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Platform Advantages
  • Easy to use
  • Excellent accuracy
  • Real-time information
  • Minimal staff interaction
  • Intracavitary dose monitoring
PSD Sensor Advantages
  • No correction factors
  • Consistent repeatability
  • Non-angular dependent
  • Water equivalent sensors
  • Monitors IGRT/IMRT & SBRT

Excellent Accuracy
OARtrac’s® accuracy is comparable to TLDs, MOSFETs, and Diodes, without requiring manual correction factors or recalibration before the subsequent treatment. Monitor with confidence in virtual real-time.

Consistent Repeatability
The OARtrac® System offers excellent precision and repeatable data without degradation of dose response or additional correction factors.

Water Equivalent
Plastic Scintillating Detector (PSD) Sensors are designed and manufactured using water equivalent materials. The PSD sensor will not perturb the energy desposition process and exhibits excellent energy independence.

Real-Time Dose Monitoring
The scintillation light is emitted within nanoseconds down the fiber optic cable to the clinical detector unit for real-time measurement.

Low Energy Dependence
Monitor dose from a low to high energy range (0.2Mev to 20Mev) including SBRT, Cyberknife, IGRT, and 3D-CRT platforms with no correction factors.

How It Works
during patient treatments
Clinical Studies BannerClinical Studies Banner

Real-Time in Vivo Dosimetry for SBRT Prostate Treatment Using Plastic Scintillating Dosimetry Embedded in a Rectal Balloon: A Case Study
Purpose: A novel FDA approved in vivo dosimetry device system using plastic scintillating detectors placed in an endorectal balloon to provide real-time in vivo dosimetry for prostatic rectal interface was tested for use with stereotactic body radiotherapy (SBRT). The system was used for the first time ever to measure dose during linear accelerator based SBRT. A single patient was treated with a total dose of 36.25Gy given in 5 fractions. Delivered dose was measured for each treatment with the detectors placed against the anterior rectal wall near the prostate rectal interface. Measured doses showed varying degrees of agreement with computed/planned doses, with average combined dose found to be within 6% of the expected dose. The variance between measurements is most likely due to uncertainty of the detector location, as well as variation in the placement of a new balloon prior to each fraction. Distance to agreement for the detectors was generally found to be within a few millimeters, which also suggested that the differences in measured and calculated doses were due to positional uncertainty of the detectors during the SBRT, which had sharp dose falloff near the penumbra along the rectal wall. Overall, the use of a real time in vivo dosimeter provided a level of safety and improved confidence in treatment delivery. We are evaluating the device further in an IRB-approved prospective partial prostate SBRT trial, and believe further clinical investigations are warranted.

