X-ray detector with x-ray fluorescent components

Abstract

X-ray detector efficiency and / or resolution are improved by using x-ray fluorescent materials in the x-ray detector design. The x-ray fluorescent components capture incident x-rays and emit fluorescent x-rays that are detected by adjacent x-ray detector elements. A component of the x-ray detector may be made x-ray fluorescent by incorporating a high-Z element-containing material into the x-ray detector component design. The x-ray fluorescent detector component may be an interpixel reflecting septa, a backing layer, and / or an interdetector reflector for multilayer detectors.

Description

X-RAY DETECTOR WITH X-RAY FLUORESCENT COMPONENTSCROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims the benefit of U.S. Provisional Patent Application Serial No. 63 / 579,868, filed on August 31. 2023, and entitled "X-RAY DETECTOR WITH X-RAY FLUORESCENT COMPONENTS,” which is herein incorporated by reference in its entirety.BACKGROUND

[0002] Conventional energy-integrating detectors in diagnostic CT systems contain interpixel reflecting septa between pixels in order to optically isolate them from one another. The typical pitch of these pixels has been fixed at approximately 1 mm for multiple decades, with minimal improvements in the in-plane direction. This is because the thickness of these reflecting septa has not been improved. The reflecting septa create dead space that has limited the spatial resolution of these detectors. Further shrinking of pixel pitch, without corresponding reductions of reflecting septa width, would increase the amount of radiation dose needed to acquire equivalent quality data.SUMMARY OF THE DISCLOSURE

[0003] The present disclosure addresses the aforementioned drawbacks by providing an x-ray detector that includes a plurality of x-ray detector elements to detect x-rays impinging on the each of the plurality of x-ray detector elements, and a plurality of interpixel reflecting septa, each of the plurality of interpixel reflecting septa being arranged between adjacent ones of the plurality’ of x-ray detector elements. The plurality of interpixel reflecting septa are composed of an optically reflective material that also emits a fluorescent x-ray when irradiated by an incident photon.

[0004] It is another aspect of the present disclosure to provide an x-ray detector that includes a semiconductor layer to convert incident x-rays into electrical charge signals, a plurality of anodes to measure the electrical charge signals, and a backing layer to capture x- rays passing through the semiconductor layer and return the captured x-rays back into the semiconductor layer.

[0005] It is yet another aspect of the present disclosure to provide an x-ray detector that includes a plurality’ of x-ray detector layers and a plurality of x-ray fluorescent layers, where each of the plurality of x-ray fluorescent layers is arranged between at least two of the pluralityof x-ray detector layers. X-rays incident on one of the plurality of x-ray fluorescent layers generate fluorescent x-rays that are subsequently detected by an adjacent one of the plurality of x-ray detector layers.

[0006] The foregoing and other aspects and advantages of the present disclosure will appear from the following description. In the description, reference is made to the accompanying drawings that form a part hereof, and in which there is shown by way of illustration one or more embodiments. These embodiments do not necessarily represent the full scope of the invention, however, and reference is therefore made to the claims and herein for interpreting the scope of the invention.BRIEF DESCRIPTION OF THE DRAWINGS

[0007] FIG. 1 illustrates an example x-ray detector having a plurality of interpixel reflecting septa that emit fluorescent x-rays when irradiated by incident x-rays.

[0008] FIG. 2 illustrates an example x-ray detector having a backing layer to reflect photons and / or emit fluorescent x-rays back into the x-ray detector layer.

[0009] FIG. 3 illustrates an example x-ray detector having a backing layer composed of two different layers to reflect photons and / or emit fluorescent x-rays back into the x-ray detector layer.

[0010] FIG. 4 illustrates an example multilayer x-ray detector having a plurality of x- ray detector layers interspersed with a plurality' of x-ray fluorescent layers that emit fluorescent x-rays back into adjacent ones of the x-ray detector layers when irradiated by incident x-rays.

[0011] FIG. 5 shows relative variance of detectors w ith reflectors loaded with different elements, under the simplified model. Lines starting at “1 mm’’ and ‘‘0.5 mm” correspond to fill factors of 81% and 64%, respectively. “Probability of XRF” is probability of x-ray fluorescence for a photon incident on the reflector. Lines terminate at a circle when this probability is equal to the fraction of the spectrum above the k-edge.

[0012] FIG. 6 shows absorbed energy distribution s(E) of the interpixel reflecting septa (labeled here as “reflectors”) and the scintillator under irradiation by monoenergetic 80 keV photons. Note that the energy' absorbed within the septa would not be integrated in a real energy -integrating detector. Left column shows with gadolinium loading for x-ray fluorescence, right column shows without. Top row shows x-rays incident at the center of a reflecting septa only, bottom row shows x-rays uniformly distributed over the whole pixel. Gd Kacharacteristic x-rays are 42 keV. Pixel pitch is 0.5 mm and fill factor is 64%.

[0013] FIG. 7 shows variance of the detector at different pitches and Gd loading factors at 120 kVp, relative to an ideal single-bin photon-counting detector.

