MicroCT with energy-resolved photon-counting detectors

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MicroCT with energy-resolved photon-counting detectors
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  MicroCT with energy-resolved photon-counting detectors X Wang 1 , D Meier  2 , S Mikkelsen 2 , G E Maehlum 2 , D J Wagenaar  3 , BMW Tsui 1 , B E Patt 3 ,and E C Frey 1E C Frey: efrey1@jhmi.edu 1 Department of Radiology and Radiological Science, Johns Hopkins University, Baltimore, MD,USA 2 Gamma Medica-Ideas (AS), Oslo, Norway 3 Gamma Medica-Ideas, Northridge, CA, USA  Abstract The goal of this paper was to investigate the benefits that could be realistically achieved on amicroCT imaging system with an energy-resolved photon-counting x-ray detector. To this end, we built and evaluated a prototype microCT system based on such a detector. The detector is based oncadmium telluride (CdTe) radiation sensors and application-specific integrated circuit (ASIC)readouts. Each detector pixel can simultaneously count x-ray photons above six energy thresholds, providing the capability for energy-selective x-ray imaging. We tested the spectroscopic performance of the system using polychromatic x-ray radiation and various filtering materials withKabsorption edges. Tomographic images were then acquired of a cylindrical PMMA phantomcontaining holes filled with various materials. Results were also compared with those acquired using an intensity-integrating x-ray detector and single-energy (i.e. non-energy-selective) CT. This paper describes the functionality and performance of the system, and presents preliminaryspectroscopic and tomographic results. The spectroscopic experiments showed that the energy-resolved photon-counting detector was capable of measuring energy spectra from polychromaticsources like a standard x-ray tube, and resolving absorption edges present in the energy range used for imaging. However, the spectral quality was degraded by spectral distortions resulting fromdegrading factors, including finite energy resolution and charge sharing. We developed a simplecharge-sharing model to reproduce these distortions. The tomographic experiments showed thatthe availability of multiple energy thresholds in the photon-counting detector allowed us tosimultaneously measure target-to-background contrasts in different energy ranges. Compared withsingle-energy CT with an integrating detector, this feature was especially useful to improvedifferentiation of materials with different attenuation coefficient energy dependences. 1. Introduction Conventional x-ray detectors integrate the total electrical current produced in the radiationsensor and disregard the charge amplitude from individual photon detection events. Thecharge amplitude from each event is proportional to the photon’s detected energy. Duringthis integration, both the detector leakage current and charges resulting from x-ray detectionare summed and measured, and provide no information about the energy of individual photons or the dependence of the attenuation coefficients in the object.A novel alternative to these detectors is photon-counting x-ray detectors that countindividual x-ray photons interacting in each pixel of the detector. Energy-resolved photoncounting with multiple energy thresholds provides the additional capability of counting photons based on their detected energies (Bert et al  2003, Beuville et al  1998, Schlomka et al  2008, Shikhaliev et al  2005). These capabilities have been made possible by the  NIH Public Access Author Manuscript Phys Med Biol . Author manuscript; available in PMC 2011 May 17. Published in final edited form as: Phys Med Biol  . 2011 May 7; 56(9): 2791–2816. doi:10.1088/0031-9155/56/9/011. NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t    availability of fast semiconductor radiation sensor materials (e.g., Si, CdTe, GaAs), whichsupport relatively short charge collection periods (e.g., 50 ns or less) and provide good energy resolution, combined with application-specific integrated circuits (ASICs) suitablefor multi-pixel parallel readout and fast counting.Photon-counting detectors have a number of potential advantages. First, photon counting provides more optimal weighting than intensity integrating, resulting in lower noise and  better contrast (Frey et al  2007a). This capability can benefit either conventional (i.e. non-energy-selective) CT or energy-selective (e.g., dual-energy) CT.Second, the ability to resolve energies allows energy-selective imaging with a single x-rayexposure. Such detectors acquire simultaneous measurements of the x-ray photon flux aboveone or more user-defined energy thresholds. These data can be used to obtain the x-ray photon flux in a set of non-overlapping energy