[go: up one dir, main page]

WO2025050041A1 - Détecteur de rayons x avec composants fluorescents à rayons x - Google Patents

Détecteur de rayons x avec composants fluorescents à rayons x Download PDF

Info

Publication number
WO2025050041A1
WO2025050041A1 PCT/US2024/044866 US2024044866W WO2025050041A1 WO 2025050041 A1 WO2025050041 A1 WO 2025050041A1 US 2024044866 W US2024044866 W US 2024044866W WO 2025050041 A1 WO2025050041 A1 WO 2025050041A1
Authority
WO
WIPO (PCT)
Prior art keywords
ray detector
ray
rays
layer
septa
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Pending
Application number
PCT/US2024/044866
Other languages
English (en)
Inventor
Scott S. Hsieh
Current Assignee (The listed assignees may be inaccurate. Google has not performed a legal analysis and makes no representation or warranty as to the accuracy of the list.)
Mayo Foundation for Medical Education and Research
Mayo Clinic in Florida
Original Assignee
Mayo Foundation for Medical Education and Research
Mayo Clinic in Florida
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Application filed by Mayo Foundation for Medical Education and Research, Mayo Clinic in Florida filed Critical Mayo Foundation for Medical Education and Research
Publication of WO2025050041A1 publication Critical patent/WO2025050041A1/fr
Pending legal-status Critical Current
Anticipated expiration legal-status Critical

Links

Classifications

    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/16—Measuring radiation intensity
    • G01T1/20—Measuring radiation intensity with scintillation detectors
    • G01T1/2002—Optical details, e.g. reflecting or diffusing layers

