the active matrix pixel array 20. The charge amplifiers 24 send the pixel charge data to an analog-to-digital converter (A / D) 26, which converts the analog signal to a digital representation. The digital representation is then stored in memory 28 awaiting transmission to the computer 16 at a time determined by control logic 29. The charge amplifiers may also perform multiplexing functions in addition to their amplification function.

[0030] Referring to FIG. 1, a schematic diagram of a multilayer radiography detector element or device is shown. In this embodiment, the detector element 14 includes three distinct sensor layers, seen as a top layer 102, a middle or central layer 104, and a bottom layer 106. As will be appreciated, in a preferred embodiment, each of the top layer 102, middle layer 104, and bottom layer 106 are identical to one another. Each of the sensor layers can be seen as an individual layer of the multilayer radiography detector element or 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 in the individual layers. In another embodiment, shown in FIG. 2a, the detector may include any number of stacked sensor layers (all labeled 102a through 102n, where n can be any number), each with its own indirect or direct conversion material. In operation, each layer generates an output that can be used by the methods of the present disclosure to derive further virtual outputs.

[0031] Alternatively, the radiography device may be part of a multi-shot radiography system. In this case, the detector contains only one sensor layer, but multiple images are obtained by re-exposing the detector to different x-ray source characteristics (e.g., but not limited to, kVp and / or filtering). Each of these images can be considered an output from the detector and then used in the presented method to obtain additional virtual outputs representing other source characteristics. A schematic diagram of a radiography detector for use in a multi-shot radiography system is shown in Figure 2b.

[0032] Referring to FIG. 3a, a flowchart outlining the basic steps of the disclosed method and its use in a multi-energy X-ray imaging device or system to generate at least one virtual output is shown. FIG. 3b is a flowchart outlining a method for determining the output of at least one virtual layer. In this embodiment, the method may be used with X-ray detector elements or X-ray imaging devices having two or more sensor layers. In one embodiment, the disclosed method and device overcomes the challenges of using X-ray detector imaging devices with different X-ray absorber thicknesses. In one embodiment, the method may enable the design of simpler multi-layer detectors with greater versatility and improved multi-energy imaging capabilities.

[0033] First, the x-ray imager is exposed to an x-ray source such that the output from each layer is read out to a processor by readout electronics, such as, but not limited to, a readout array. In other words, the system receives inputs (viewed as layer outputs) from a multi-energy imager that may be categorized as being produced by different x-ray absorption spectra (200).

[0034] Next, based on the application for which the radiography device is being used, the processor can preferably input or substitute inputs into a predetermined or preselected generic algorithm or equation to determine a virtual output algorithm for the radiography device (204). This means calculating or determining parameters for the generic algorithm. The generic algorithm may be selected based on either the application for the radiography device, the physical characteristics of the radiography device or system, and / or the settings of the specific x-ray source used in one or more exposures. Once these parameters are calculated, they may be input into or used in the generic algorithm to determine or generate a virtual output algorithm. The virtual output algorithm can then be used to calculate expected (or virtual) outputs, such as images, for other virtual layers of the radiography device (204).

[0035] To aid in understanding this method, an exemplary embodiment of the method is provided, which provides an overview of the amount of signal remaining after an x-ray beam has passed through an object and has been absorbed by a single, infinitely thick scintillator. The amount of signal remaining at any point in the beam path is given by

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[0036] As shown in Figure 4a, the signal decays exponentially as it passes through the absorber. By considering a multi-layer detector embodiment with layers of the same scintillator material and thickness, the signal (s) obtained in each layer can be calculated. i ) can be used to generate an equation that describes the trend. The signal at each layer is expected to decay exponentially, and the rate of exponential decay is expected to change as the amount of signal in the beam decreases. This is because the signal at each layer is the difference between the values ​​of two points on the curve shown in Figure 4a. Therefore, the equation chosen for this example is

