T R E A T M E N T   D E L I V E R Y   V E R I F I C A T I O N   U S I N G   A    
P E L V I C   A N T H R O P O R M O P H I C   P H A N T O M 
 
 
 
 
 
 
 
 
 
Nsangu Augustine Mwale 
 
 
 
 
 
 
 
 
 
A research report submitted to the Faculty of Science, University of the Witwatersrand, 
Johannesburg, in partial fulfillment of the requirements for the degree of Master of 
Science. 
 
 
 
 
 
 
Johannesburg, 2009
 ii 
DECLARATION 
 
I declare that this research is my own, unaided work. It is submitted for the degree of 
Master of Science at the University of the Witwatersrand, Johannesburg. It has not been 
submitted before for any degree or examination in any other University. 
 
 
 
________________________________________ 
(Signature of candidate) 
 
________________ day of _________________________2009 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 iii 
ABSTRACT 
 
Aim: A series of measurements in a pelvic anthropomorphic phantom was performed as 
part of the validation of the entire radiotherapy treatment chain. Two treatment planning 
calculation algorithms were used: the Pencil Beam (PB) and the Collapsed Cone (CC). 
The dose calculation algorithms of Radiotherapy Treatment Planning Systems (RTPS) 
were validated to ensure that the dose delivered to a treatment target was accurately 
predicted. An anthropomorphic phantom, designed and manufactured locally at the 
Pretoria Academic Hospital, was employed in this study. The phantom was fabricated 
with locally available materials as human tissue substitutes based on the attenuation 
coefficients, electron densities and effective atomic numbers.  
 
Materials: Pelvic anthropomorphic phantom, Thermoluminescent Dosimeters (TLD_100 
chips), X-Omat V film, film processors, a densitometer and 6 MV and 15 MV linear 
accelerator photon beams with beam quality (TPR20,10) of 0.674 and 0.763 respectively. 
 
Results: Two treatment planning techniques were studied, a four field ?box? and parallel 
opposed beams using a local cancer of the cervix protocol. Point doses calculated by the 
RTPS were compared with equivalent point dose values measured with 
thermoluminescent dosimeters (TLDs). Three dose regions emerged for the four field 
technique, those of low, intermediate and high dose gradient. The four field technique for 
6 MV gave a dose deviation from -4.9% to -32.1% and at 15 MV from -1.5% to -20.6%. 
For 6 MV and 15 MV parallel opposed beams, percentage dose deviations from -2.7% to 
-11.8% and from +0.2% to -11.5 were observed. The mean value of the ratios of 
measured to calculated dose values was 0.91?0.05 for the four field technique and 
0.94?0.02 for the AP/PA. Radiographic film was used to compare the predicted 2D 
isodose distributions to the actual dose distribution in the phantom. The 2D isodose 
distributions obtained were not meaningful in comparing the doses predicted by the 
planning system. A smaller field size of 7 cm x 7 cm was also employed and results of 
both TLD and film obtained were comparable to those predicted by the planning system. 
 
 iv 
Conclusion: The stated goal of dose delivery accuracy (ICRU, 1987) to within 5% was 
not generally met in this study. On average the measured doses using TLDs and film at a 
field size of 7 cm x 7 cm were lower than the point doses predicted by the RTPS dose 
calculation algorithms whereas the film over-responded when a local cancer of the cervix 
protocol was employed. At a field size of 7 cm x 7 cm, film dosimetry was comparable to 
the TLD results. Film and TLDs were calibrated perpendicularly and exposed parallel. 
The phantom is unsuitable for film dosimetry studies at field sizes more than 14 cm x 14 
cm. 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 v 
In Memory of my late parents,  
Elson Joseph Mwale (1956 -1999) and Elizabeth Sakala Mwale (1956 -2000) 
 
who made tremendous sacrifices to ensure that I had a good education, for this and much 
more, I am forever in their debt. 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 vi 
ACKNOWLEDGEMENTS 
 
I would like to thank the International Atomic Energy Agency (I.A.E.A) for the financial 
support during my two year internship training at the Johannesburg Hospital, the Canon 
Collins Trust, my supervisor Professor Debbie Van Der Merwe for suggesting this work 
and helping with developing the treatment plans. I am grateful to The Division of 
Radiation Oncology, Johannesburg Hospital for making the equipment available for this 
experiment, to Prof A. J. Van Rensburg, Pretoria Academic Hospital for helping with the 
film scanner, to Nico Van Der Merwe for providing the densitometer, to Lindiwe for 
developing the cast, and the Johannesburg Hospital Radiation Oncology staff, who have 
been absolutely kind. 
 
