Proton beam output measurement with an extrapolation chamber

Medical Dosimetry, Vol. 23, No. 4, pp. 288-291, 1998 Copyright © 1998 American Association of Medical Dosimetrists Printed in the USA. All rights reserved 0958-3947/98/$–see front matter

PII S0958-3947(98)00024-7

PROTON BEAM OUTPUT MEASUREMENT WITH AN EXTRAPOLATION CHAMBER COREY ZANKOWSKI1, STANISLAV VATNITSKY2, JEFFREY SIEBERS2, and ERVIN B. PODGORSAK1, 1

Department of Medical Physics, McGill University, Montre´al, Quebe´c; 2Department of Radiation Sciences, Loma Linda University Medical Center, Loma Linda, CA

Abstract—A variable air-volume, parallel-plate extrapolation chamber forming an integral part of a polystyrene phantom was used in measurement of dose rate in a 250 MeV clinical proton beam. The sensitive air-volume of the extrapolation chamber is controlled through the movement of the chamber piston by means of a micrometer mounted on the phantom body. The relative displacement of the piston is monitored by a calibrated mechanical distance travel indicator. The proton beam dose rate determined with the uncalibrated extrapolation chamber was 5% lower than the dose rate determined with a calibrated Farmer-type thimble chamber at the same depth in the polystyrene phantom. Despite the current 5% discrepancy, uncalibrated extrapolation chambers may offer a simple and practical alternative to current techniques used in output measurements of proton beam machines. © 1998 American Association of Medical Dosimetrists.

INTRODUCTION

trapolation chambers in the given phantom material agreed within 6 1% with the doses determined with calibrated thimble chambers following the AAPM TG 214 and TG 255 protocols, respectively. Alternative techniques for proton beam dose rate measurements are important because they may help to improve the reliability of current basic physics dosimetry data. To test the suitability of extrapolation chambers in proton beam dose rate measurements, we took the polystyrene version of the extrapolation chamber to LLUMC and carried out a basic dose output measurement procedure on the 250 MeV monoenergetic proton beam. In comparison with standard ionization chamber techniques in practice at LLUMC, the extrapolation chamber measurements resulted in a 5% lower dose rate. We cannot explain this discrepancy but, based on our work with photon and electron beams and the preliminary results on the LLUMC proton beam, we believe that the extrapolation chamber could provide a simple and reliable method for the calibration of clinical, high-energy proton beams.

Because they offer several advantages over conventional photon and electron beams, clinical proton beams would gain a widespread acceptance in modern radiotherapy were it not for the extreme costs related to installing and maintaining a hospital-based proton beam facility. The high costs notwithstanding, there are currently several hospital-based proton beam facilities in clinical operation around the world, and new ones are being built or planned. In contrast to the situation with photon and electron beams, absolute dosimetry of clinical proton beams is more ambiguous because of inherent uncertainties in basic physics data related to proton beam dosimetry. Recently, Vatnitsky et al.1 published results of an intercomparison of proton beam dosimetry techniques from several institutions around the world involved in clinical use of proton beams. Representatives from various institutions used their own ionization chambers and calibration protocols to determine the proton beam output at the Loma Linda University Medical Center (LLUMC). An intercomparison of results showed discrepancies of up to 6% between various institutions, resulting from variations in basic physics data sets that were used in dose calculations. Recently we built two extrapolation chambers for use in dosimetry of photon and electron beams. Both chambers were described in the literature;2,3 one forms an integral part of a polystyrene phantom2 and the other is built into a solid water phantom.3 The photon and electron doses determined with the two uncalibrated ex-

MATERIALS AND METHODS Proton beam dosimetry Models for determining the absorbed dose in clinical proton beams are based on protocols developed by the Task Group 20 of the American Association of Physicists in Medicine6 (AAPM-TG20) and the European Heavy Particle Dosimetry Group7,8 (ECHED). Both groups recommend the use of air-filled thimble ionization chambers which are calibrated in air in terms of exposure or air kerma in a cobalt-60 gamma ray beam and trace their calibration factors to a national standards laboratory. Similarly to the photon and electron beam dosimetry, proton beam dosimetry is based on the Bragg-

Reprint requests to: Dr. C. Zankowski, McGill University, Montreal General Hospital, Department of Medical Physics, 1650 Cedar Avenue, Montre´al, P.Q, H3G-1A4 288

Proton beam output measurement with an extrapolation chamber ● DR. C. ZANKOWSKI et al.