[0014] FIG. 8 shows sensitivity of the detector variance to different Gd loading factors at 120 kVp. Improvements are limited beyond 0.5 g / cm3of Gd and saturate near 1 g / cm3.

[0015] FIG. 9 shows variance at different kVp for different loading elements, normalized to an ideal photon-counting detector. Pixel pitch is 0.5 mm, and 1.0 g / cm3of each element is added. All spectra were filtered by 30 cm of water. "PMMA" is acrylic only reflecting septa. Additionally, the septa, including the PMMA example, may include a means to reflect optical photons and therefore reduce crosstalk between pixels. In this example, it is assumed that the material used to isolate pixels constitutes a small percentage of total atoms and their contribution to the x-ray attenuation properties is neglected here.

[0016] FIGS. 10A and 10B illustrate an example CT system that can implement the x- ray fluorescent x-ray detector designs described in the present disclosure.DETAILED DESCRIPTION

[0017] Described here are systems and method for improving the resolution and / or stopping power of x-ray detectors using x-ray fluorescent materials in the x-ray detector design to capture incident x-rays and emit fluorescent x-rays that are detected by adjacent x-ray detector elements, thereby improving the dose efficiency and / or resolution of the x-ray detector. In general, any component of the x-ray detector may be made x-ray fluorescent by incorporating a high-Z element-containing material into the x-ray detector component design. While all elements have some probability of x-ray fluorescence, elements with low atomic number have a smaller probability and generate very low energy fluorescence x-rays. Carbon, for example, has a fluorescence yield <1% and is not considered “fluorescent” here. As a nonlimiting example, a high-Z element may include an element wi th an atomic number of at least 58. For example, elements with an atomic number in the range of 58 to 69, inclusive, may be referred to as high-Z elements. Cerium, with an atomic number of 58, has a fluorescence yield of >90% and is considered “fluorescent.”

[0018] As one non-limiting example, the x-ray detector may be an energy-integrating detector whose resolution may be improved without sacrificing dose efficiency from the fill factor loss. In some implementations, interpixel reflecting septa arranged between adjacent x- ray detector elements are composed of the x-ray fluorescent material. The x-ray fluorescent materials may be composed of a high-Z material, such as a high-Z element mixed into the septamaterial. This high-Z element will emit characteristic x-rays (i.e., fluorescent x-rays) when bombarded by radiation and some of those characteristic x-rays may then be absorbed in adjacent x-ray detector elements, thereby improving the resolution of the x-ray detector. Additionally or alternatively, x-ray detector components other than interpixel reflecting septa can be composed of high-Z materials, such as a metallic backing layer arranged behind the x- ray detector elements, or metallic fluorescence layers arranged between layers of a multi-layer detector. Alternatively, the x-ray detector may be an energy-discriminating detector, such as a photon-counting detector.

[0019] The arrival of photons onto an x-ray detector is a Poisson process. The signal that each photon deposits into the detector is a random variable that depends on the photon’s energy, location, and the random interactions with the detector material. Some possible models for the energy deposited are briefly described. For an ideal single-bin photon-counting detector the probability distribution function of the signal deposited, s(E), is a delta function; that is, all photons increment the counter once. For an ideal energy -integrating detector, s(E) is equal to the incident spectrum; that is. all photons add their kinetic energy to the photodiode signal. For an energy-integrating detector that has a fill factor, F, but is otherwise ideal, s(E) has a probability' F to be sampled from the incident spectrum, and the remaining probability' (1 — F) to be zero.

[0020] The total integrated signal follows a compound Poisson distribution. For a rate A and time duration t, the expected number of arriving photons is At, and the mean signal is A.t E(s(E)), where E(x) is the expected value. The variance is At E(s(E)2). Detector output may be normalized to the same mean during the process of gain calibration. After normalization, the variance is:

[0021] As a convention, the expected number of arriving photons may be selected as At = 1. This has the effect of normalizing all reported variances to an ideal photon-counting detector, which achieves cr2= 1.

[0022] When a high-Z, x-ray fluorescent element is loaded into the septa, the signal probability distribution function s(E) is changed. In particular, the probability’ of sampling a zero is reduced and is replaced with the probability of sampling the energy of a characteristic x-ray. The high-Z element may be mixed or otherwise loaded into the septa material, or acompound containing the high-Z element may be mixed or otherwise loaded into the septa material. Additionally or alternatively, the high-Z material (e.g., high-Z element, compound containing a high-Z element) may be coupled to the septa, such as by depositing a thin film or layer of high-Z material on the septa, or the like.

[0023] It is an aspect of the present disclosure to load, mix, or otherwise couple a high- Z element or material to interpixel reflecting septa or other components of an x-ray detector. In this way, the x-ray detector components may become x-ray fluorescent: absorbing some of the incident x-ray photons and re-emitting them as characteristic x-rays to be detected in adjacent detector elements or pixels. This yields partial recovery of detector dead space and increases the effective fill factor. These improvements can also enable higher spatial resolution, or to make more dose efficient detectors.