I would also like to thank my family for moral support, especially my brothers Steve and 
Chitalu, my sisters Ann and Linda. I love you. You are the semantics behind all that I do. 
 
 
 
 
 
 
 
 
 
 
 
CONTENTS 
 
ABSTRACT....................................................................................................................... iii 
ACKNOWLEDGEMENTS............................................................................................... vi 
LIST OF TABLES........................................................................................................... viii 
LIST OF FIGURES ........................................................................................................... ix 
1.0 INTRODUCTION???????????????????????????????????.1 
1.1 Objectives/Rationale of Study .................................................................................. 3 
2.0 LITERATURE REVIEW ............................................................................................. 4 
2.1 Interaction of photons with matter............................................................................ 4 
2.1.1 Energy deposition by photon beams .................................................................. 5 
2.1.2 Kerma................................................................................................................. 6 
2.1.3 Absorbed dose.................................................................................................... 7 
2.1.4 Primary and scattered dose components ............................................................ 7 
2.2 Experimental quantities ............................................................................................ 7 
2.3 Treatment planning system....................................................................................... 8 
2.3.1 Monitor Unit (MU) Calculation....................................................................... 10 
2.4 Film dosimetry........................................................................................................ 11 
2.5 Thermoluminescent dosimetry................................................................................ 14 
2.6 Computed Tomography .......................................................................................... 17 
2.7 The radiotherapy procedure .................................................................................... 19 
2.7.1 Beam combinations.......................................................................................... 19 
2.7.2 Single photon beam.......................................................................................... 19 
2.7.3 Parallel opposed beams.................................................................................... 20 
2.7.4 Four field ?box? ............................................................................................... 20 
2.7.5 Treatment verification...................................................................................... 20 
3.0 METHODOLOGY ..................................................................................................... 22 
3.1 Phantom fabrication ................................................................................................ 22 
3.2 TLD Procedure........................................................................................................ 24 
3.3 Calibration and irradiation process ......................................................................... 24 
3.4 Film calibration....................................................................................................... 27 
3.5 Dose calculation on Helax-TMS............................................................................. 29 
4.0 RESULTS ................................................................................................................... 36 
5.0 DISCUSSION............................................................................................................. 41 
6.0 CONCLUSION AND RECOMMENDATIONS ....................................................... 44 
REFERENCES.??????????????????????..????????.47 
APPENDIX A?????????????????????????????51 
APPENDIX B?????????????????????????????59 
 
  
 
 
 
 
 
 viii 
LIST OF TABLES 
 
TABLE 3.1:  Treatment protocol showing the number of MUs needed to deliver 
0.7 Gy to the isocenter for no inhomogeneity, manual calculation, 
PB algorithm and CC algorithm. 
TABLE A.1: Individual TLD results for 6 MV AP/PA compared to point doses 
predicted by the planning system when the local cancer of the cervix 
protocol is used. 
TABLE A.2: Individual TLD results for 6 MV four field ?box? compared to 
point doses predicted by the planning system when the local cancer 
of the cervix protocol is used. 
TABLE A.3: Individual TLD results for 15 MV AP/PA compared to point doses 
predicted by the planning system when the local cancer of the cervix 
protocol is used. 
TABLE A.4: Individual TLD results for 15 MV four field ?box? compared to 
point doses predicted by the planning system when the local cancer 
of the cervix protocol is used.   
TABLE A.5 Individual TLD and film results for 6 MV AP/PA treatment plan 
based on figure 4.9 compared to point doses predicted by the 
planning system. 
TABLE A.6 Individual TLD and film results for 6 MV four field ?box? 
treatment plan based on figure 4.10 compared to point doses 
predicted by the planning system. 
TABLE B.1 This table shows Kanduza?s results at 6 MV using the AP/PA  
technique Kanduza (2005). 
 