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[0037] Once fitted, the discovered parameters for each pixel can be used to generate an image of a virtual detector layer of any chosen thickness and with any chosen amount of pre-filtering. In this way, a virtual output algorithm for the radiography device and the application in which the radiography device is being used can be found, which can then be used to calculate the values ​​of the virtual layer. For example, an infinitely thick underlayer would be

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[0038] Thus, an advantage of the present disclosure is that it facilitates the calculation of virtual multi-layer detector elements with any number of layers of any thickness, as well as physically impossible detector configurations such as overlapping layers or infinitely thick layers. This can be a boon or advantage for both dual-energy techniques, where the virtual thickness can be adjusted to produce the best possible tissue-removal image, and digital radiography techniques, where image quality may be improved by generating a single impractically thick virtual layer or by intelligently reducing noise using more complex fitting methods.

[0039] 3b, a flowchart outlining a method for determining virtual layer outputs of a multi-layer radiography device is shown. First, inputs (such as outputs from multi-layer x-ray detectors exposed to an x-ray source) are received from each layer of the multi-layer radiography device (206). These inputs (or outputs) are then substituted as inputs, pixel by pixel, into a general algorithm to determine parameters for the virtual output algorithm and generate the virtual output algorithm (208). The virtual output algorithm can then be used to generate a full or partial image that would be generated by the virtual layer (210).

[0040] Referring to Figure 3c, a flowchart is shown outlining a method for determining the virtual energy output of a multi-shot radiography device. First, the outputs from each exposure of the multi-shot radiography device are received (212). These outputs are then substituted as inputs into a general algorithm for each pixel to determine parameters for the virtual output algorithm and generate the virtual output algorithm (214). The virtual output algorithm can then be used to determine a full or partial image of the virtual exposure (216).

[0041] While several mathematical implementations or equations describing signal changes are disclosed with respect to Figures 3a, 3b, or 3c, any number of equations or algorithms may be used as a general algorithm. These general equations or algorithms may require different numbers of fitting parameters and may have different quality fits. Some may fit the input signal exactly, while others may use the signal as a reference to approximate a new signal curve. However, they are all similar in that they receive as input or signals the output or energy exposure of different layers, as well as physical information about the detector and its operation, such as the thickness and material of the scintillator layers, or different source voltages and filtering used.

[0042] Furthermore, while the disclosed embodiments discuss using a multi-layer detector with all equal absorbers to obtain the required fit, it should be noted that other configurations of various sensor types and thicknesses are contemplated and may improve the accuracy of the fit and enable more sophisticated fitting algorithms. The method of the flowchart in FIG. 3b may be beneficial even when only two layers are used. Similarly, the method of the flowchart in FIG. 3c may be used with the multi-shot switching detector system shown in FIG. 2b, where any number of exposures at different source voltages, currents, and / or filtration may be used as input to an algorithm capable of generating a virtual exposure image.

[0043] Multilayer detectors with fewer layers and therefore fewer outputs available to the general algorithm may result in lower algorithm fitting accuracy. However, this can be improved by, for example, using known materials as intermediate filters to spectrally separate the beam spectrum between detector layers and broaden the spectral range of the signal to the algorithm. As long as the physical configuration of the detector device is known, the general algorithm can be adapted to generate an appropriate virtual output algorithm that corresponds to any configuration and allows for the calculation of a virtual layer signal. Similarly, as long as the exposure settings (voltage, current, filtration, etc.) are known in a multi-shot imaging system, the general algorithm can be selected to generate a virtual output algorithm that corresponds to the selected parameters and allows for the calculation of a virtual exposure signal.

[0044] The embodiments presented above are examples that help to illustrate this approach. As previously mentioned, the details of embodiments of the disclosed methods can be modified to obtain better results in a particular application or given a particular detector system. The simplest modification to the examples provided is to change the general equation or algorithm to:

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[0045] Another example is to use a multi-layer detector with scintillators of the same or different thicknesses, and by fitting to the amount of signal in the beam rather than the absorbed signal, the curve in FIG. 4a can be approximated with a fitting equation, and by assuming that the signal in each layer is a definite integral of the curve, a general algorithm can be obtained, for example, as follows:

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[0046] Furthermore, the disclosed method can be modified for use with multi-layer detectors having scintillators of both different materials and thicknesses. In this case, the input X-ray spectrum at each pixel can be fit to a parameterized function. This is possible because the signal in each layer is known to be proportional to the product of the spectrum remaining in each layer and the absorption efficiency of the layer.