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Gray cavity theory, and the absorbed dose to the medium, D med is given by: ¯ gp D med 5 Q corrN ygasW

SD S¯ r

med

(1) air

where: Q corr is the charge collected in the cavity air, corrected for temperature, pressure, and ion recombinag tion; N gas is the cavity-gas calibration factor determined # gp is the ratio of mean in a cobalt-60 gamma ray beam; W energies required to produce an ion pair in air for protons [34.3 eV/ion pair9] and for cobalt-60 gamma rays [33.97 eV/ion pair10]; and (S# / r ) med air is the ratio of mean mass collisional stopping powers of the medium to air, averaged over the spectrum of protons at the point of measurement in the medium. Proton beam treatment facility In the early 1990s, the Department of Radiation Sciences at Loma Linda University Medical Center (LLUMC) designed and constructed the first dedicated, hospital-based proton treatment facility in the world. LLUMC operates three proton beam treatment rooms having rotating proton beam gantries, one treatment room with two horizontal beam delivery systems, and a fifth beam room dedicated to physics and radiobiological research. Calibration of the unmodulated monoenergetic 250 MeV proton beam with our extrapolation chamber was conducted using one of the horizontal beam lines of the LLUMC proton therapy facility. Proton beams at LLUMC are generated by a zero gradient synchrotron, capable of accelerating protons in a continuously variable manner to kinetic energies ranging from 70 to 250 MeV. The proton beam is initially accelerated to 2 MeV by a radiofrequency quadrupole linear accelerator and subsequently injected into the synchrotron ring. Shortly after the ring is filled, the beam is accelerated to its final energy in an rf cavity located diametrically opposite to the injection point. The proton beam is then extracted from the synchrotron and directed toward the beam transport lines which guide the beam into the appropriate treatment room. This method permits the treatment of patients with a low-frequency pulsemode duty cycle, providing a beam for approximately 300 ms every 2.2 sec. Beam intensity is approximately 5 3 1010 protons per pulse. The horizontal proton beam enters the treatment room as a narrowly focused monoenergetic pencil beam that is unsuitable for most clinical applications. The pencil beam was made to traverse two lead scattering foils to produce a 15 cm diameter uniform proton field. For clinical use, the proton beam Bragg peak is spread out by modulating the beam energy using a variable thickness Lucite propeller.11 The propeller contains nine sections of Lucite with 0.8 cm thickness increments between adjacent sections. Each section reduces the range of protons in water by approximately 1 cm. The

FIG. 1. Relative electrode separation z as a function of the inverse capacitance (DQ/DV) 21. The solid line represents a least-squares fit to the data and has a slope m of 2.976 mm pF. The electrode area is determined from m/« 0 and equals (3.36 6 0.02) cm2.

