Stefaan Vynckier

Code of practice for clinical proton dosimetry

The objective of this document is to make recommendations for the determination of absorbed dose to tissue for clinical proton beams and to achieve uniformity in proton dosimetry. A Code of Practice has been chosen, providing specific guidelines for the choice of the detector and the method of determination of absorbed dose for proton beams only. This Code of Practice is confined specifically to the determination of absorbed dose and is not concerned with the biological effects of proton beams. It is recommended that dosimeters be calibrated by comparison with a calorimeter. If this is not available, a Faraday cup, or alternatively, an ionization chamber, with a 60Co calibration factor should be used. Physical parameters for determining the dose from tissue-equivalent ionization chamber measurements are given together with a worksheet. It is recommended that calibrations be carried out in water at the centre of the spread-out-Bragg-peak and that dose distributions be measured in a water phantom. It is estimated that the error in the calibrations will be less than +/- 5% (1 S.D.) in all cases. Adoption and implementation of this Code of Practice will facilitate the exchange of clinical information.

Fast 2D phantom dosimetry for scanning proton beams

A quality control system especially designed for dosimetry in scanning proton beams has been designed and tested. The system consists of a scintillating screen (Gd2O2S:Tb), mounted at the beam-exit side of a phantom, and observed by a low noise CCD camera with a long integration time. The purpose of the instrument is to make a fast and accurate two-dimensional image of the dose distribution at the screen position in the phantom. The linearity of the signal with the dose, the noise in the signal, the influence of the ionization density on the signal, and the influence of the field size on the signal have been investigated. The spatial resolution is 1.3 mm (1 s.d.), which is sufficiently smaller than typical penumbras in dose distributions. The measured yield depends linearly on the dose and agrees within 5% with the calculations. In the images a signal to noise ration (signal/1 s.d.) of 10(2) has been found, which is in the same order of magnitude as expected from the calculations. At locations in the dose distribution possessing a strong contribution of high ionization densities (i.e., in the Bragg peak), we found some quenching of the light output, which can be described well by existing models if the beam characteristics are known. For clinically used beam characteristics such as a Spread Out Bragg peak, there is at most 8% deviation from the NACP ionization chamber measurements. The conclusion is that this instrument is a useful tool for quick and reliable quality control of proton beams. The long integration-time capabilities of the system make it worthwhile to investigate its applicability in scanning proton beams and other dynamic treatment modalities.

Monte Carlo evaluation of the AAA treatment planning algorithm in a heterogeneous multilayer phantom and IMRT clinical treatments for an Elekta SL25 linear accelerator

The Anisotropic Analytical Algorithm (AAA) is a new pencil beam convolution/superposition algorithm proposed by Varian for photon dose calculations. The configuration of AAA depends on linear accelerator design and specifications. The purpose of this study was to investigate the accuracy of AAA for an Elekta SL25 linear accelerator for small fields and intensity modulated radiation therapy (IMRT) treatments in inhomogeneous media. The accuracy of AAA was evaluated in two studies. First, AAA was compared both with Monte Carlo (MC) and the measurements in an inhomogeneous phantom simulating lung equivalent tissues and bone ribs. The algorithm was tested under lateral electronic disequilibrium conditions, using small fields (2 x 2 cm(2)). Good agreement was generally achieved for depth dose and profiles, with deviations generally below 3% in lung inhomogeneities and below 5% at interfaces. However, the effects of attenuation and scattering close to the bone ribs were not fully taken into account by AAA, and small inhomogeneities may lead to planning errors. Second, AAA and MC were compared for IMRT plans in clinical conditions, i.e., dose calculations in a computed tomography scan of a patient. One ethmoid tumor, one orophaxynx and two lung tumors are presented in this paper. Small differences were found between the dose volume histograms. For instance, a 1.7% difference for the mean planning target volume dose was obtained for the ethmoid case. Since better agreement was achieved for the same plans but in homogeneous conditions, these differences must be attributed to the handling of inhomogeneities by AAA. Therefore, inherent assumptions of the algorithm, principally the assumption of independent depth and lateral directions in the scaling of the kernels, were slightly influencing AAAs validity in inhomogeneities. However, AAA showed a good accuracy overall and a great ability to handle small fields in inhomogeneous media compared to other pencil beam convolution algorithms.