Materials/Methods:The patient was a 55-year-old male with localized prostate cancer clinical stage T1cN0M0, Gleason 7(3+4) initial PSA 6.6 ng/mL Stage IIa with ECOG performance status of 0. He presented for consultation inquiring about advanced radiotherapy techniques including proton therapy and stereotactic radiosurgery. His main concerns were late effects of the radiotherapy for quality of life. After consultation with an urologist and radiation oncologist to discuss management, he elected linear-accelerator-based stereotactic body radiation therapy (SBRT) to deliver 36.25 Gy in 5 fractions of 7.25 Gy per fraction. Fiducial markers were placed in the operating room transperineally by his urologist, and at the time of fiducial marker placement a hydrogel device (SpaceOAR, Augmenix Inc. Waltham, ME) was placed to separate the prostate and the rectal wall to reduce the risk for rectal toxicity related to the radiation exposure.(7) The OARtrac system (RadiaDyne, Houston, TX) is a new in vivo scintillation dosimetry system designed to measure rectal wall dose during prostate radiotherapy procedures. The OARtrac system uses a single-use prostate immobilization endorectal balloon (ERB) embedded with two independent plastic scintillation radiation detectors that provide near real-time dose verification for external beam irradiation of prostatic cancer.(7) Two plastic scintillating detectors (PSDs) are installed on the anterior surface and along the length of an endorectal balloon (labeled as proximal and distal, respectively).(8,9,10) These PSDs measure the dose at the prostatic rectal interface where the dose gradient is steep as the patient is being irradiated with megavoltage X-rays. The rectal balloon reduces the motion of the prostate gland to a minimum while at the same time maintains a constant shape of the rectum; see Fig. 1 for more details about the system. Use of the system is simplified by system-specific software. The system simply needs a few minutes to warm up, input the sensor used, and take a background measurement before each measurement. After each measurement, the user has the option to create a PDF report, and the measurement is saved in the system automatically. The sensors were precalibrated at a Dosimetry Laboratory, but did require an on-site correction for SBRT treatments. During installation of the system, a dose verification test was performed to assess the accuracy of the pair of PSD detectors using a solid-water phantom and the patient-specific plan. These measurements were compared against the expected machine output under the same irradiation condition. To reduce the variability of detector response to radiation during the SBRT treatment, a total of five different PSDs pairs were placed sequentially in the solid-water phantom and measured under the same conditions. Using the measured values, a single system adjustment was made to ensure the dose measurement accuracy. The PSD sensors used during patient treatment were then tested using the solid-water phantom and the patient-specific plan, which allowed for a controlled test with minimal positional uncertainty. Differences between the measured dose and the planned dose were found to be within 2% and 1%, on average, for the proximal and distal sensors respectively when using the solid-water phantom. This was done to test the accuracy of the various sensors used for dose measurements during treatment. For patient treatments, the system was used to measure real-time dose delivered to the patient prostatic rectal interface for each fraction. The measured dose was compared to the computed dose to the rectal wall for SBRT from the treatment planning system (Pinnacle, Philips Healthcare, Madison, WI). The computed dose was found by creating a region of interest (ROI) and determining the mean dose to the ROI. PSD location was determined using the location of the fiducial in the sensor and the known distances between the fiducial and the sensors. In addition, fraction-specific computed dose was found to compare with the measured dose. This was done using the cone-beam CT (CBCT) taken between the two treatment arcs. This CBCT was chosen because it was thought to be the most representative of the total treatment fraction. The CBCT was exported to MIM (MIM Software, Cleveland, OH), a software that allowed the CBCT to be fused to the original treatment planning CT and the treatment planning dose transferred to the CBCT. Rigid registration was performed based upon fiducials in the prostate, as this was the method used during patient treatment since the distance between the anterior rectal wall and the prostate was minimal. Each of the system’s ERB sensors included a fiducial between the two PSDs which was visible in each CBCT, and allowed a better approximation of the placement of the PSD sensors, which might suffer from interfractional positional variation. This information was used to determine a more appropriate predicted dose for each individualfraction. For each treatment fraction, a new endorectal balloon (ERB) and sensor was used. Residual air was removed from the balloon to prevent gas pockets and it was then filled with water before insertion into the rectal cavity. The ERB was placed with lubricating gel with the PSD devices positioned to press on the anterior rectal wall to improve heterogeneity. The balloon was filled to a total of 40 cc of water to help immobilize the gland without exerting excess pressure against the prostate since the excess pressure could move the rectal wall closer to the PTV. Once the balloon was inserted and filled, it was retracted to hold against the anal sphincter. The external rectal stopper was locked onto the shaft of the balloon at the same distance from the tip each day for reproducibility to match the daily treatment distance from the anal verge. Radiation dose measured for each treatment fraction consisted of two treatment arcs and a CBCT taken between the two arcs. Any dose measured from the CBCT was subtracted from the final reading so that only the treatment dose was considered. The only change to the workflow of a typical prostate treatment using a rectal balloon was the presence of a physicist trained to use the OARtrac system. Setup of the system (machine warm-up, connection of sensors, and background measurements) was able to be performed during patient positioning so that treatment was not prolonged or delayed. At the time of the final consultation, informed consent was obtained to proceed with SBRT for the prostate gland using fiducial markers and hydrogel placement prior to treatment planning, as well as for daily use of the PSD and ERB device during treatment to record the in vivo rectal dose. While the hydrogel was placed to provide distance between the rectal wall and the prostate, the use of the PSD and ERB system was not only to measure rectal wall dose in vivo, but also to prevent prostate motion during the SBRT delivery to maximize the benefit from image guidance during the procedure in a complimentary fashion for each device to improve patient safety and treatment delivery. Utilizing both technologies allowed us to limit average rectal wall dose to 36.9% of prescribed daily dose to the prostate gland and verify this in vivo. All procedures were performed in accordance with the ethical standards set forth by the IRB committee and with the Helsinki Declaration of 1975, as revised in 2000.