 
 
 
 
 
 
 ix
 LIST OF FIGURES 
 
FIGURE 2.1  Gurney and Mott model of latent image. AgBr remains in the ionic 
form (Ag+Br?) in the crystal of the grain .Radiation produces 
ionization of Br? to Br+e. These electrons make the speck 
negatively charged. The Ag+ migrate to neutralize the speck and 
forms a lump of Ag (aggregate) on the speck (Pai et al, 2007). 
FIGURE 2.2 A simplified energy diagram to illustrate thermoluminescence 
process (Khan, 2003). 
FIGURE 2.3  Diagram showing apparatus for measuring thermoluminescence 
(Khan, 2003). 
FIGURE 3.1   Diagnostic imaging insert 1 (Kanduza, 2005). 
FIGURE 3.2   Dosimetric insert 2 (Kanduza, 2005). 
FIGURE 3.3   Position of film and TLDs in the dosimetric insert (Kanduza, 2005) 
FIGURE 3.4  Variation of TLD output (Gy) with prescribed dose (Gy) at 6 MV. 
The error bars represent the overall deviation in the measurements 
FIGURE 3.5  Variation of TLD output (Gy) with prescribed dose (Gy) at 15 MV.      
The error bars represent the overall deviation in the measurements. 
FIGURE 3.6  Sensitometric curves at 6 MV and 15 MV for the perpendicular 
setup. The uncertainties were very small, hence error bars not 
visible. 
FIGURE 3.7  Pencil Beam parallel opposed dose distribution at the level of the 
isocenter using 6 MV beams. 
FIGURE 3.8 Collapsed Cone parallel opposed dose distribution at the level of 
the isocenter using 6 MV beams. 
FIGURE 3.9  Pencil Beam parallel opposed dose distribution at the level of the 
isocenter using 15 MV beams. 
FIGURE 3.10  Collapsed Cone parallel opposed dose distribution at the level of 
the isocenter using 15 MV beams. 
FIGURE 3.11  Pencil Beam four field ?box? dose distribution at the level of the 
isocenter and the position of the cut film using 6 MV Beams. 
 
 x
 FIGURE 3.12  Collapsed Cone four field ?box? dose distribution at the level of 
the isocenter using 6 MV beams. 
FIGURE 3.13  Pencil Beam four field ?box? dose distribution at the level of the 
isocenter using 6 MV beams.  
FIGURE 3.14  Collapsed Cone four field ?box? dose distribution at the level of 
the isocenter using 15 MV beams.  
FIGURE 3.15 Pencil Beam parallel opposed dose distributions of a 7 x 7 cm field 
size at the level of the isocenter using 6 MV beams.  
FIGURE 3.16 Pencil Beam four field ?box? dose distributions (7 x 7 cm box) at 
the level of the isocenter using 6 MV beams. 
FIGURE 4.1  The comparison between the dose measured by the TLD?s using 
the AP/PA technique to the dose calculated by the PB algorithm at 
6 MV. 
FIGURE 4.2  A comparison between the dose measured by the TLDs and that 
calculated by the PB algorithm using the four field ?box? 
technique at 6 MV. Three dose regions emerged, those of low, 
intermediate and high dose gradient. 
FIGURE 4.3  A comparison of measured dose with that calculated by the 
planning system for the AP/PA technique at 15 MV using TLDs. 
The PB algorithm was used to calculate dose. 
FIGURE 4.4  A comparison between the dose measured by the TLDs and that 
calculated by the PB algorithm using the four field ?box? 
technique at 15 MV. Three dose region emerged, those of low, 
intermediate and high dose gradients. 
FIGURE 4.5  Ratio of measured to the calculated doses vs calculated doses for 
AP/PA at 6 MV using TLDs. The PB algorithm was used to 
calculate the doses. 
FIGURE 4.6  Ratio of measured to the calculated doses vs calculated doses for 
the four field ?box? at 6 MV using TLDs. The PB algorithm was 
used to calculate doses. 
 
 xi
 FIGURE 4.7  Ratio of measured to the calculated doses vs calculated doses for 
AP/PA at 15 MV using TLDs. The PB algorithm was used to 
calculate doses. 
FIGURE 4.8  Ratio of measured to the calculated doses vs calculated doses for 
four field ?box? at 15 MV using TLDs. The PB algorithm was 
used to calculate doses. 
FIGURE A.1 Figure A.1 2-D isodose distribution at 15 MV for the four field box 
technique, film was cut and positioned as shown in figure 4.5 
FIGURE A.2 Figure A.2 Shows an exposed film (using a 7 x 7 cm field size) at 6 
MV for the four field box technique, film was cut and positioned 
figure 4.10. 
 