Dosimetry of beta-ray ophthalmic applicators: Comparison of different measurement methods

An international intercomparison of the dosimetry of three beta particle emitting ophthalmic applicators was performed, which involved measurements with radiochromic film, thermoluminescence dosimeters (TLDs), alanine pellets, plastic scintillators, extrapolation ionization chambers, a small fixed-volume ionization chambers, a diode detector and a diamond detector. The sources studied were planar applicators of 90Sr-90Y and 106Ru-106Rh, and a concave applicator of 106Ru-106Rh. Comparisons were made of absolute dosimetry determined at 1 mm from the source surface in water or water-equivalent plastic, and relative dosimetry along and perpendicular to the source axes. The results of the intercomparison indicate that the various methods yield consistent absolute dosimetry results at the level of 10%-14% (one standard deviation) depending on the source. For relative dosimetry along the source axis at depths of 5 mm or less, the agreement was 3%-9% (one standard deviation) depending on the source and the depth. Crucial to the proper interpretation of the measurement results is an accurate knowledge of the detector geometry, i.e., sensitive volume and amount of insensitive covering material. From the results of these measurements, functions which describe the relative dose rate along and perpendicular to the source axes are suggested.

The radial dose function of low-energy brachytherapy seeds in different solid phantoms: comparison between calculations with the EGSnrc and MCNP4C Monte Carlo codes and measurements.

The use of low-energy photon emitters for brachytherapy applications, as in the treatment of the prostate or of eye tumours, has drastically increased in the last few years. New seed models for 103Pd and 125I have recently been introduced. The American Association of Physicists in Medicine recommends that measurements are made to obtain the dose rate constant, the radial dose function and the anisotropy function. These results must then be compared with Monte Carlo calculations to finally obtain the dosimetric parameters in liquid water. We have used the results obtained during the characterization of the new InterSource (furnished by IBt, Seneffe, Belgium) palladium and iodine sources to compare two Monte Carlo codes against experiment for these low energies. The measurements have been performed in three different media: two solid water plastics, WT1 and RW1, and polymethylmetacrylate. The Monte Carlo calculations were made using two different codes: MCNP4C and EGSnrc. These codes use photon cross-section data of a different origin. Differences were observed between both sets of input data below 100 keV, especially for the photoelectric effect. We obtained differences in the radial dose functions calculated with each code, which can be explained by the difference between the input data. New cross-section data were then tested for both codes. The agreement between the calculations using these new libraries is excellent. The differences are within the statistical uncertainties of the calculations. These results were compared with the experimental data. A good agreement is reached for both isotopes and in the three phantoms when the measured values are corrected for the presence of the TLDs in the phantom.

Is it possible to verify directly a proton-treatment plan using positron emission tomography?

A PET camera is used to visualize the positron activity induced during protonbeam therapy in order to verify directly the proton-treatment plans. The positron emitters created are predominantly the 15O and 11C, whose total activity amounts to 12 MBq after an irradiation with 85 MeV protons, delivering 3 Gy in a volume of approximately 300 cm3. Although this method is a useful verification of patient set-up, care must be taken when deriving dose distributions from activity distributions. Correlation between both quantities is difficult, moreover at the last millimeters of their range, protons will no longer activate tissue. Due to the short half-lives the PET camera must be located close to the treatment facility.

Entrance and exit dose measurements with semiconductors and thermoluminescent dosemeters: a comparison of methods and in vivo results.

BACKGROUND AND PURPOSE In order to compare diodes and TLD for in vivo dosimetry, systematic measurements of entrance and exit doses were performed with semiconductor detectors and thermoluminescent dosemeters for brain and head and neck patients treated isocentrically with external photon beam therapy. MATERIAL AND METHODS Scanditronix EDP-20 diodes and 7LiF thermoluminescent chips, irradiated in a 8 MV linac, were studied with similar build-up cap geometries and materials in order to assure an equivalent electronic equilibrium. Identical calibration methodology was applied to both detectors for the dose determination in clinical conditions. RESULTS For the entrance dose evaluation over 249 field measurements, the ratio of the measured dose to the expected dose, calculated from tabulated tissue maximum ratios, was equal to 1.010 +/- 0.028 (1 s.d.) from diodes and 1.013 +/- 0.041 from thermoluminescent crystals. For the exit dose measurements, these ratios were equal to 0.998 +/- 0.049 and 1.016 +/- 0.070 for diodes and TLDs, respectively, after application of a simple inhomogeneity correction to the calculation of the expected exit dose. CONCLUSIONS Thermoluminescence and semiconductors led to identical results for entrance and exit dose evaluation but TLDs were characterised by a lower reproducibility inherent to the TL process itself and to the acquisition and annihilation procedures.