[0047] In another embodiment, a multi-layer detector with two or more layers may be used, and the resulting signal may be used to:

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[0048] In a further embodiment, a two-layer detector may be used with an intermediate filter made of the same scintillator material, dividing the signal by the equation

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[0049] In another embodiment, a four-layer detector is used to fit the signal to one of the general equations above, or

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[0050] Through these examples, it is clear that different types of mathematical methods can be used in conjunction with any multi-layer X-ray detector to generate virtual layer signals. The disclosed method can be extended to any multi-energy detector system, including but not limited to multi-shot imaging systems, where separate image exposures are performed with different source voltages, currents, and / or filtering. This method can fit trends between different input spectra, allowing for extrapolation to other input source voltages and further understanding of the material being imaged. Clearly, the approach adopted by the disclosed method is equally valid for additional applications, such as multispectral 3D computed tomography or real-time imaging.

[0051] Furthermore, the disclosed methods can be used to algorithmically transfer information between layers or exposures while maintaining local contrast. This allows for correction of defective array pixels, lines, or regions, and for correction of other issues commonly encountered in radiography, such as reducing electronic or quantum noise. Array defect correction allows for relaxation of the requirements for low or minimum defect density in individual sensor layers. Similar improvements can be achieved in noise reduction, allowing data from multiple layers or multiple exposures to reduce the uncertainty in the measurement of true signals.

[0052] One method for correcting defective array pixels, lines, or regions in individual sensor layers of a multi-layer X-ray detector device using the method of the present disclosure involves first identifying each defective pixel, or all pixels belonging to the defective line or region, in one sensor layer, obtaining outputs corresponding to these pixels or regions from all other sensor layers in the multi-layer detector device, where outputs from one layer correspond to outputs from another layer if their values ​​correspond to a similar portion of the object being imaged, and fitting these outputs to a general algorithm to generate a virtual output algorithm, using the virtual algorithm to obtain virtual outputs for all defective pixels or regions that match the physical characteristics of the original sensor layer, and replacing the values ​​of the defective pixels in the original sensor layer with the virtual outputs. It will be apparent that this method can be reproduced for each individual sensor layer to remove all defective pixel values ​​from some or all layers of the multi-layer detector device.

[0053] Noise reduction of sensor output data may be achieved by utilizing the methods of the present disclosure. This may be achieved by selecting a general algorithm that requires fewer fitting parameters than the number of layers in a multi-layer imaging device or the number of exposures in a multi-shot imaging system, or by selecting an algorithm that does not equally weight all output data. Once a virtual output algorithm for this general algorithm is found, a virtual output layer or exposure can be generated that has the same or similar physical characteristics as one of the device outputs. Depending on the nature of the selected general algorithm, this virtual output may have similar local contrast to the original device output, but with a smaller noise component. It may also be possible to replace only specific regions or spatial frequency components of the original output to achieve better results.

[0054] One further application of the disclosed method is the measurement of bone mineral density by dual-energy X-ray absorptiometry. Either the parameters specified for the virtual output algorithm or the generated virtual layer or exposure images may be used in combination with additional information about the X-ray imaging device, the exposure settings used, or the X-ray system configuration to calculate the density or areal density in some or all regions of the imaged bone.

[0055] A further application of this method of the present disclosure is object scatter correction. X-ray radiation is typically scattered from the object being imaged, resulting in overall poor image quality. The differences in the spectral characteristics of typical object scatter can be exploited by the method of the present disclosure to separate and therefore remove it from the final output image, thereby improving image quality.