Definitions

  • 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.
  • 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.
  • 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 plurality of 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.
  • 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.
  • 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.
  • 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.
  • 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.
  • 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.
  • 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 K a characteristic x-rays are 42 keV.
  • Pixel pitch is 0.5 mm and fill factor is 64%.
  • 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.
  • FIG. 8 shows sensitivity of the detector variance to different Gd loading factors at 120 kVp. Improvements are limited beyond 0.5 g/cm 3 of Gd and saturate near 1 g/cm 3 .
  • 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/cm 3 of 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.
  • 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.
  • 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.
  • a high-Z element may include an element wi th an atomic number of at least 58.
  • 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.”
  • the x-ray detector may be an energy-integrating detector whose resolution may be improved without sacrificing dose efficiency from the fill factor loss.
  • 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 septa material. 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.
  • 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.
  • the x-ray detector may be an energy-discriminating detector, such as a photon-counting detector.
  • 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.
  • the probability distribution function of the signal deposited, s(E) is a delta function; that is, all photons increment the counter once.
  • s(E) is equal to the incident spectrum; that is. all photons add their kinetic energy to the photodiode signal.
  • F fill factor
  • s(E) has a probability' F to be sampled from the incident spectrum, and the remaining probability' (1 — F) to be zero.
  • the total integrated signal follows a compound Poisson distribution.
  • 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:
  • the signal probability distribution function s(E) is changed.
  • 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 a compound containing the high-Z element may be mixed or otherwise loaded into the septa material.
  • the high-Z material e.g., high-Z element, compound containing a high-Z element
  • 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.
  • 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).
  • a high-Z element e.g., gadolinium (Gd). cerium (Ce), thulium (Tm), tungsten (W), or the like.
  • 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.
  • 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.
  • 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.
  • 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-ray source, which are of much greater energy.
  • 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.
  • 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.
  • 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.
  • 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%.
  • 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.
  • 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.
  • the high-Z element may be gadolinium (Gd), cerium (Ce), thulium (Tm), tungsten (W), or the like.
  • the interpixel 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.
  • 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.
  • the interpixel reflecting septa 14 may be composed of a material into which a high-Z element or material is mixed.
  • 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.
  • the x-ray detector elements 12 may be composed of solid-state x-ray detector elements or pixels.
  • 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.
  • a semiconductor material e.g., CdTe
  • 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.
  • 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.
  • 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.
  • 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.
  • 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 distribution where /r E is 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.
  • Recapture can be improved with a multi-material backing layer.
  • 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.
  • the backing layer 24 may implement a 2- material design: a front layer 26 adjacent the plurality of anodes 22 and a backing layer 28 arranged after the front layer 26. thereby sandwiching the front layer 26 between the plurality of anodes 22 and the backing layer 28.
  • 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.
  • 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%.
  • the backing layer 24 may be composed of more than two layers, such as three or more layers.
  • a signal may also originate from holes and this signal may drift towards the cathode.
  • the x-ray detector 10 may include a plurality of cathodes in addition or as an alternative to the plurality of anodes 22.
  • the x-ray detector 10 may be a multilayer x-ray detector.
  • the x-ray detector 10 is composed of a plurality of x-ray detector layers 30.
  • 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.
  • 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.
  • 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.
  • 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 today are 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.
  • 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.
  • tungsten is an adequate element for fluorescence.
  • certain ceramic substrates may also be used.
  • the ceramic substrate may incorporate tungsten carbide into the ceramic.
  • other alloys that have high atomic number and good thermal conductivity may be used.
  • 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.
  • the atomic number, Z, of the element loaded into the septa affects s(F) in two competing ways.
  • the k-edge of the element determines the minimum energy of photons that are eligible for x-ray fluorescence.
  • 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.
  • 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.
  • 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.
  • 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.
  • GEANT4 is a set of general purpose Monte Carlo codes that track the passage of particles through matter.
  • 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.
  • 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.
  • the anti-scatter grid was not modeled. The antiscatter grid presents another limit to the fill factor.
  • each pixel is bordered by interpixel reflecting septa, which may be assumed to be 0. 1 mm in width in most instances.
  • 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.
  • a binder is used to provide structure, and the binder (not the pigment) occupies most of the volume of the septa.
  • the binding agent may be a plastic, such as acrylic (e.g.. PMMA) of density 1.19 g/cm 3 or a transparent epoxy.
  • 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 is also marked and has higher variance because of the non-uniform (energy weighting) of the photons, which has sometimes been described as radiation Swank noise.
  • 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.
  • 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.
  • 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.
  • 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 K a or 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.
  • 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 / cm 3 of 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.
  • 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/cm 3 . Beyond 1 g/cm 3 , the incremental losses from septa absorption of characteristic x-rays within the septa appear to cancel out incremental improvements in x-ray attenuation.
  • 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.
  • Gd is optimal or nearly optimal at all kVp 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.
  • 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.
  • 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.
  • each x-ray detector 110 is capable of counting the number of x-ray photons that impinge upon the detector 110.
  • the g an tty 102 and the components mounted thereon rotate about a center of rotation 114 located within the CT system 100.
  • 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.
  • the operator workstation 116 is in communication with a data store server 124 and an image reconstruction system 126.
  • 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.
  • the communication system 128 may include both proprietary or dedicated networks, as well as open networks, such as the internet.
  • 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 source 104 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.
  • 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.
  • 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.
  • 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.
  • 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.
  • the CT system 100 may also include one or more networked workstations 142.
  • 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 facility 7 as the operator workstation 116, or in a different facility, such as a different healthcare institution or clinic.
  • the networked workstation 142 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.
  • TCP transmission control protocol
  • IP internet protocol

Landscapes

  • Physics & Mathematics (AREA)
  • Health & Medical Sciences (AREA)
  • Life Sciences & Earth Sciences (AREA)
  • General Physics & Mathematics (AREA)
  • High Energy & Nuclear Physics (AREA)
  • Molecular Biology (AREA)
  • Spectroscopy & Molecular Physics (AREA)
  • Measurement Of Radiation (AREA)

Abstract

Selon l'invention, l'efficacité et/ou la résolution de détecteur de rayons X sont améliorées en utilisant des matériaux fluorescents à rayons X dans la conception de détecteur de rayons X. Les composants fluorescents à rayons X capturent des rayons X incidents et émettent des rayons X fluorescents qui sont détectés par des éléments détecteurs de rayons X adjacents. Un composant du détecteur de rayons X peut être rendu fluorescent par rayons X par incorporation d'un matériau contenant un élément à Z élevé dans la conception de composant de détecteur de rayons X. Le composant de détecteur fluorescent de rayons X peut être un septa réfléchissant interpixel, une couche de support et/ou un réflecteur inter-détecteur pour des détecteurs multicouches.
PCT/US2024/044866 2023-08-31 2024-08-30 Détecteur de rayons x avec composants fluorescents à rayons x Pending WO2025050041A1 (fr)

Applications Claiming Priority (2)