Conclusion: While this manuscript is exploratory in reporting the first ever use of this novel device in a patient for treatment delivery using SBRT to treat prostate cancer, the clinical implications are very pertinent to improving patient care. By providing an in vivo reading of actual delivered daily dose, it may help to reduce treatment errors in daily setup or initial dose calculations. Patient safety and treatment efficacy are improved through the use of the technology, especially in the setting of hypofractionated treatments, where a daily error can result in a larger deviation in total delivered dose. We are now utilizing this technology in a novel partial prostate SBRT protocol and should anticipate the ability to provide further updates in patient reported outcomes. This is the first reported case using both a hydrogel and the in vivo dosimetry system with PSDs and ERB to maximally reduce dose to the rectal wall and minimize prostate motion during SBRT to reduce late rectal toxicity. This will now be further clinically evaluated on an IRB-approved prospective study in 12 patients using OARtrac system without a hydrogel device. The goal of this study will be to treat a limited volume of the prostate gland as defined through a combination of both anatomic and functional MRI sequencing and correlated with tracked histopathological evaluation in and around the index lesion to define a planning target volume for a 3-fraction regimen of SBRT of 9.75 Gy per fraction to a total dose of 29.25 Gy. This trial will be using a quality of life endpoint to evaluate treatment tolerance and side effects in addition to biochemical response with PSA and serial MRI imaging. The use of OARtrac system in this study was initiated as a result of a recommendation by the protocol review and monitoring committee to track daily delivered dose to assure patient safety and will be correlated to patient reported toxicities. The results from this trial will be the subject of a future manuscript.

A Review of Recent Advances in Optical Fibre Sensors for In-Vivo Dosimetry During Radiotherapy
Abstract: This article presents an overview of the recent developments and requirements in radiotherapy dosimetry, with particular emphasis on the development of optical fibre dosemeters for radiotherapy applications, focusing particularly on in vivo applications. Optical fibres offer considerable advantages over conventional techniques for radiotherapy dosimetry, owing to their small size, immunity to electromagnetic interferences, and suitability for remote monitoring and multiplexing. The small dimensions of optical fibre-based dosemeters, together with being lightweight and flexible, mean that they are minimally invasive and thus particularly suited to in vivo dosimetry. This means that the sensor can be placed directly inside a patient, for example, for brachytherapy treatments, the optical fibres could be placed in the tumour itself or into nearby critical tissues requiring monitoring, via the same applicators or needles used for the treatment delivery thereby providing real-time dosimetric information. The article outlines the principal sensor design systems along with some of the main strengths and weaknesses associated with the development of these techniques. The successful demonstration of these sensors in a range of different clinical environments is also presented.

Conclusion: Optical fibres have been demonstrated to be able to perform accurate radiotherapy dosimetric measurements comparable to existing dosemeters, such as TLDs and diodes, and have the potential for advantages over these conventional systems. They allow for remote monitoring and, in most cases, real-time measurements of the radiation dose. Their small size, lightweight and flexibility have allowed their dosimetric performance to be demonstrated successfully in a range of measurement scenarios. These include both external beam and HDR brachytherapy radiotherapy delivery techniques, in addition to verifying complex treatment deliveries such as IMRT, VMAT and SRS. Their immunity to the intense magnetic field and radio frequency pulses present in the MRI environment may give optical fibre-based dosemeters a significant advantage over conventional dosemeters for real-time monitoring in MRI-guided linacs. The performance of a number of different fibres have been assessed against a wide range of typical radiotherapy dosimetry equipment, including TLDs, diodes, film, 2D arrays and ionization chambers. There is increasing interest in the development of such fibres to allow in vivo dosimetric verification of treatments such as HDR and prostate seed brachytherapy. The majority of optical fibre dosimetry systems are IVD is recommended by many national and international organizations as a safety tool to avoid major errors. However, except for its use during special treatments, IVD during EBRT is not widely routinely applied for the verification of dose delivery to individual patients. The main reasons for this situation are the confidence in other methods of QA, the relatively low frequency of occurrence of serious errors in dose delivery, and the workload involved. In addition to regulatory compliance and reimbursement issues, the rationale for in vivo dose measurements is to provide an accurate and effective independent verification of the overall treatment procedure. It will enable the identification of potential errors in dose calculation, data transfer, dose delivery, patient setup, and changes in patient anatomy. The recent developments in reducing the workload involved in performing IVD using point detectors, for instance by implementing modern methods of data collection and analysis, will stimulate its large-scale implementation. In the coming decade also a substantial increase in 3D in vivo dosimetry may be expected during verification of IMRT and VMAT, as well as during treatment with proton and ion beams. In an era where we are moving toward adaptive RT, 3D in vivo dosimetry could also be used as a tool for plan adaptation. Implementation of IVD in EBRT has the potential to impact clinical outcome and should therefore in principle be used in each radiotherapy center in addition to other QA tools. However, more work is needed to make existing technology more robust and simple for routine and large-scale clinical implementation.