[0024] Thus, using the techniques described in the present disclosure, dose efficiency can be improved at the same pixel size or dose efficiency can be maintained while reducing the pixel size. In general, one or more components of an x-ray detector may be composed of material containing a high-Z (i.e., high atomic number) element (e.g., gadolinium (Gd). cerium (Ce), thulium (Tm), tungsten (W), or the like). As one example, the high-Z material may be incorporated into interpixel reflecting septa in order to make them fluorescent to x-rays (in addition to their existing role of individual detector elements, usually by redirecting visible light), so that more x-rays can be detected by the detector. Additionally or alternatively, the high-Z material may be incorporated into one or more metallic sheets arranged after the detector, to recapture x-ray photons that penetrate or otherwise pass through the detector elements and convert at least some of those captured x-ray photons to fluorescent x-rays, some of which are emitted back to the detector. In still other examples, the high-Z material may be incorporated into one or more metallic sheets arranged between the layers of a multi-layer detector, fluorescent to x-rays.

[0025] FIG. 1 illustrates an example x-ray detector 10 that includes a plurality of x-ray detector elements 12 and interpixel reflecting septa 14 between adjacent x-ray detector elements 12. The septa are referred to as interpixel reflecting septa 14 because their usual purpose is to isolate detector elements from each other (i.e., to reduce crosstalk) and they are engineered to reflect the visible light that is emitted from the scintillator. Visible light here may refer to the light that is emitted from the scintillator material in response to x-ray irradiation, which is often in the visible range but may extend into infrared or ultraviolet regimes. Visible light photons may be differentiated from the x-ray photons that are emitted from the x-raysource, which are of much greater energy. As described above, the interpixel reflecting septa 14 are composed of a high-Z material, or otherwise incorporate a high-Z material or element, so as to generate fluorescent x-rays from x-rays that are incident on the septa 14. In the illustrated example, the x-ray detector 10 is an energy-integrating detector. In other examples, the x-ray detector 10 may take other forms, including a photon-counting detector.

[0026] Energy -integrating detectors are pixelated into a plurality of x-ray detector elements 12 composed of a scintillator 16 that converts x-rays into visible light and a photodiode 18 or other suitable light sensor (e.g., a photomultiplier tube (PMT)) that receives the visible light generated by the scintillator 16 and converts the visible light into an electrical signal that is a measurement of the x-rays impinging on the x-ray detector element 12. Interpixel reflecting septa 14 are arranged between adjacent x-ray detector elements 12 and coated with a reflective material (e g., a reflective paint or metal sheet) or otherwise incorporate or are composed of a reflective material to trap the visible light. In general, about 20% of the total x-ray detector 10 area is lost to the interpixel reflecting septa 14, and upwards of 36% of the area may be lost for higher-resolution detectors.

[0027] It is an advantage of the systems and methods described in the present disclosure that by mixing high-Z materials or elements into the interpixel reflecting septa 14, some of the dead area can be recovered because the high-Z atoms can absorb x-rays and re-emit them (e.g., by fluorescence) for capture in nearby scintillators 16. It is contemplated that CT resolution can be improved by about 25-40% using this technique. Additionally or alternatively, in a photon counting detector with a semiconductor detecting layer that does not provide sufficient x-ray attenuation, adding a backing layer containing a high-Z material can allow recapture (that is, absorption of x-ray photons followed by characteristic emission of a fluorescent photon in an isotropic direction) of some photons back into the detector. The improvement in stopping power may be about 5%.

[0028] Thus, as one example, the interpixel reflecting septa 14 can be coated in a reflective paint or other coating intended to reflect visible light, and which has also been doped with a high-Z element or material (e.g., a high-Z element-containing compound) to fluoresce x-rays. For instance, a high-Z element-containing powder can be mixed with a reflective paint epoxy, which may then be applied to the surface of the interpixel reflecting septa 14 or may be incorporated into the bulk of the septa, in cases where the septa are composed of a material that is otherwise transparent to visible light. As described above, the high-Z element may be gadolinium (Gd), cerium (Ce), thulium (Tm), tungsten (W), or the like. Alternatively, theinterpixel reflecting septa 14 may be composed of a reflective high-Z material, such as a reflective metal alloy containing one or more high-Z elements, or could include a reflective pigment including a high-Z element, such as lead-based paints. For example, the interpixel detectors 14 could be composed of a metal alloy composed of 90% aluminum and 10% high- Z element (e.g., Gd) or high-Z element containing compound. In still other examples, the interpixel reflecting septa 14 may be composed of a material into which a high-Z element or material is mixed. For instance, the interpixel reflecting septa 14 may be composed of a polymer that contains both a high-Z material and an optically reflective substance, such as titanium dioxide.