  
 
 1 
1.0 INTRODUCTION 
 
Radiation oncology is a clinical process that uses radiation for the treatment of cancer. 
The probability of success in radiotherapy with external photon beams depends strongly 
on delivering an adequately high dose to the intended target volume and at the same time, 
limiting dose to normal tissues close to the tumor (Delich et al, 1999). This results in the 
eradication of the disease, the propagation of life and/or the improvement in the quality 
of life. For some tumors and normal tissues, the response curve is very steep and a small 
change in dose can result in a large change in the probability of tumor control and normal 
tissue complications. This leads to strong demands on the accuracy in dose delivery. The 
International Commission on Radiation Units and Measurements (ICRU) (ICRU, 1987) 
has recommended that the dose must be delivered to within 5% of the treatment 
prescription. In order to satisfy this recommendation, each step involved in the 
radiotherapeutic process has to be performed at an accuracy better than 2% (ICRU, 
1976). Other reviews have suggested accuracies in dose delivery of ?3.5% (Brahme, 
1984). As part of the overall uncertainty arises from the process of dose calculations in 
treatment planning, the tolerances for the accuracy of a Radiotherapy Treatment Planning 
Systems (RTPS) have to be appropriately smaller (Mijnheer et al, 1987). To meet these 
delivery standards, quality assurance (QA) for RTPS is required. One of the most 
relevant parts of the RTPS QA is verification of the dose calculation. 
 
An accurate knowledge of the distribution of the dose within patients undergoing external 
beam radiotherapy has been an essential requirement for the efficient planning and 
verification of treatment regimes for many years. As a result, many different materials 
have been used to simulate the body/ body sections during dose measurements involving 
therapeutic radiation beams. Modern treatment planning techniques use highly advanced 
tools to conform the dose distribution to the target, e.g., asymmetric beams, irregular 
beams shaped by multileaf collimators (MLC) and noncoplanar beams. A RTPS requires 
input of several physical and dosimetric data related to the radiotherapy treatment 
machine (Percentage Depth Dose (PDD) curves, off axis scans, etc) before actual use in 
patient treatment dosimetric evaluations. The RTPS has to be tested thoroughly by means 
 
 2 
of both homogeneous and anthropomorphic phantoms, so that knowledge of the 
simulation capabilities of the computerized system can be obtained (Dunscombe et al, 
1996). The testing procedure is mandatory because the accuracy in computing the dose 
distribution is influenced by the raw data describing the characteristics of the input 
radiation beams.  
 
Ideally, confirmation of adequate performance of a RTPS should be based on the 
comparison of measured and calculated doses in patients receiving treatment as this is the 
situation for which the system is designed. In practice of course this is not feasible 
directly and validation has to rely on confirmation of the system?s accuracy by comparing 
calculation and measurement under experimental conditions that, to a greater or lesser 
degree, reflect the clinical situation. Uncertainties in dose delivery may be introduced at 
the treatment phase (including machine calibration) or during the process of deriving 
monitor unit (MU) or timer settings from the dose prescription determined during 
treatment preparations. The latter type, arising at the treatment planning phase, could 
potentially affect the whole course of treatment and therefore are of particular concern. 
The MU settings required to deliver the prescribed dose are often calculated by a 
computerized treatment planning system using methods and quantities different from 
those in manual MU calculations. For computerized calculations of MU, whether or not 
accompanied by a dose distribution, uncertainties may be further categorized as arising 
from the input data, the calculation algorithm, incorrect use of the system or data transfer 
to the treatment sheet (Leszczynski and Dunscombe, 2000). MU calculations must also 
be compared with point/absolute dose (e.g. Gy) calculations as significant discrepancies 
of up to 5% have been observed (Starkschall et al, 2000). When comparing RTPS-
 calculated dose values with measured dose values, the discrepancies observed may be 
attributed to measurement error or to the RTPS calculation algorithms. Possible problems 
with the RTPS calculation include physical limitation of the beam model used in the 
dose-calculation matrix size, incorrect beam data and limitations in the computer code of 
the dose-calculating algorithm (Herrick et al, 1999). 
 
 
 3 
There are several sources of uncertainties which affect the accuracy of planning and 
treatment. These are: (a) Patient localization, (b) Imaging, (c) Definition of anatomy,    
(d) Establishment of beam geometry (e) Dose calculations, (f) Dose display and plan 
evaluation, (g) Plan implementation, (h) Treatment delivery (Fraass et al, 1998). 
 
In routine clinical practice, the more likely sources of systematic dose errors for 
individual patients result from misuse systems e.g. through inadequate understanding of 
normalization protocols, misinterpretation of the system output or data transfer errors 
(Leszczynski and Dunscombe, 2000). 
 