Toward a Real-Time In Vivo Dosimetry System Using Plastic Scintillation Detectors
Purpose: In the present study, we have presented and validated a plastic scintillation detector system designed for real-time multi-probe in vivo measurements.

Materials/Methods: The PSDs were built with a dose-sensitive volume of 0.4 mm(3). The PSDs were assembled into modular detector patches, each containing five closely packed PSDs. Continuous dose readings were performed every 150 ms, with a gap between consecutive readings of <0.3 ms. We first studied the effect of electron multiplication. We then assessed system performance in acrylic and anthropomorphic pelvic phantoms.

Results: The PSDs were compatible with clinical rectal balloons and were easily inserted into the anthropomorphic phantom. With an electron multiplication average gain factor of 40, a twofold increase in the signal/noise ratio was observed, making near real-time dosimetry feasible. Under calibration conditions, the PSDs agreed with the ion chamber measurements to 0.08%. Precision, evaluated as a function of the total dose delivered, ranged from 2.3% at 2 cGy to 0.4% at 200 cGy.

Conclusion: Real-time PSD measurements are highly accurate and precise. These PSDs can be mounted onto rectal balloons, transforming these clinical devices into in vivo dose detectors without modifying current clinical practice. Real-time monitoring of the dose delivered near the rectum during prostate radiotherapy should help radiation oncologists protect this sensitive normal structure.

In Vivo Dosimetry In External Beam Radiotherapy
Abstract: In vivo dosimetry (IVD) is in use in external beam radiotherapy (EBRT) to detect major errors, to assess clinically relevant differences between planned and delivered dose, to record dose received by individual patients, and to fulfill legal requirements. After discussing briefly the main characteristics of the most commonly applied IVD systems, the clinical experience of IVD during EBRT will be summarized. Advancement of the traditional aspects of in vivo dosimetry as well as the development of currently available and newly emerging noninterventional technologies are required for large-scale implementation of IVD in EBRT. These new technologies include the development of electronic portal imaging devices for 2D and 3D patient dosimetry during advanced treatment techniques, such as IMRT and VMAT, and the use of IVD in proton and ion radiotherapy by measuring the decay of radiation-induced radionuclides. In the final analysis, we will show in this Vision 20/20 paper that in addition to regulatory compliance and reimbursement issues, the rationale for in vivo measurements is to provide an accurate and independent verification of the overall treatment procedure. It will enable the identification of potential errors in dose calculation, data transfer, dose delivery, patient setup, and changes in patient anatomy. It is the authors’ opinion that all treatments with curative intent should be verified through in vivo dose measurements in combination with pretreatment checks.

Conclusion: The last decade has seen a number of developments that have made IVD a relatively simple and accurate method to detect errors in dose delivery during EBRT. Both with respect to IVD applying point detectors and the use of EPID-based 2D and 3D dose verification, great progress has been made in understanding the dosimetric characteristics of these detector systems and the development of procedures for their clinical use. Furthermore, IVD is recommended by many national and international organizations as a safety tool to avoid major errors. However, except for its use during special treatments, IVD during EBRT is not widely routinely applied for the verification of dose delivery to individual patients. The main reasons for this situation are the confidence in other methods of QA, the relatively low frequency of occurrence of serious errors in dose delivery, and the workload involved. In addition to regulatory compliance and reimbursement issues, the rationale for in vivo dose measurements is to provide an accurate and effective independent verification of the overall treatment procedure. It will enable the identification of potential errors in dose calculation, data transfer, dose delivery, patient setup, and changes in patient anatomy. The recent developments in reducing the workload involved in performing IVD using point detectors, for instance by implementing modern methods of data collection and analysis, will stimulate its large-scale implementation. In the coming decade also a substantial increase in 3D in vivo dosimetry may be expected during verification of IMRT and VMAT, as well as during treatment with proton and ion beams. In an era where we are moving toward adaptive RT, 3D in vivo dosimetry could also be used as a tool for plan adaptation. Implementation of IVD in EBRT has the potential to impact clinical outcome and should therefore in principle be used in each radiotherapy center in addition to other QA tools. However, more work is needed to make existing technology more robust and simple for routine and large-scale clinical implementation.