[0029] As shown in FIG. 2, in a photon-counting detector, the x-ray detector elements 12 may be composed of solid-state x-ray detector elements or pixels. For example, a photoncounting detector may include a semiconductor layer 20 composed of a semiconductor material (e.g., CdTe) that converts x-rays into electric charges that can be measured by an x-ray detector layer (e.g., by an array of pixelated anodes 22) to count the number of x-ray photons incident on the x-ray detector 10. Unlike energy-integrating detectors, which stop 99% or more of incident x-ray photons, CdTe sensors of 2 mm thickness may only stop approximately 90% of a 140 kVp spectrum fdtered by 25 cm of water, and this kVp is often used to improve spectral separation. Increasing the thickness of the CdTe layer may stop more incident x-ray photons, but would further increase charge sharing and therefore is not used as a viable solution for improving photon-counting detector efficiency. Thicker CdTe layers may also increase the cost of the x-ray detector.

[0030] It is another aspect of the present disclosure that the efficiency of photoncounting detectors can be improved by using a backing layer 24, which acts as a guardrail, placed beyond the x-ray detector layer 22 to capture punch-through photons. Advantageously, the backing layer 24 is easy to integrate on an existing PCD system and with little risk. It is desirable to have a radiation shield behind the PCD to reduce leakage radiation to the room. As an added advantage, the backing layer 24 can be used as both a radiation shield and to recapture incident photons back into the semiconductor layer 20 to be counted by the x-ray detector layer 22. As a result, quantum efficiency can be improved by several percent because the effective stopping power is improved.

[0031] A photon with energy E that punches through a CdTe sensor interacts with a high-Z backing layer 24 at a depth z following the exponential distributionwhere / rEis the linear attenuation coefficient of the backing layer 24. Assuming E is above the k-edge.most interactions will result in photoelectric absorption and emission of an isotropic characteristic x-ray. Some will re-enter the CdTe sensor and be reabsorbed. The probability of reabsorption can be calculated by Monte Carlo simulations. For 100 keV photons incident on a lead backing layer 24 as a pencil beam, the probability of recapture is 24%, and the average distance between the pencil beam and the point of recapture is 0.5 mm, implying that resolution loss is small.

[0032] Recapture can be improved with a multi-material backing layer. Consider lead, which emits characteristic x-rays of about 85 keV. If such a characteristic x-ray is randomly emitted away from the sensor, it could be separately recaptured in a second layer of tungsten, with k-edge at 70 keV. Some of these photons could then be given a second chance to be reemitted upwards, back towards the sensor.

[0033] The resolution loss noted above will only impact recaptured photons, which compose a minority of the signal. The magnitude of the location error leading to this resolution loss depends on the distance from the semiconductor sensor. If the backing layer is directly beneath the semiconductor sensor, the error in location will be smaller. If the backing layer is separated by some distance, the error will grow larger. In either case, the question of whether it is better to detect the photon at the incorrect location via the backing layer or not can be answered by considering the frequency-dependent detective quantum efficiency (DQE).

[0034] Consider a one-dimensional numerical model where 10% of the photons are either undetected (i.e.. no backing layer) or detected at the wrong location (i.e.. with a backing layer), with error following an exponential distribution with characteristic distance of 0.5 mm. The backing layer improves the DQE at zero frequency by 10%, but may decrease DQE at higher frequencies by 10%. The balance point where DQE is neutral was experimentally observed to be approximately 3.3 line pairs per cm. This suggests that the backing layer is advantageous for detecting low-contrast lesions that are a few mm in diameter or larger, but may be disadvantageous for classifying certain structures that are sub-mm in width. It is worth noting that the recaptured photons will only be at certain, characteristic energies. A detector with sufficient spectral resolution can additionally or alternatively be used to isolate these recaptured events and use them only at lower frequencies through a frequency blending technique. In these instances, the impact of reduced DQE at higher frequencies can be eliminated.

[0035] As one example, shown in FIG. 3, the backing layer 24 may implement a 2- material design: a front layer 26 adjacent the plurality of anodes 22 and a backing layer 28arranged after the front layer 26. thereby sandwiching the front layer 26 between the plurality of anodes 22 and the backing layer 28. Assuming a 140 kVp spectrum and 30 cm water thickness, the thickness of the back layer 28 may be selected as 3 mm (i.e., thick enough to stop nearly all photons), and the elemental composition of the back layer 28 can be selected to match the characteristic x-rays of the front layer 26. As a non-limiting example, the backing layer 24 can be constructed to recapture 40% of incident photons, improving the effective stopping power of the PCD from 90% to 94%. In other examples, the backing layer 24 may be composed of more than two layers, such as three or more layers.

[0036] Additionally or alternatively, a signal may also originate from holes and this signal may drift towards the cathode. In this way, the x-ray detector 10 may include a plurality of cathodes in addition or as an alternative to the plurality of anodes 22.