windows. Therefore measurements from suchdetectors can be used to perform energy-selective x-ray CT imaging and provideinformation about the energy dependence of the attenuation coefficients of the materials inthe object and, through these, about the chemical elements present (Feuerlein et al  2008,Alvarez and Macovski 1976). For example, different target materials can be more efficientlyand quantitatively identified and differentiated (Schlomka et al  2008, Cormode et al  2010)using an energy-resolved photon-counting CT system.In our previous work (Frey et al  2007b), simulations have shown that the advantage from photon counting (without energy resolving capability) was small for conventional CT and  basis material decomposition. The largest advantage was from the energy resolvingcapability where better weighting improved the contrast-to-noise ratio and the multipleenergy windows resulted in improved basis decomposition. Thus this paper will concentrateon applications that use the energy resolving capability in an effort to demonstrate the extentto which current photon-counting detectors can reach the potential seen in simulations and todetermine challenges and problems that must be addressed to realize these potentials in thefuture.In this study we used a detector that we developed and that is based on cadmium telluride(CdTe) radiation sensors and an ASIC for energy-resolved photon counting (Mikkelsen et al 2008). This paper will describe the detector hardware and system configuration and presentresults from studies characterizing its spectral performance. In previous work (Wang et al 2011) we concentrated on one method to use the energy information, i.e. the use of multipleenergy windows to allow simultaneous separation of materials having K-absorption edges inthe energy range used for imaging. Thus in this work we concentrate on the other potentialuse of the energy information, i.e. the potential to use the energy information to provideimproved contrast between materials when appropriate energy windows are used.Finally, in studying the spectral performance we observed some degradation in the energyspectrum at low energies that could reduce the utility of the energy information. We believethese degradations are due to charge-sharing effects, i.e. where energy from a photonincident in one detector pixel results in a count in both that pixel and a neighboring pixelwith both counts having incorrect recorded energies. We developed a simple model of thiseffect that is able to reproduce the spectral distortions, and we discuss this model in thecontext of the physical effects that could result in charge sharing. Wang et al.Page 2 Phys Med Biol . Author manuscript; available in PMC 2011 May 17. NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t    2. Methods 2.1. Energy-resolved photon-counting x-ray detector  The energy-resolved photon-counting detector used in this work is housed in a sealed aluminum enclosure.Figure 1 shows a picture of the radiation sensor and the electronics inside the detector enclosure with the aluminum cover and the radiation collimator removed. The specificationsof the detector are summarized in table 1. The interior temperature of the detector isregulated by a Peltier-based cooler to maintain a stable interior air temperature of ~25 °C. 2.1.1. Radiation sensor and collimator— The radiation sensor of the photon-countingdetector is made from 3mmthick CdTe layer (Acrorad Co. Ltd, Japan) and forms a line witha sensitive length of 204.8 mm. The line detector has 512 imaging pixels, formed by a set of strip anodes with a pitch of 0.4 mm along the crystal length direction, and a width of 1.6mm.The incident x-ray beam is collimated by a lead slit collimator (not shown in figure 1)centered above the radiation sensor. The slit width is adjustable in the range of 0.5 to 1.5mm. In this study we used a width of 1.2 mm. The aluminum-protective cover (not shown infigure 1) is milled down to a thickness of 2 mm where the opening slit of the collimator islocated. Therefore the incident x-ray beam is always attenuated by a 2 mm aluminum layer.In the future we plan to remove this 2 mm aluminum layer from the detector entrancewindow and replace it with less attenuating materials. 