A Systematic Characterization of the Low-Energy Photon Response of Plastic Scintillation Detectors
Abstract: To characterize the low energy behavior of scintillating materials used in plastic scintillation detectors (PSDs), 3 PSDs were developed using polystyrene-based scintillating materials emitting in different wavelengths. These detectors were exposed to National Institute of Standards and Technology (NIST)-matched low-energy beams ranging from 20 kVp to 250 kVp, and to 137Cs and 60Co beams. The dose in polystyrene was compared to the dose in air measured by NIST-calibrated ionization chambers at the same location. Analysis of every beam quality spectrum was used to extract the beam parameters and the effective mass energy-absorption coefficient. Monte Carlo simulations were also performed to calculate the energy absorbed in the scintillators’ volume. The scintillators’ expected response was then compared to the experimental measurements and an energy-dependent correction factor was identified to account for low-energy quenching in the scintillators. The empirical Birks model was then compared to these values to verify its validity for low-energy electrons. The clear optical fiber response was below 0.2% of the scintillator’s light for x-ray beams, indicating that a negligible amount of fluorescence contamination was produced. However, for higher-energy beams (137Cs and 60Co), the scintillators’ response was corrected for the Cerenkov stem effect. The scintillators’ response increased by a factor of approximately 4 from a 20 kVp to a 60Co beam. The decrease in sensitivity from ionization quenching reached a local minimum of about 11%±1% between 40 keV and 60 keV x-ray beam mean energy, but dropped by 20% for very low-energy (13 keV) beams. The Birks model may be used to fit the experimental data, but it must take into account the energy dependence of the kB quenching parameter. A detailed comprehension of intrinsic scintillator response is essential for proper calibration of PSD dosimeters for radiology.

Materials and Methods: The PSDs were developed using an optical fiber model to optimize light collection and optical coupling. The scintillators considered in this study were 3 polystyrene-based plastic scintillating fibers emitting at different wavelengths: the BCF-10 and the BCF-12 in blue and the BCF-60 in green (Saint-Gobain Crystals, Nemours, France). Their emission spectra are shown in figure 1. The scintillators were 0.97 mm in diameter and 10 mm long and were surrounded by a 0.015 mm-thick poly(methyl methacrylate) (PMMA) cladding. Each one was optically coupled to a 15 m-long PMMA optical fiber (Eska GH-4001, Mitsubishi International Corporation, NY) of the same diameter. Both the scintillating and the PMMA fibers were covered by a 0.6 mm-thick light-tight polyethylene jacket, and the scintillators’ ends were sealed with a mix of black acrylic paint and epoxy glue. Every interface was polished with successive grain sizes of 30 μm, 9 μm, 3 μm, and 0.3 μm using an automated optical fiber polisher (SpecPro, Krell Technologies, Morganville, NJ). The clear fiber end was inserted into an SMA connector and connected through a fiber adapter to an H10721-20 photomultiplier tube (PMT) (Hamamatsu Photonics, Hamamatsu, Japan). This compact photodetector was chosen for its high photosensitivity (0.075 A W−1) at the BCF-60 peak emission wavelength (530 nm) and its adjustable gain of 5×103 to 5×106. This high sensitivity is desirable for the lowdose- rate measurements involved in low-energy beams (Boivin et al 2015b). The PMT was operated in current mode, and its signal was read in real time by a 2-channel SuperMax electrometer (Standard Imaging, Middleton, WI), which was controlled by a computer. A fourth, scintillator-free PMMA fiber was also prepared to account for fluorescence and Cerenkov contamination. 2.2. Radiation sources The fibers were exposed to 13 types of National Institute of Standards and Technology (NIST)-matched beam qualities including: 11 x-ray beams from the University of Wisconsin Accredited Dosimetry Calibration Laboratory’s (UWADCL) Advanced x-ray (AXI) constant potential x-ray system with a Gulmay CP 320 generator and a Comet 320 tungsten anode, a 137Cs source from the UWADCL G10 irradiator (Hopewell Designs Inc, Alpharetta, GA) holding a 416 Ci 137Cs source (DeWerd et al 2014), and a 60Co source from a Theratron 780C unit (Theratronics, Ottawa, Ontario, Canada). The x-ray beams were carefully matched for first and second half-value layers (HVLs) with comparable x-ray beams at NIST (Lamperti et al 1988). The beam characteristics and additional filtration are listed in table 1. The x-ray measured spectra were also used to extract the HVL, the homogeneity coefficient (HC), and the beam mean energy (Emean), which is the product of every spectrum bin multiplied by its energy and summed over the whole spectrum. The x-ray beams were calibrated in terms of air-kerma rates using NIST-calibrated ionization chambers: the LE-0.8 (Precise Radiation Measurement Inc, Nashville, TN) for the UW20-M to UW40-M beams and an Exradin A3 (Standard Imaging) for the remaining x-ray beam qualities. The chambers were positioned at a distance of 100 cm from the x-ray tube focal spot. The consistency of the air-kerma rate throughout all x-ray measurements was monitored with a transmission monitor chamber that was connected to the electrometer’s second channel. A NIST-calibrated ionization chamber was also used to determine the daily air-kerma rate for both the 137Cs and the 60Co sources. The 137Cs and 60Co mean energy listed in table 1 are the mean of their nominal emission lines, but their actual mean energy would be slightly lower because of scatter from the head and shielding.