[0037] As shown in FIG. 4, in some embodiments the x-ray detector 10 may be a multilayer x-ray detector. In these instances, the x-ray detector 10 is composed of a plurality of x-ray detector layers 30. For instance, the x-ray detector layers 30 may be energy -integrating detector layers, energy-discriminating detector layers, or combinations thereof. A plurality of x-ray fluorescent layers 32 are arranged between the x-ray detector layers 30. In this way, when an x-ray punches through a particular x-ray detector layer 30 it has a probability of being absorbed by the subsequent x-ray fluorescent layer, from which a fluorescent x-ray is then emitted into an adjacent one of the x-ray detector layers 32. The incident x-ray may otherwise pass through the fluorescent x-ray layer 32 to the next x-ray detector layer 30. Accordingly, the overall efficiency of the x-ray detector 10 is improved by capturing a higher percentage of incident x-rays that may otherwise punch through the x-ray detector 10 in its entirety. In the illustrated embodiment, the last layer in the x-ray detector 10 is an x-ray fluorescent layer 32, similar to the backing layer embodiments shown in FIGS. 2 and 3. In other examples, however, the final layer in the x-ray detector may be an x-ray detector layer 30, a radiation shielding layer, or other such layer. Similar to the embodiment shown in FIG. 3, in some examples one or more of the x-ray fluorescent layers 32 may be composed of two or more layers. The orientation of the layers 30 and 32 need not be perpendicular to the incident x-ray beam, but may be aligned to be parallel or at an angle to the incident x-ray beam.

[0038] In some embodiments, the backing layer may serve multiple functions. Most photon counting detectors integrate a circuit placed directly behind the semiconductor sensor. The circuit can be made thin, such as with a thickness of approximately 1 mm. Even when the circuit is thin, it can generate heat that must be removed from the device. Most circuits todayare coupled directly with some kind of thermal management mechanism, such as a heatsink or a substrate with high thermal conductivity. These thermal management mechanisms are directly bonded to the circuit to provide good thermal conductivity.

[0039] In some embodiments, these thermal management mechanisms can be modified to have their material incorporate fluorescent elements such that the thermal management mechanisms serve the dual purpose of providing the fluorescent backing layer. Heatsinks today are often made of copper, but tungsten-based heatsinks are available. Advantageously, tungsten is an adequate element for fluorescence. In other embodiments, certain ceramic substrates may also be used. As one non-limiting example, the ceramic substrate may incorporate tungsten carbide into the ceramic. Additionally or alternatively, other alloys that have high atomic number and good thermal conductivity may be used. As an example, Rose’s metal has been used as a solder and as a heat transfer medium, and could be used in the thermal management mechanisms described above. Rose’s metal contains approximately 50% bismuth and 25% lead. Both of these elements are excellent from a fluorescence standpoint. A thin layer of Rose’s metal could provide enough fluorescence to improve the detection efficiency of the detector.

[0040] The atomic number, Z, of the element loaded into the septa affects s(F) in two competing ways. First, the k-edge of the element determines the minimum energy of photons that are eligible for x-ray fluorescence. Second, the characteristic x-rays of the element determine the amount of signal that would be contributed. According to Eqn. (1), it is desirable for the characteristic x-rays to be near the center of the incident spectrum, but characteristic x- rays are always less energetic than the k-edge, which makes many incident photons ineligible for x-ray fluorescence. The optimal choice of atomic number reflects a balance between these two effects and it is also dependent on the incident spectrum. Therefore, a high-Z element can be selected as an element that works reasonably well across the range of tube voltages that are encountered in clinical practice.

[0041] As one example, numerical experiments using a simplified model can be conducted to choose an atomic number. In these instances, the signal probability distribution function s(E) can be estimated using incident spectra of 80, 120, or 140 kVp after filtration by 30 cm of water. In one example study, fill factors, F, of either 81% or 64%, corresponding to pixel pitches of 1.0 or 0.5 mm, respectively, were considered together with a 0.1 mm septa width. If a photon is incident onto the active area of the pixel (with probability equal to fill factor), the energy of the photon will be integrated perfectly. Otherwise, it can be assumed thatthe photon will be incident onto the septa, and the photon can be assumed to undergo photoelectric effect and re-emission of a characteristic x-ray with some variable probability, p. This variable probability, p, can be allowed to vary from zero to the fraction of the spectrum above the k-edge. Note that this upper limit may not be achievable in practice because not all interactions above the k-edge may lead to the emission of a characteristic x-ray (i.e., photons might interact with other shells or by Compton scattering). Also, in some instances characteristic x-rays may not be reabsorbed within the septa. Modeling these effects requires more comprehensive Monte Carlo simulations, such as by using GEANT4.

[0042] GEANT4 is a set of general purpose Monte Carlo codes that track the passage of particles through matter. In an example Monte Carlo simulation for selecting a high-Z material, an array of 10 x 10 pixels were modeled, but only a single pixel touching the center vertex was irradiated with photons. The size of the irradiation field was set to be equal to the pixel pitch. Other pixels were included in the model only to capture intradetector scattering, which was not expected to extend further than a few millimeters. In this example Monte Carlo simulation, each pixel was composed of gadolinium oxysulfide, with density and chemical composition taken from the NIST tables. Other components of the pixel, including the photodiode, were not modeled. In particular, the anti-scatter grid was not modeled. The antiscatter grid presents another limit to the fill factor.