propeller rotates at a speed of 310 rpm in order to avoid resonance effects arising from the 60 Hz or higher harmonic structure in the extracted beam. Extrapolation chamber The polystyrene-embedded extrapolation chamber and its use in calibration of photon and electron beams have been described in detail before.2 A 7 cm diameter polystyrene piston was fashioned to move inside a cylinder bored along the center of a 30 3 30 3 8 cm3 polystyrene phantom. Graphite dag was painted on the top surface of the piston and a 1.5 mm deep and 0.04 mm wide groove was cut through the graphite surface into the piston to form the measuring electrode (nominal diameter of 2 cm) and the guard ring of the chamber. The measuring electrode was connected to ground through an electrometer (model 35616; Keithley, Cleveland, OH) and the guard electrode was connected to ground directly. The polarizing electrode consists of a 0.5 mm thick polystyrene disk painted with graphite dag and fastened to the top of the large phantom. The separation between the polarizing and measuring electrodes can vary between 0.5 mm and 10 mm and is controlled by a micrometer mounted to the phantom body. The movement of the piston (i.e., change in the air sensitive volume) is monitored by a calibrated distance travel indicator which is accurate to within 0.002 mm. The sensitive area of the measuring electrode was determined experimentally by plotting the relative electrode separation z as a function of the measured inverse chamber capacitance C. As shown in Fig. 1, the slope of the z vs 1/C relationship yields an effective collecting electrode area of (3.36 6 0.02) cm2 for our extrapolation chamber. In the proton beam dose measurement at the LLUMC the sensitive volume of the extrapolation cham-

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Medical Dosimetry

ber was positioned at a water equivalent depth of 10 cm in a polystyrene phantom. Thus the calibration depth was located in the flat entrance dose region of the 250 MeV monoenergetic proton beam which has a range of 31.8 cm in polystyrene. The field size at the depth of measurement was 15 3 15 cm2. Measurements of the ionization resulting from the proton beam irradiations were carried out at various electrode separations with the electric field maintained at ;400 V/mm. A variable voltage power supply (model 412B; John Fluke, Seattle, WA) capable of providing up to 6 2100 V was used to establish the desired chamber polarizing potential. Measurements were concentrated in the range of electrode separations from 0.50 to 2.50 mm, except for one measurement which was made with an electrode separation of 5.50 mm. The separation of 5.5 mm corresponds to the maximum electrode separation which could be biased at 400 V/mm by our power source. Measurements were repeated for both positive and negative chamber polarities and all readings were corrected for ambient temperature and pressure. The polystyrene extrapolation chamber was subsequently replaced by a Farmer-type cylindrical ionization chamber (model W 30001, PTW, Freiburg, Germany) at the same depth in a polystyrene phantom for independent dose verification and comparison. The cobalt-60 calibration factor of the cylindrical chamber was traceable to a standards laboratory. Ion recombination loss for both the extrapolation chamber and the PTW chamber was estimated using the standard two-voltage technique for continuous radiation.12,13 At LLUMC the proton beam output is reported in terms of dose to muscle; therefore, for the purposes of this study, all doses will be reported as such. RESULTS Determination of proton beam dose rate ˙ poly when The absorbed dose rate in polystyrene D the extrapolation chamber is exposed to proton beam irradiation is given by the modified Bragg-Gray relationship: ˙ poly 5 D

SD

˙ S¯ dQ ¯ pair W dm r

poly

(2) air

˙ /dm is the measured ionization charge gradiwhere: dQ ent per 106 monitor units (mu) corrected for temperature, # pair is the mean energy pressure and recombination loss; W required to produce an ion-pair in air by proton irradiation (34.3 eV/ion-pair); and dm 5 r Adz where: r is the density of air in the chamber sensitive volume, A the area of the collecting electrode, and z the relative electrode separation. To convert the dose absorbed in polystyrene to that which would be absorbed in muscle under identical beam conditions, the dose rate to polystyrene is multiplied simply by the ratio of mass stopping powers of muscle to

Volume 23, Number 4, 1998

FIG. 2. Ionization response of the extrapolation chamber to a 250 MeV monoenergetic proton beam, measured with a field size of 15315 cm2 at a water equivalent depth of 10 cm in polystyrene. Data represent the mean response of the chamber operated at both positive and negative polarities, corrected for temperature, pressure, and ion recombination loss using the two-voltage technique for continuous radiation.