Results: The PSDs were developed using an optical fiber model to optimize light collection and optical coupling. The scintillators considered in this study were 3 polystyrene-based plastic scintillating fibers emitting at different wavelengths: the BCF-10 and the BCF-12 in blue and the BCF-60 in green (Saint-Gobain Crystals, Nemours, France). Their emission spectra are shown in figure 1. The scintillators were 0.97 mm in diameter and 10 mm long and were surrounded by a 0.015 mm-thick poly(methyl methacrylate) (PMMA) cladding. Each one was optically coupled to a 15 m-long PMMA optical fiber (Eska GH-4001, Mitsubishi International Corporation, NY) of the same diameter. Both the scintillating and the PMMA fibers were covered by a 0.6 mm-thick light-tight polyethylene jacket, and the scintillators’ ends were sealed with a mix of black acrylic paint and epoxy glue. Every interface was polished with successive grain sizes of 30 μm, 9 μm, 3 μm, and 0.3 μm using an automated optical fiber polisher (SpecPro, Krell Technologies, Morganville, NJ). The clear fiber end was inserted into an SMA connector and connected through a fiber adapter to an H10721-20 photomultiplier tube (PMT) (Hamamatsu Photonics, Hamamatsu, Japan). This compact photodetector was chosen for its high photosensitivity (0.075 A W−1) at the BCF-60 peak emission wavelength (530 nm) and its adjustable gain of 5×103 to 5×106. This high sensitivity is desirable for the lowdose- rate measurements involved in low-energy beams (Boivin et al 2015b). The PMT was operated in current mode, and its signal was read in real time by a 2-channel SuperMax electrometer (Standard Imaging, Middleton, WI), which was controlled by a computer. A fourth, scintillator-free PMMA fiber was also prepared to account for fluorescence and Cerenkov contamination. 2.2. Radiation sources The fibers were exposed to 13 types of National Institute of Standards and Technology (NIST)-matched beam qualities including: 11 x-ray beams from the University of Wisconsin Accredited Dosimetry Calibration Laboratory’s (UWADCL) Advanced x-ray (AXI) constant potential x-ray system with a Gulmay CP 320 generator and a Comet 320 tungsten anode, a 137Cs source from the UWADCL G10 irradiator (Hopewell Designs Inc, Alpharetta, GA) holding a 416 Ci 137Cs source (DeWerd et al 2014), and a 60Co source from a Theratron 780C unit (Theratronics, Ottawa, Ontario, Canada). The x-ray beams were carefully matched for first and second half-value layers (HVLs) with comparable x-ray beams at NIST (Lamperti et al 1988). The beam characteristics and additional filtration are listed in table 1. The x-ray measured spectra were also used to extract the HVL, the homogeneity coefficient (HC), and the beam mean energy (Emean), which is the product of every spectrum bin multiplied by its energy and summed over the whole spectrum. The x-ray beams were calibrated in terms of air-kerma rates using NIST-calibrated ionization chambers: the LE-0.8 (Precise Radiation Measurement Inc, Nashville, TN) for the UW20-M to UW40-M beams and an Exradin A3 (Standard Imaging) for the remaining x-ray beam qualities. The chambers were positioned at a distance of 100 cm from the x-ray tube focal spot. The consistency of the air-kerma rate throughout all x-ray measurements was monitored with a transmission monitor chamber that was connected to the electrometer’s second channel. A NIST-calibrated ionization chamber was also used to determine the daily air-kerma rate for both the 137Cs and the 60Co sources. The 137Cs and 60Co mean energy listed in table 1 are the mean of their nominal emission lines, but their actual mean energy would be slightly lower because of scatter from the head and shielding.

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