[0043] Each pixel is bordered by interpixel reflecting septa, which may be assumed to be 0. 1 mm in width in most instances. In some examples, the active ingredient of these septa may be TiO2, CnCh. or the like, which are pigments that reflect light from the scintillator and thereby reduce depth-dependent optical Swank noise. To hold the septa in place, a binder is used to provide structure, and the binder (not the pigment) occupies most of the volume of the septa. As a non-limiting example, the binding agent may be a plastic, such as acrylic (e.g.. PMMA) of density 1.19 g / cm3or a transparent epoxy. In an example Mote Carlo simulation, it was assumed that the high-Z element can be superimposed on top of an existing binder with attenuation characteristics comparable to acrylic, and the simulation did not model reductions in the density of acrylic when adding new material.

[0044] FIG. 5 shows the results of a simplified numerical model, where photons that land on the scintillator are perfectly integrated and photons that land in the septa have some probability to undergo x-ray fluorescence. Four elements were chosen with k-edges that are approximately 10 keV apart. A perfect photon-counting detector with 100% fill factor would have a variance of 1.0 on these plots. An energy -integrating detector with 100% fill factor isalso marked and has higher variance because of the non-uniform (energy weighting) of the photons, which has sometimes been described as radiation Swank noise.

[0045] FIG. 5 shows that for any fixed probability of x-ray fluorescence, elements with higher k-edges (among the four choices shown here) yield lower noise because the emitted energy is closer to the average energy of the spectrum. However, lower atomic numbers can capture and re-emit more of the spectrum. Among our examined choices. Ce (k-edge, 40 keV) has the best maximum performance at 80 kVp and Tm (k-edge, 59 keV) has the best performance at 140 kVp. Gd (k-edge, 50 keV) is the best choice at 120 kVp and is the best averaged over all three spectra. Conveniently, Gd is also a familiar choice to x-ray physicists because of its use in scintillators and contrast agents, but this is purely coincidental: elements adjacent to Gd on the periodic table would work similarly well. FIG. 5 shows that a variety of high-Z materials could be used and would be effective at reducing the variance of the detector. While there are some differences between the four elements chosen here, all of them are effective compared to a detector that does not use any of them.

[0046] FIG. 5 shows that a detector loaded with Gd at a pixel pitch of 0.5 mm has the potential to outperform a detector without Gd at a pixel pitch of 1.0 mm, but FIG. 5 does not model all losses that occur, which will be included in our GEANT4 simulations.

[0047] FIG. 6 show s the energy deposited by photons incident on either a point at the center of septa, or distributed across an entire pixel. To make the spectral effects clear in this plot, only monoenergetic 80 keV radiation was used. On the top left of this figure, it can be observed that with Gd loading, approximately half the photons incident on the septa result in characteristic x-rays, either Kaor Kp lines, that are deposited in adjacent scintillator. On average, an 80 keV photon incident on the septa deposits 24 keV of energy into the scintillator. Two sources of losses are present: (1) -30% of the photons are nearly fully absorbed within the scintillator, either because the characteristic x-ray does not escape the septa or because of competing photelectric effect from the L-shell; and (2) -20% of the photons penetrate the septa without depositing energy. This is consistent with the linear attenuation coefficient of Gd at 80 keV, which is 5.57 cm2 / g. The Beer-Lambert law predicts a 18% chance of non-interaction. On the top right, it can be seen that without Gd loading, nearly all the photons penetrate the septa. On the bottom right, it is observed that without Gd loading, the pixel has a -35% chance of not receiving any energy, consistent with the 64% fill factor, whereas with Gd loading on the bottom left, the pixel has a -18% chance of not receiving any energy. Overall, half the photons that normally contribute zero signal because they are incident on the septa can be recoveredand contribute some signal, albeit at reduced energy. K-escape effects can also be seen: without Gd, there is a -10% probability of k-escape. With Gd, this probability increases slightly, probably because of k-escape to adjacent pixels that is now blocked by the increased attenuation of the septa.

[0048] From the probability distribution functions in FIG. 6, it is possible to calculate using Eqn. (1) that an energy-integrating detector without Gd has variance 1.61 times greater than an ideal detector, whereas the detector with 1 g / cm3Gd has variance 1.32 times greater. While the geometric fill factor with Gd loading remains unchanged at 64%, the quantum noise penalty is reduced almost in half by x-ray fluorescence.

[0049] FIGS. 7 and 8 illustrate the effect of pixel pitch and Gd loading concentration on the variance of the detector under Monte Carlo simulation studies. These figures illustrate the same kind of information, but FIG. 7 uses finer granularity in pixel pitch while FIG. 8 uses finer granularity in Gd concentration. FIG. 7 shows that at 1 g / cm3of Gd, a detector with pixel pitch of 0.6 mm can attain the same noise level as a conventional detector with pixel pitch of 1.0 mm without Gd. We reiterate that the anti-scatter grid is not modeled here, and this will decrease the relative advantage of Gd. FIG. 8 predicts that most of the benefit of Gd loading is achieved at 0.5 g / cm3. Beyond 1 g / cm3, the incremental losses from septa absorption of characteristic x-rays within the septa appear to cancel out incremental improvements in x-ray attenuation.