Application Number Priority Date Filing Date Title
US202363579868P 2023-08-31 2023-08-31
US63/579,868 2023-08-31

Publications (1)

Publication Number Publication Date
WO2025050041A1 true WO2025050041A1 (fr) 2025-03-06

Family

ID=92900102

Family Applications (1)

Application Number Title Priority Date Filing Date
PCT/US2024/044866 Pending WO2025050041A1 (fr) 2023-08-31 2024-08-30 Détecteur de rayons x avec composants fluorescents à rayons x

Country Status (1)

Country Link
WO (1) WO2025050041A1 (fr)

Citations (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4725734A (en) * 1984-05-10 1988-02-16 Kabushiki Kaisha Toshiba Radiation-detecting device for computed tomography
US5773829A (en) * 1996-11-05 1998-06-30 Iwanczyk; Jan S. Radiation imaging detector
US5808306A (en) * 1993-04-05 1998-09-15 Cardiac Mariners, Inc. X-ray detector
JP2005121528A (ja) * 2003-10-17 2005-05-12 Rigaku Corp ２次元イメージ素子及びそれを利用した２次元イメージ検出装置並びにｘ線分析装置
US9285488B2 (en) * 2012-02-14 2016-03-15 American Science And Engineering, Inc. X-ray inspection using wavelength-shifting fiber-coupled scintillation detectors

Patent Citations (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4725734A (en) * 1984-05-10 1988-02-16 Kabushiki Kaisha Toshiba Radiation-detecting device for computed tomography
US5808306A (en) * 1993-04-05 1998-09-15 Cardiac Mariners, Inc. X-ray detector
US5773829A (en) * 1996-11-05 1998-06-30 Iwanczyk; Jan S. Radiation imaging detector
JP2005121528A (ja) * 2003-10-17 2005-05-12 Rigaku Corp ２次元イメージ素子及びそれを利用した２次元イメージ検出装置並びにｘ線分析装置
US9285488B2 (en) * 2012-02-14 2016-03-15 American Science And Engineering, Inc. X-ray inspection using wavelength-shifting fiber-coupled scintillation detectors

Similar Documents

Publication Publication Date Title
Shikhaliev et al. Photon counting computed tomography: concept and initial results
Shikhaliev Energy-resolved computed tomography: first experimental results
Taguchi et al. Vision 20/20: single photon counting x‐ray detectors in medical imaging
Seibert et al. X-ray imaging physics for nuclear medicine technologists. Part 2: X-ray interactions and image formation
EP2052279B1 (fr) Dispositif et procede pour tomographie spectrale informatisee
Åslund et al. Scatter rejection in multislit digital mammography
US7352840B1 (en) Micro CT scanners incorporating internal gain charge-coupled devices
CN101375798B (zh) 具有第二射线管/检测器修补的ct成像系统和方法
Boone et al. A Monte Carlo study of x‐ray fluorescence in x‐ray detectors
JP2010513908A (ja) エネルギー分解検出システム及び撮像システム
Star‐Lack et al. A piecewise‐focused high DQE detector for MV imaging
CN109073768A (zh) 用于x射线检测器的系统和方法
Moses Scintillator requirements for medical imaging
Dunning et al. Optimization of a table-top x-ray fluorescence computed tomography (XFCT) system
JP7418334B2 (ja) Ｘ線イメージングのためのディテクター
Chan et al. Monte Carlo simulation in diagnostic radiology
Howansky et al. An apparatus and method for directly measuring the depth‐dependent gain and spatial resolution of turbid scintillators
Toia et al. Approaches, advantages, and challenges to photon counting detector and multi-energy CT
Whitman et al. Digital mammography: a practical approach
Lundqvist Silicon strip detectors for scanned multi-slit x-ray imaging
Day et al. The detective quantum efficiency of cadmium telluride photon‐counting x‐ray detectors in breast imaging applications
US20230375727A1 (en) Combined imaging detector and imaging system
Michail et al. Radiation detectors and sensors in medical imaging
Larsson et al. Characterization of scintillator‐based detectors for few‐ten‐keV high‐spatial‐resolution x‐ray imaging
WO2025050041A1 (fr) Détecteur de rayons x avec composants fluorescents à rayons x

Legal Events

Date Code Title Description
121 Ep: the epo has been informed by wipo that ep was designated in this application

Ref document number: 24777072

Country of ref document: EP

Kind code of ref document: A1