1.1 Objectives/Rationale of Study 
 
The high incidence of cervical cancer in sub-Saharan Africa has led to the need for a 
comprehensive set of quality assurance (QA) guidelines that can be applied to relatively 
simple clinical treatment planning. The aim of this study was to employ a pelvic 
anthropomorphic phantom (Kanduza, 2005) to verify the quality and consistency of 
treatments by mimicking the many steps involved in a radiotherapy process. The RTPS 
(Helax TMS, 2003) dose calculation algorithms were compared with measured doses in 
the pelvic anthropomorphic phantom for parallel opposed and four field ?box? techniques 
at two different energies. Should this result contribute to verifying the overall integrity of 
the radiation therapy chain, a procedure could be developed for international dosimetry 
audits in developing countries for conventional techniques. The latter have to date, been 
restricted to absolute dosimetry protocols only. 
 
 
 
 
 
 
 
 
 
 4 
2.0 LITERATURE REVIEW 
 
2.1 Interaction of photons with matter 
 
When a beam of photons enters a patient, energy is transferred from the beam and 
deposited within the patient as a result of photon interactions. Photon interaction in 
radiotherapy is dominated by three competing processes involving the atoms of the 
tissues: the photoelectric effect, the Compton effect and pair production. In water-like 
tissues with effective atomic number (Zeff) of approximately 7.5, the probability of 
Compton events dominates (>80%) for energies between 0.05 MeV and 10 MeV thus the 
accurate modeling of Compton scattering events is an essential ingredient of any method 
of inhomogeneity correction for megavoltage photon beams. For higher atomic number 
materials, such as bone (Zeff=13), the energy range of Compton dominance is reduced 
(0.08 to 7 MeV). The photoelectric effect depends on the photon energy and the atomic 
number (Z) of the material. The probability of a photoelectric interaction decreases as 
(1/Eo3) and Z3. As a result the dose to bone relative to the dose to water will be increased, 
for a given photon fluence (Khan, 2003., Papanikalaou et al, 2004). The initial photon 
interactions within tissue are characterized by a linear attenuation coefficient (?) which is 
the probability of a photon interacting per unit distance. The probability depends on: (a) 
the incident photon energy (MeV), (b) the tissue density (g/cm3) and the (Zeff) of the 
tissue (Khan, 2003). 
 
The main aim of dosimetry in radiotherapy is ?to measure or predict the absorbed dose in 
various tissues of a patient undergoing radiotherapy? (Metcalfe et al, 1997). This 
involves two steps: first, the assessment of the absorbed dose by a radiation detector in a 
suitable phantom and second, the use of beam and patient data such as tumor location and 
patient contours to predict dose at any point in the patient. It is generally assumed that a 
5% difference in dose delivered to a patient does make a clinically detectable difference 
in treatment outcome (Metcalfe et al, 1997). For example an under dosage of a tumor by 
5% may reduce the tumor control probability by more than 15% (Karzmark, 1993). 
 
 5 
Absorbed dose is strictly defined as the mean energy imparted (by ionizing radiation) per 
mass (ICRU, 1998) i.e. dose is decoupled from the radiation used to deliver it. Therefore, 
absorbed dose is the fundamental physical quantity of interest for relating radiation 
treatment to its outcome. The broad spectrum of events that impart energy to matter 
implies that direct measurement of dose from its definition is not a trivial task. 
Calorimeters and ionization chambers can be used to measure dose in absolute terms but 
are not suitable for in vivo dosimetry. TLDs and diodes, placed on the patient?s surface or 
within cavities, are used to check entrance or exit-dose measurements and the delivered 
dose within patients, but they are not suitable for obtaining an in-vivo map of the dose. 
Entrance dose measurements can provide confirmation of machine calibration and patient 
position relative to the accelerator, whilst exit dose measurements can also confirm beam 
alignment and the radiological thickness of the patient. 
 
2.1.1 Energy deposition by photon beams 
 
The photons from a treatment machine yield a cascade of interactions, not only in the 
patient but also in the treatment machine itself before the energy is absorbed as dose. 
Irradiation within the treatment head adds a secondary photon and electron component to 
the primary beam. As such the build up region is a contribution from contaminant 
charged particles in the treatment machine head and the air column between the head and 
the irradiated medium. Starting at the source (electron beam target), most photons 
entering the patient have not been subject to any interactions before entering the patient 
and will serve as the originators of the primary and phantom scatter dose distributions. 
Particles interacting in the treatment head yield two dose categories: charged particle 
contamination and head scatter dose. Head scatter dose accounts for approximately 5 ? 
15% of the total dose depending on the energy (Ahnesjo, 1994). The dosimetric data used 
in treatment planning are mainly derived from measurement in water. Fano?s theorem 
states that when an object of varying density but constant atomic composition is exposed 
to a uniform photon fluence (under charged particle equilibrium (CPE)), differential in 
energy and direction, then the fluence of charged particles launched by the photons is 
also constant and independent of density variations (Fano, 1954). This constant fluence of 
 