[0050] X-ray fluorescence in the septa could reduce resolution, because x-rays incident on the septa could be redirected as characteristic x-rays to adjacent pixels. At a pitch of 0.5 mm, when a single pixel (meaning a square region 0.5 mm wide that contains the whole scintillator pixel as well as half of each septum that borders it) is irradiated uniformly, the average signal received by the four Cartesian neighbors is 2.9% of the signal received by the irradiated pixel. This neighboring pixel signal is due to travel of characteristic x-rays away from the irradiated pixel, which usually transmits through the acrylic septa. With 1 g / cm3of Gd, the average signal received by the four Cartesian neighbors increases to 4.3% of the signal received by the irradiated pixel. A formal analysis of the frequency-dependent modulation transfer function is outside the scope of this work, but these preliminary results suggest that the incremental impact of Gd loading on spatial resolution is small.

[0051] FIG. 9 shows the variance at different kVp. The variance of the detector with acrylic only septa increases slightly with kVp, probably because of increases in radiation Swank noise. Among the three elements compared here, Gd is optimal or nearly optimal at allkVp stations except for 80 kVp, where Ce shows slightly better performance. While we used 120 kVp for the simulations shown here because it is the most typical kVp station in CT. FIG. 9 shows that we can expect slightly better results at lower kVp and slightly worse results at higher kVp.

[0052] Referring particularly now to FIGS. 10A and 10B, an example of an x-ray computed tomography (“CT”) imaging system 100 is illustrated. The CT system includes a gantry 102, to which at least one x-ray source 104 is coupled. The x-ray source 104 projects an x-ray beam 106, which may be a fan-beam or cone-beam of x-rays, towards a detector array 108 on the opposite side of the gantry 102. The detector array 108 includes a number of x-ray detector elements 110. Together, the x-ray detector elements 110 sense the projected x-rays 106 that pass through a subject 112, such as a medical patient or an object undergoing examination, that is positioned in the CT system 100. Each x-ray detector element 110 produces an electrical signal that may represent the intensity of an impinging x-ray beam and, hence, the attenuation of the beam as it passes through the subject 112. In some configurations, each x- ray detector 110 is capable of counting the number of x-ray photons that impinge upon the detector 110. During a scan to acquire x-ray proj ection data, the g an tty 102 and the components mounted thereon rotate about a center of rotation 114 located within the CT system 100.

[0053] The CT system 100 also includes an operator workstation 116, which ty pically includes a display 118; one or more input devices 120, such as a keyboard and mouse; and a computer processor 122. The computer processor 122 may include a commercially available programmable machine running a commercially available operating system. The operator workstation 116 provides the operator interface that enables scanning control parameters to be entered into the CT system 100. In general, the operator workstation 116 is in communication with a data store server 124 and an image reconstruction system 126. By way of example, the operator workstation 116, data store sever 124, and image reconstruction system 126 may be connected via a communication system 128, which may include any suitable network connection, whether wired, wireless, or a combination of both. As an example, the communication system 128 may include both proprietary or dedicated networks, as well as open networks, such as the internet.

[0054] The operator workstation 116 is also in communication with a control system 130 that controls operation of the CT system 100. The control system 130 generally includes an x-ray controller 132, a table controller 134, a gantry controller 136. and a data acquisition system 138. The x-ray controller 132 provides power and timing signals to the x-ray source104 and the gantry controller 136 controls the rotational speed and position of the gantry 102. The table controller 134 controls a table 140 to position the subject 112 in the gantry 102 of the CT system 100.

[0055] The DAS 138 samples data from the detector elements 110 and converts the data to digital signals for subsequent processing. For instance, digitized x-ray data is communicated from the DAS 138 to the data store server 124. The image reconstruction system 126 then retrieves the x-ray data from the data store server 124 and reconstructs an image therefrom. In some implementations, the image reconstruction system 126 may perform an image reconstruction in which knowledge of the x-ray fluorescence generated using the x-ray detectors 10 descirbed in the present disclosure is used to improve signal, such as by applying frequency weighting. As one example, a notch filter may be used (e.g., as part of the DAS 138) to identify fluorescence events. The image reconstruction system 126 may include a commercially available computer processor, or may be a highly parallel computer architecture, such as a system that includes multiple-core processors and massively parallel, high-density computing devices. Optionally, image reconstruction can also be performed on the processor 122 in the operator workstation 116. Reconstructed images can then be communicated back to the data store server 124 for storage or to the operator workstation 116 to be displayed to the operator or clinician.

[0056] The CT system 100 may also include one or more networked workstations 142. By way of example, a networked workstation 142 may include a display 144; one or more input devices 146, such as a keyboard and mouse; and a processor 148. The networked workstation 142 may be located within the same facility7as the operator workstation 116, or in a different facility, such as a different healthcare institution or clinic.