[0056] Different multi-layer detectors that can be used in the methods of the present disclosure are shown schematically in Figure 5 (n-layer), Figures 6a and 6b (two-layer approach), and Figures 7a and 7b (three-layer approach) for both indirect scintillator-based X-ray detectors and direct photoconductor-based X-ray approaches. Given the nature of the materials used, it is expected that when the detector is exposed to light, there will be some scattering or fluorescence emission (here grouped under the first term) from one layer to another, which may change the signal output from each layer and affect the method of determining the virtual output presented herein.

[0057] 5, the detector 14 includes "n" sensor layers 500a, 500b, ..., 500n. It is understood that "n" represents any number. In the case of a direct multi-layer X-ray detector, each sensor layer 500 includes a photoconductor layer 502 and a substrate layer 504. In the case of an indirect multi-layer X-ray detector, each sensor layer 500 includes a scintillator layer 506 and a substrate layer 508.

[0058] 6a, the detector includes a first sensor layer 500a, an intermediate filter layer 510, and a second sensor layer 500b. In the case of a direct multi-layer X-ray detector, each sensor layer 500 includes a photoconductor layer 502 and a substrate layer 504. In this embodiment, the intermediate filter layer 510 may be another photoconductor layer 512. In the case of an indirect multi-layer X-ray detector, each sensor layer 500 includes a scintillator layer 506 and a substrate layer 508, where the intermediate filter layer 510 may be another scintillator layer 514.

[0059] The embodiment shown in Figure 6b is similar to the embodiment of Figure 6a, with the positions of the photoconductor layer 502 and substrate layer 504 (direct) and the positions of the scintillator layer 506 and substrate layer 508 (indirect) switched within the sensor layer 500.

[0060] 7a, the detector includes a first sensor layer 500a, a second sensor layer 500b, and a third sensor layer 500c. In the case of a direct multi-layer X-ray detector, each sensor layer 500 includes a photoconductor layer 502 and a substrate layer 504. In the case of an indirect multi-layer X-ray detector, each sensor layer 500 includes a scintillator layer 506 and a substrate layer 508.

[0061] The embodiment shown in Figure 7b is similar to the embodiment of Figure 7a, with the addition of an intermediate filter layer between the second sensor layer 500b and the third sensor layer 500c. As will be appreciated, an intermediate filter layer may also be disposed between the first sensor layer 500a and the second sensor layer 500b. Alternatively, an intermediate filter layer 510 may be disposed between both the first and second sensor layers and the second and third sensor layers.

[0062] Various strategies can be employed to overcome the challenge of reducing or minimizing radiation scattered by the X-ray absorption layer. One strategy is to select a material with a low-k absorption edge (such as an amorphous selenium photoconductor) whose k-fluorescence X-ray energy is less than 12 keV and therefore does not travel far. Alternatively, a CsI scintillator with a 33 keV x-ray fluorescence can be used to reduce the effects of scattered radiation. Furthermore, the orientation of the sensor layer can be modified as shown schematically in Figures 6a, 7a, and 7b, where the sensor layer 500a is oriented to minimize the distance between it and the photoconductor layer 502 (direct) or the scintillator layer 506 (indirect), thereby reducing the scattering distance associated with X-ray k-fluorescence.

[0063] Further techniques can be used to reduce cross-scattering between layers, including adding an anti-scatter grid between the sensor layers in any of the configurations described above, as shown in Figures 11a, 11b, and 11c, which disproportionately absorbs scattered radiation and therefore reduces the proportion of the layer signal value that corresponds to scatter (known as the scatter-to-primary dose ratio).

[0064] 11a is a schematic diagram of a multi-layer detector 500 including multiple sensor layers 500a, 500b, ..., 500n, where "n" can be any number. Located between the sensor layers 500 is an anti-scatter grid layer 516. As with the previous embodiment, 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.

[0065] Figure 11b is a schematic diagram of a multi-layer detector 500 including three sensor layers 500a, 500b, and 500c and a single anti-scatter grid layer 516 between the first and second sensor layers. Figure 11c is a schematic diagram of a multi-layer detector 500 including three sensor layers 500a, 500b, and 500c and a single anti-scatter grid layer 516 between the second and third sensor layers.