2.1.2. ASICs— The row of 512 pixels is comprised of two detector modules eachcontaining eight 32-channel custom-made ASICs. Figure 2 shows one piece of CdTe crystaland two ASICs. The crystal is bump bonded to a ceramic printed circuit board that hasaluminum traces leading to a fine pitch connector. The connector plugs into a matingconnector that is soldered to an FR4 printed circuit board. The ASICs are wire bonded ontothis FR4 printed circuit board. Digital frame data can be transmitted from the ASICs atregular intervals to an external computer via an FPGA resident in the detector.A schematic of the readout circuitry of a detector pixel is illustrated in figure 3. Each pixel isconnected to a preamplifier followed by a CR-RC shaping amplifier with pole-zerocompensation to improve the signal-to-noise ratio. The shaping time for the shapingamplifier was set to approximately 50 ns to provide high-count-rate performance.The output of the analog stage is connected to six amplitude discriminator circuits. Thediscriminator thresholds are globally programmable via a global digital-to-analog converter (DAC). In addition, each threshold for each pixel can be further adjusted via a trim-DAC.These trim-DACs allow one to individually adjust the thresholds so that they are aligned relative to each other. This allows compensation for differences in pixel gains and offsetvoltages arising from process variations during manufacturing process. The output of eachdiscriminator is connected to a one-shot circuit with a pulse length set to ~250 ns. The netresult is that when the input amplitude exceeds the threshold a single pulse is produced.Another pulse can be produced only after an interval of 250 ns and when the previous pulsefalls below the threshold.The pulses produced by the threshold circuit are counted by a 17-bit counter (the 17th bit islatched to detect overflow). Thus the counter increments only if the pulse height out of theanalog stage exceeds the discriminator threshold, and at most one count is registered every250 ns. There is a separate counter for each pixel and threshold. The digital counter values Wang et al.Page 3 Phys Med Biol . Author manuscript; available in PMC 2011 May 17. NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t    can be read out upon request. All the counters can be reset, started and stopped at the sametime via a common control signal. Prior to readout to the computer, all counter values aretransferred at high speed to a readout buffer so that the ASIC can continue counting duringthe time the counter values are read out by the computer.The six discriminators for each pixel act in parallel, allowing simultaneous acquisition of counts above six independent energy thresholds. Subtracting the counts from pairs of thresholds provides the number of counts in an energy window defined by the twothresholds. As a result, six energy windows are available, with the highest window havingno upper threshold. In this work we typically set the highest threshold at an energycorresponding to the x-ray tube voltage; the counts from the corresponding counter thusrepresented pulse pileup only and were discarded. The default energy thresholds werefactory calibrated using photons emitted by Co-57 and Am-241 radionuclide sources. 2.2. X-ray microCT system prototype We installed the energy-resolved photon-counting detector in a benchtop microCT system(figure 4). This system consisted of a SourceRay SB-120–350 x-ray generator system(SourceRay Inc., Bohemia, NY) capable of tube currents and voltages of 350 µA and 120kV, respectively. The tube had a 75 µm focal spot and a 0.75 mm thick beryllium exitwindow. Since the detector consisted of only one row of active pixels, the x-ray tube wascollimated by a 1.6 mm wide lead slit to provide a fan beam of x-ray irradiation. We aligned the fan beam with the radiation sensor to maximize direct x-ray illumination of the detector.For some experiments, additional external filters were used to shape the output spectra asneeded for the specific application.A commercial motor-controlled rotary stage (Velmex Inc., Bloomfield, NY) was positioned  between the x-ray generator system and the detector in order to rotate the imaged object and acquire projections for tomographic reconstruction. In most acquisitions we positioned thedetector 88.8 cm from the x-ray tube focal spot, while the phantom rotational stage was 20.3cm from the tube focal spot. Hence magnification was ~4.37 and the expected spatialresolution at isocenter was ~0.09 mm. The field of view at isocenter was ~ 4.7 cm. 2.3. Phantom We used a cylindrical phantom with a 2.54 cm diameter circular cross section (figure 5). The phantom was made of polymethyl methacrylate (PMMA, or acrylic) and had five 4.8 mmdiameter holes drilled at equiangular intervals parallel to the cylinder’s axis on a radius of 6.25 mm. The holes could be filled with various solid or liquid materials. Table 2summarizes the materials that were imaged, which included the following three categories of materials: 1. soft-tissue-like materials: water, PMMA, polytetrafluoroethylene (PTFE), polyoxymethylene (POM), glycol-modified polyethylene terephthalate (PET) and  polycarbonate; 2.  bone-like material: bone-equivalent plastic; 3. contrast agents having K-edges in the energy range of interest: aqueous solutions of an iodated contrast agent, Omnipaque™350. 