Calculation of beta-ray dose distributions from ophthalmic applicators and comparison with measurements in a model eye

Dose distributions throughout the eye, from three types of beta-ray ophthalmic applicators, were calculated using the EGS4, ACCEPT 3.0, and other Monte Carlo codes. The applicators were those for which doses were measured in a recent international intercomparison [Med. Phys. 28, 1373 (2001)], planar applicators of 106Ru-106Rh and 90Sr-90Y and a concave 106Ru-106Rh applicator. The main purpose was to compare the results of the various codes with average experimental values. For the planar applicators, calculated and measured doses on the source axis agreed within the experimental errors (<10%) to a depth of 7 mm for 106Ru-106Rh and 5 mm for 90Sr-90Y. At greater distances the measured values are larger than those calculated. For the concave 106Ru-106Rh applicator, there was poor agreement among available calculations and only those calculated by ACCEPT 3.0 agreed with measured values. In the past, attempts have been made to derive such dose distributions simply, by integrating the appropriate point-source dose function over the source. Here, we investigated the accuracy of this procedure for encapsulated sources, by comparing such results with values calculated by Monte Carlo. An attempt was made to allow for the effects of the silver source window but no corrections were made for scattering from the source backing. In these circumstances, at 6 mm depth, the difference in the results of the two calculations was 14%-18% for a planar 106Ru-l06Rh applicator and up to 30% for the concave applicator. It becomes worse at greater depths. These errors are probably caused mainly by differences between the spectrum of beta particles transmitted by the silver window and those transmitted by a thickness of water having the same attenuation properties.

Fluence correction factors in plastic phantoms for clinical proton beams

In recent codes of practice for reference dosimetry in clinical proton beams using ionization chambers, it is recommended to perform the measurement in a water phantom. However, in situations where the positioning accuracy is very critical, it could be more convenient to perform the measurement in a plastic phantom. In proton beams, a similar approach as in electron beams could be applied by introducing fluence correction factors in order to account for the differences in particle fluence distributions at equivalent depths in plastic and water. In this work, fluence correction factors as a function of depth were determined for proton beams with different energies using the Monte Carlo code PTRAN for PMMA and polystyrene with reference to water. The influence of non-elastic nuclear interaction cross sections was investigated. It was found that differences in proton fluence distributions are almost entirely due to differences in non-elastic nuclear interaction cross sections between the plastic materials and water. For proton beams with energies lower than 100 MeV, for which the contributions from non-elastic interactions become small compared to the total dose, the fluence corrections are smaller than 1%. For beams with energies above 200 MeV, depending on the cross sections dataset for non-elastic nuclear interactions, fluence corrections of 2-5% were found at the largest depths. The results could, with an acceptable accuracy, be represented as a correction per cm penetration of the beam, yielding values between 0.06% and 0.15% per cm for PMMA and 0.06% to 0.20% per cm for polystyrene. Experimental information on these correction factors was obtained from depth dose measurements in PMMA and water. The experiments were performed in 75 MeV and 191 MeV non-modulated and range-modulated proton beams. From the experiments, values ranging from 0.03% to 0.15% per cm were obtained. A decisive answer about which dataset for non-elastic nuclear interactions would result in a better representation of the measurements could not be given. We conclude that below 100 MeV, dosimetry could be performed in plastic phantoms without a dramatic loss of accuracy. On the other hand, in clinical high-energy proton beams, where accurate positioning in water is in general not an issue, substantial correction factors would be required for converting dose measurements in a plastic phantom to absorbed dose to water. It is therefore not advisable to perform absorbed dose measurements nor to measure depth dose distributions in a plastic phantom in high-energy proton beams.

RBE variation as a function of depth in the 200-MeV proton beam produced at the National Accelerator Centre in Faure (South Africa).

BACKGROUND AND PURPOSE Thorough knowledge of the RBE of clinical proton beams is indispensable for exploiting their full ballistic advantage. Therefore, the RBE of the 200-MeV clinical proton beam produced at the National Accelerator Centre of Faure (South Africa) was measured at different critical points of the depth-dose distribution. MATERIAL AND METHODS RBEs were determined at the initial plateau of the unmodulated and modulated beam (depth in Perspex = 43.5 mm), and at the beginning, middle and end of a 7-cm spread-out Bragg peak (SOBP) (depths in Perspex = 144.5, 165.5 and 191.5 mm, respectively). The biological system was the regeneration of intestinal crypts in mice after irradiation with a single fraction. RESULTS Using 60Co gamma-rays as the reference, the RBE values (for a gamma-dose of 14.38 Gy corresponding to 10 regenerated crypts) were found equal to 1.16 +/- 0.04, 1.10 +/- 0.03, 1.18 +/- 0.04, 1.12 +/- 0.03 and 1.23 +/- 0.03, respectively. At all depths, RBEs were found to increase slightly (about 4%) with decreasing dose, in the investigated dose range (12-17 Gy). No significant RBE variation with depth was observed, although RBEs in the SOBP were found to average a higher value (1.18 +/- 0.06) than in the entrance plateau (1.13 +/- 0.04). CONCLUSION An RBE value slightly larger than the current value of 1.10 should be adopted for clinical application with a 200-MeV proton beam.