[0057] The networked workstation 142, whether within the same facility or in a different facility as the operator workstation 116, may gain remote access to the data store server 124 and / or the image reconstruction system 126 via the communication system 128. Accordingly, multiple networked workstations 142 may have access to the data store server 124 and / or image reconstruction system 126. In this manner, x-ray data, reconstructed images, or other data may be exchanged between the data store server 124, the image reconstruction system 126, and the networked workstations 142, such that the data or images may be remotely processed by a networked workstation 142. This data may be exchanged in any suitable format, such as in accordance with the transmission control protocol (“TCP”), the internet protocol (“IP”), or other known or suitable protocols.

[0058] The present disclosure has described one or more preferred embodiments, and it should be appreciated that many equivalents, alternatives, variations, and modifications, aside from those expressly stated, are possible and within the scope of the invention.

Claims

CLAIMS1. An x-ray detector, comprising: a plurality of x-ray detector elements to detect x-rays impinging on each of the plurality of x-ray detector elements; and a plurality' of interpixel reflecting septa, each of the plurality of interpixel reflecting septa being arranged between adjacent ones of the plurality of x-ray detector elements, wherein the plurality of interpixel reflecting septa comprise an optically reflective material that emits a fluorescent x-ray when irradiated by an incident photon.

2. The x-ray detector of claim 1, wherein the plurality of x-ray detector elements are arranged in an array.

3. The x-ray detector of claim 1, wherein the plurality of interpixel reflecting septa are composed of a metal alloy that generates the fluorescent x-ray when irradiated by the incident x-ray.

4. The x-ray detector of claim 1, wherein the plurality of interpixel reflecting septa are composed of a polymer having suspended therein a fluorescent media to generate the fluorescent x-ray when irradiated by the incident x-ray.

5. The x-ray detector of claim 1, wherein each of the plurality of x-ray detector elements comprises a scintillator matched with a photodetector.

6. The x-ray detector of claim 5, wherein each of the plurality of interpixel reflecting septa is arranged between scintillators of the adjacent ones of the plurality of x-ray detectors.

7. The x-ray detector of claim 5, wherein the photodetector in each of the plurality of x-ray detectors comprises a photodiode.

8. The x-ray detector of claim 1, wherein each of the plurality of x-ray detector elements comprises an energy -discriminating detector.

9. The x-ray detector of claim 1, wherein the x-rays impinging on the each of the plurality of x-ray detector elements have energies less than 200 keV.

10. The x-ray detector of claim 1, wherein the reflective material contains a high- Z element with an atomic number of at least 58.

11. The x-ray detector of claim 10. wherein the high-Z element comprises an element with an atomic number in a range of 58 to 69, inclusive.

12. The x-ray detector of claim 10, wherein the reflective material comprises a reflective coating that has been doped with the high-Z element.

13. The x-ray detector of claim 10, wherein the reflective material comprises a metal alloy containing the high-Z element.

14. The x-ray detector of claim 10. wherein the reflective material comprises a polymer containing the high-Z element and an optically reflective material.

15. The x-ray detector of claim 10, wherein the high-Z element comprises at least one of gadolinium (Gd), cerium (Ce), thulium (Tm), or tungsten (W).

16. The x-ray detector of claim 1, further comprising a backing layer coupled to a back surface of the plurality of x-ray detector elements and composed of a high-Z material that captures x-ray photons punching through the plurality of x-ray detector elements and returns some of the x-ray photons back into the plurality of x-ray detector elements.

17. An x-ray detector, comprising: a semiconductor layer to convert incident x-rays into electrical charge signals; a plurality of anodes to measure the electrical charge signals; anda backing layer to capture x-rays passing through the semiconductor layer and to return the captured x-rays into the semiconductor layer.

18. The x-ray detector of claim 17, wherein the plurality of anodes are arranged in a layer that is arranged between the semiconductor layer and the backing layer.

19. The x-ray detector of claim 17, wherein the plurality of anodes are arranged to a side of the semiconductor layer and the backing layer.

20. The x-ray detector of claim 17. wherein the reflective backing layer comprises a first layer and a second layer, wherein the first layer is arranged between the plurality’ of anodes and the second layer.

21. The x-ray detector of claim 20, wherein the second layer has an elemental composition selected to match characteristic x-rays generated by the first layer when x-rays are incident on the first layer.

22. An x-ray detector, comprising: a plurality of x-ray detector layers; a plurality of x-ray fluorescent layers, wherein each of the plurality of x-ray- fluorescent layers is arranged between at least two of the plurality of x-ray detector layers; and wherein x-rays incident on one of the plurality of x-ray fluorescent layers generate fluorescent x-rays that are subsequently detected by an adjacent one of the plurality of x-ray detector layers.

23. The x-ray detector of claim 22, wherein each of the plurality of x-ray detector layers comprise energy-discriminating x-ray detector elements.

Images