˙ muscle 5 D ˙ poly (S# / r ) muscle polystyrene, i.e., D poly . In this conversion, it is assumed that the proton fluences at the equivalent depths of measurement in polystyrene and muscle are identical. This means that Eq. (2) can be modified to read: ˙ muscle 5 D

SD

˙ S¯ dQ ¯ pair W dm r

muscle

(3) air

˙ /dm ionization gradiwith the understanding that the dQ ent measurements were carried out in polystyrene rather than in muscle. For the 250 MeV proton beam used in our experiment, a mean stopping power ratio of muscle to air equal to 1.148 was used for the depth of 10 cm.1 Figure 2 shows a plot of the extrapolation chamber response to proton irradiation (i.e., the measured ionization charge) as a function of the relative electrode separation averaged over positive and negative chamber polarities and corrected for temperature, pressure and charge recombination loss. The slope of the line in Fig. 2 is (9.247 6 0.016) nC/(106 MU mm). Substituting this ˙ /dz into Eq. (3) results in the absorbed dose value of dQ to muscle of 93.9 cGy/106 MU at a 10 cm depth. Three ionization measurements were carried out with a Farmer thimble ionization chamber at the same depth in polystyrene phantom as in the experiment with the extrapolation chamber, and the mean value of the ionization, corrected for temperature, pressure, and ion recombination, was used in conjunction with Eq. (1). The beam output determined using this method resulted in a dose to muscle of 98.7 cGy/106 MU exceeding the extrapolation chamber result by ;5%. Discussion of uncertainties The AAPM-TG20 protocol states that ionization chambers yield absolute dose measurements in proton

Proton beam output measurement with an extrapolation chamber ● DR. C. ZANKOWSKI et al.

beams with total uncertainties of 6 3%.1 Therefore the difference of 5% between the preliminary extrapolation chamber measurements and the calibrated thimble chamber measurements is not completely unexpected. During the time of our experiments we did not investigate the ion recombination properties of the polystyrene extrapolation chamber exposed to the proton beam. Based on our subsequent work involving the ion recombination in continuous photon beams, we speculate that the use of the continuous beam two-voltage technique for the determination of saturation charge in the quasi-continuous proton beam will underestimate the actual ionization produced in the chamber. Initial recombination and ionic diffusion against the electric field in the chamber may be more important in charge recombination loss for cavities exposed to high-energy proton beams than for cavities exposed to high-energy photon or electron beams; yet initial recombination and ionic diffusion losses have not been included in the analysis of the data obtained in the 250 MeV proton beam. The omission of the contributions of initial recombination and ionic diffusion to the overall charge will certainly result in an underestimation of the actual saturation ionization; however, the exact amount of this effect is currently unknown. Other potential sources of discrepancy between the measurement of dose with the extrapolation chamber and the PTW chamber are the fluence correction factor and the perturbation factor which account for the cylindrical geometry of the PTW chamber but were not used in our dose rate calculations. However, the uncertainty in perturbation effects for the PTW chamber is much smaller than the observed 5% discrepancy, so it cannot be used alone to explain the discrepancy. Since the PTW chamber is in excellent agreement with proton doses determined with calorimetry, we conclude that the 6% discrepancy is, for the most part, attributable to the extrapolation chamber. CONCLUSIONS The uncalibrated, phantom-embedded extrapolation chamber determines the absorbed dose-to-muscle to within 5% of standard calibrated Farmer-type ionization chamber for the 250 MeV proton therapy beam. More work is needed to determine the ion recombination characteristics of the extrapolation chamber exposed to proton beams before a better comparison with standard techniques can be made. Uncertainties in the dosimetry

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of high-energy charged-particle therapy beams preclude accurate evaluation of the performance of the extrapolation chamber at this time, and further work will be required to elucidate the source of discrepancies between proton doses determined with the extrapolation chamber and Farmer chambers. However, the preliminary data indicate that uncalibrated extrapolation chambers, just like in calibrations of photon and electron beams, may prove useful for dose measurements in clinical proton beams.

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