[0066] Similarly, intermediate filters may be added between sensor layers, as shown in Figures 10a, 10b, and 10c, which are primarily at the low-energy end of the diagnostic x-ray spectrum and will therefore disproportionately absorb scattered photons. The specific material type of the intermediate filter may be selected to tailor the scattered energy absorption. In one embodiment, the material choice for one or more intermediate filters is a metal, such as copper, aluminum, or silver.

[0067] 10a is a schematic diagram of a multi-layer detector 500 including multiple sensor layers 500a, 500b, ..., 500n, where "n" can be any number. Located between the sensor layers 500 is an intermediate filter layer 518. As with the previous embodiment, 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.

[0068] Figure 10b is a schematic diagram of a multi-layer detector 500 including three sensor layers 500a, 500b, and 500c and a single intermediate filter layer 518 between the first and second sensor layers. Figure 11c is a schematic diagram of a multi-layer detector 500 including three sensor layers 500a, 500b, and 500c and a single intermediate filter layer 518 between the second and third sensor layers.

[0069] Another approach is to reduce or minimize the distance between the X-ray absorber layers by utilizing the thinnest possible substrates; their thickness can be significantly reduced by using flexible substrates. Finally, this distance can be completely eliminated by combining the substrate layer with the absorber layer, in the form of a scintillator-infused substrate.

[0070] In the foregoing description, for purposes of explanation, numerous details are set forth in order to provide a thorough understanding of the embodiments. However, it will be apparent to one skilled in the art that these specific details may not be required. In other instances, well-known structures may be shown in block diagram form in order to avoid obscuring the understanding. For example, specific details are not provided regarding whether elements of the embodiments described herein are implemented as a software routine, a hardware circuit, firmware, or a combination thereof.

[0071] Embodiments of the present disclosure, or components thereof, may be provided or expressed as a computer program product stored on a machine-readable medium (also referred to as a computer-readable medium, a processor-readable medium, or a computer-usable medium having computer-readable program code embodied therein). The machine-readable medium may be any suitable tangible, non-transitory medium, including a diskette, a compact disk read-only memory (CD-ROM), a magnetic, optical, or electrical storage medium, including memory devices (volatile or non-volatile), or similar storage mechanisms. The machine-readable medium may include various sets of instructions, code sequences, configuration information, or other data that, when executed, cause a processor or controller to perform method steps according to an embodiment of the present disclosure. Those skilled in the art will appreciate that other instructions and operations necessary to implement the described embodiments may also be stored on the machine-readable medium. The instructions stored on the machine-readable medium may be executed by a processor, controller, or other suitable processing device and may interface with circuitry to perform the described tasks.

[0072] The above-described embodiments are intended to be examples only. Alterations, modifications, and variations may be effected to the particular embodiments by those of skill in the art without departing from the scope, which is defined solely by the claims appended hereto.

Claims

1. 1. A method for determining at least one virtual energy output for a multi-energy x-ray imaging device, comprising: receiving a plurality of outputs from the multi-energy X-ray imaging device produced by a plurality of different X-ray absorption spectra; determining a general equation based on an application of the multi-energy X-ray imaging device, physical characteristics of the multi-energy X-ray imaging device, or an exposure setting of an X-ray source, the general equation representing the output of the plurality of different X-ray absorption spectra; Substituting the plurality of outputs as inputs into the general equation to determine parameters and generate a virtual energy output equation for the multi-energy radiography device and the determined application; dynamically generating the at least one virtual energy output utilizing the virtual output equation and at least one input based on x-ray spectral characteristics; the at least one virtual energy output represents an energy output of a virtual layer within the multi-energy radiography device; The method, wherein the virtual output equation is used to determine the at least one virtual output for the virtual layer.

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

3. the plurality of outputs received from the multi-energy X-ray imaging device are obtained from two or more X-ray exposures taken with different X-ray source exposure settings; the multi-energy X-ray imaging device is a multi-shot X-ray imaging device; The method of claim 1.