2.4. Spectroscopic experiments This detector could be used to measure the spectra of polychromatic x-ray radiation. Toachieve this, one or several threshold(s) were swept through a pre-set energy range with asmall, pre-set energy step. By numerically differentiating the data we obtained the energy Wang et al.Page 4 Phys Med Biol . Author manuscript; available in PMC 2011 May 17. NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t    spectrum. In this study we used this method to measure the output x-ray spectra from the x-ray tube. The x-ray tube output was filtered with a 0.794 mm thick aluminum sheet. Therewas no other filtering between the tube and the detector. A tube current of 50 µA was used and the resulting count rate was approximately 4 × 10 4  cps/pixel for the 80 kVp spectrum.For this study, we kept the count rate relatively low to reduce the effects of pulse pileup.Furthermore, filtering materials with K-absorption edges in the energy range of interest werealso used (table 3), in addition to the previously mentioned 0.794 mm aluminum sheet, inorder to investigate the detector’s ability to resolve the discontinuous changes in x-rayattenuation and the resulting spectra at K-edge energies. To use liquids (i.e. Magnevist™ or Omnipaque™350) as filters, a syringe filled with the solution was placed in front of the x-ray tube exit window, and the thickness of the filter corresponded to the internal diameter of the syringe.These spectral measurements provided important information to aid in understanding thespectral response of the detector and to model and compensate for the physical degradingfactors present in the detector. The characteristic peaks and endpoints in the polychromaticx-ray spectra (section 3.1) were investigated in terms of their potential use for calibration of the energy threshold DAC values in units of keV. This is desirable as it would be more practical than using measurements from monochromatic radiation sources such assynchrotron radiation or radionuclide sources.For comparison purposes, we also measured the spectra from the x-ray tube with the samesource filters (0.794 mm aluminum sheet, and K-edge materials listed in table 3 whenapplicable) using a commercial spectrometer (XR-100T-CdTe, Amptek Inc., Bedford, MA).This spectrometer is comprised of a 1 mm thick CdTe radiation sensor, digital pulse processor and multi-channel analyzer. The count rate capability of this device is much lower than our photon-counting detector, so we used a 200 µm diameter collimator (provided byAmptek) in front of the spectrometer to reduce the photon flux. A 2 mm thick aluminumsheet was also placed in front of the spectrometer entrance window to mimic the effect of the 2 mm thick aluminum protection cover in front of the CdTe sensor of the energy-resolved photon-counting detector. Due to its superior energy resolution (<1.2 keV FWHM)and readout electronics, the spectrometer is capable of providing much more accuratemeasurements of the x-ray spectrum than photon-counting detectors intended for high countrate x-ray imaging. Therefore, measurements made using the spectrometer are, after correction for the sensor thickness, very good approximations of the true spectra. Theseresults were compared with the spectra acquired using the energy-resolved photon-countingdetector in order to show the spectral degradations. For convenience, in the remainder of this paper, we may use “photon-counting detector” to specifically refer to the energy-resolved  photon-counting detector (although technically the spectrometer is also a photon-countingdetector). 2.5. X-ray computed tomography experiments: ring reduction and contrast analysis2.5.1. Tomographic acquisition— Figure 6 describes the procedure we used to obtaintomographic images from different x-ray energies. The x-ray tube output was filtered with a0.794 mm thick aluminum sheet, as in the spectroscopic experiments. There was no other filtering between the tube and the phantom. During experiments involving contrast agents inthe phantom, we usually set one energy threshold at each K-absorption edge, one at the tubevoltage, and distributed the other thresholds such that the number of counts betweenadjacent energy thresholds was approximately equal. The counter for each energy threshold counted the number of photons with energies greater than the energy corresponding to thethreshold voltage. Thus, subtracting the detected counts acquired using a higher energy Wang et al.Page 5 Phys Med Biol . Author manuscript; available in PMC 2011 May 17. NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t  NI  H-P A A  u t  h  or M an u s  c r i   p t  
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