4. The method of claim 3 , wherein the exposure settings of the X-ray source comprise a source voltage, a source current, or a source filter.

5. Determining the general equation determining an x-ray application for which the multi-energy x-ray imaging device is being used; and selecting the general equation based on the determined application. The method of claim 1 , comprising:

6. Selecting the general equation As the general equation, for a multi-layer X-ray imaging device, [Equation 1] (where a, b, c are parameters, s ｉ is the signal of each layer, l ｉ is the defined layer number) The method of claim 5 , comprising:

7. Selecting the general equation As the general equation, for a multi-layer X-ray imaging device, [Equation 2] (where b and c are parameters, s ｉ is the signal of each layer, l ｉ is the defined layer number) The method of claim 5 , comprising:

8. Selecting the general equation As the general equation, for a multi-layer X-ray imaging device [Equation 3] (where b and c are parameters, [Equation 4] is the pre-filtering thickness of each scintillator layer, t ｉ is the thickness of the scintillator layer) The method of claim 5 , comprising:

9. Determining the general equation selecting a minimization equation as the general equation; The method of claim 5 , comprising:

10. Utilizing the virtual output equation obtaining an output having a noise component smaller than a virtual output obtained from the multi-energy X-ray imaging device; The method of claim 1 , comprising:

11. Utilizing the virtual output equation obtaining an output having a smaller object scatter component than a virtual output obtained from the multi-energy X-ray imaging device; The method of claim 1 , comprising:

12. 3. The method of claim 2, wherein some or all of the at least one virtual output generated by the virtual output equation is used to correct defective array pixels, lines, or regions in one or more sensor layers of the single-shot multi-layer radiography device.

13. The method of claim 1 , wherein some or all of the at least one virtual output generated by the virtual output equation is used to obtain a measurement of bone mineral density or bone mineral areal density.

14. 1. An X-ray imaging system for determining at least one virtual output for 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 multi-energy X-ray imaging device generated by a plurality of different X-ray absorption spectra and for determining at least one virtual output of the multi-energy X-ray imaging device, the processor, when executed, causing the processor to: determining a general equation based on an application of the multi-energy X-ray imaging device, physical characteristics of the multi-energy X-ray imaging device, and / or exposure settings of the X-ray source, the general equation representing the output of the plurality of different X-ray absorption spectra; Substituting a plurality of outputs of the multi-energy X-ray imaging device as inputs into the general equation to determine parameters of a virtual output equation for the multi-energy X-ray imaging device and the determined application; a processor further including a computer readable medium having stored thereon instructions for generating the at least one virtual output utilizing the virtual output equation and at least one input based on x-ray spectral characteristics. Preparation, the at least one virtual output represents an output of a virtual layer within the multi-layer radiography device; The virtual output equation is used to determine the at least one virtual output for the virtual layer.

15. The multi-energy X-ray imaging device A series of sensor layers The X-ray imaging system according to claim 14,

16. The multi-energy X-ray imaging device At least two sensor layers The X-ray imaging system according to claim 15,

17. The multi-energy X-ray imaging device at least one intermediate filter layer between at least two of the at least two sensor layers; The X-ray imaging system of claim 16 further comprising:

18. 20. The radiography system of claim 17, wherein the intermediate filter layer comprises a metallic material filter, a photoconductor layer, or a scintillator layer.

19. The multi-energy X-ray imaging device At least one anti-grid layer is disposed between at least two of the at least two sensor layers. The X-ray imaging system of claim 16 further comprising:

20. Each of the at least one sensor layer Photoconductor layer or scintillator layer The X-ray imaging system according to claim 15,

21. 21. The radiography system of claim 20, wherein the photoconductor or scintillator layers of adjacent sensor layers are adjacent to each other.

22. 17. The radiography system of claim 16, wherein at least one of the sensor layers comprises a scintillator-infused glass substrate layer.

23. 17. The radiography system of claim 16, wherein at least one of the sensor layers comprises a flexible substrate layer and an x-ray absorber.

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