 6 
secondary electrons equals the fluence in CPE for a given fluence of photons. This 
implies that the absorbed dose across any area of density variation would be constant. 
Fano?s theorem is an important test of dose calculation algorithms (Nilson and Knoos, 
1992). To apply Fano?s theorem to external radiation beams, the primary photon 
attenuation, stopping power density effect and release of secondary photons can be 
neglected in an equilibrium layer surrounding a dose point of interest. While Fano?s 
theory applies to situations of charged particle equilibrium, the density scaling theorem 
by O?Connor (O?Connor, 1957) relates the dose in two media of different density but 
equal atomic composition, both irradiated by the same external beam, to each other. 
According to this theorem, the ratio of the secondary scattered photon fluence to that of 
primary photon fluence is constant in the two media provided all geometric distances, 
including field sizes are scaled inversely with density. This means that the dose at 
corresponding points in the two media is the same if all dimensions in the irradiation 
geometries are scaled inversely with density.  Both Fano?s and O?Connor?s theorems rely 
on a common assumption that the interaction probability (per electron) is independent of 
density variations between media. 
 
2.1.2 Kerma 
 
The deposition of energy in tissue from a photon beam is fundamentally a two-step 
process; KERMA and absorbed dose: 
 
1. The photons interact in the medium to impart kinetic energy to charged particles 
(e.g. the TERMA-total energy released in medium) 
2. Charged particles then deposit their given energy through ionization and 
excitation events along a finite track (e.g. the dose step) 
 
Kerma stands for the Kinetic Energy Released in the Medium 
 
( / )tr
 energyK dE dm
 mass
 
 
 =    
 
 
 7 
where dEtr is the kinetic energy transferred from photons to electrons in a volume 
element of mass dm. It is the quantity that most directly connects the description of the 
radiation beam with its effects (Johns and Cunningham, 1983). 
 
2.1.3 Absorbed dose 
 
This is the quantity that is of more interest in radiotherapy and radiobiology. The 
absorbed dose is the energy actually retained in the medium and is brought about through 
the ionization and excitation events along the finite secondary particle tracks (Johns and 
Cunningham, 1983). 
 
2.1.4 Primary and scattered dose components 
 
One basic physical principle which has been used for many years in photon dose 
calculations involves partitioning the dose delivered to each volume element into two 
components ? primary and scatter. The photons that enter through the surface of the 
patient or phantom are referred to as incident or primary photons. The first time such 
photons interact within the medium, they contribute to primary dose through secondary 
electrons. The scattered component is that part of the dose which is deposited by photons, 
which have previously interacted at least once in the medium. These two components 
obey the same physical laws (divergence, attenuation, generation of secondary photons 
and electrons), but have very different source distributions and energy spectra (Swinnen, 
2005). Dose calculation models which achieve separation of these components have 
gained considerable success when calculating complex treatment plans. 
 
2.2 Experimental quantities 
 
To describe the penetration characteristics of photon beams, three quantities have been 
widely used: the percentage depth dose (PDD), the tissue phantom ratio (TPR) and the 
tissue maximum ratio (TMR). PDD data are impractical for direct reconstruction of dose 
distributions since they depend on the source to surface distance (SSD). Instead TPR?s 
 
 8 
being independent of SSD, have gained popularity. Another quantity, TMR has been used 
in some dosimetry systems (Khan, 2003). TMR is renormalized TPR data such that the 
specified reference depth is the depth of maximum dose. The uncertainties due to electron 
contamination of measurements at the depth of dose maximum is a complication, and use 
of a reference depth away from the build up region as in the TPR definition, is a better 
choice for dosimetry systems (Dutriex et al, 1997). Early dosimetric systems have tried to 
separate scatter from the primary dose using scatter factors to express the ratio of the total 
to primary dose at a point. 
 
2.3 Treatment planning system 
 
The functionality and quality of any RTPS is dependent on the type of algorithms used in 
the different steps of the planning process. Knowledge at some level of the various 
algorithms used within the RTPS can help the user understand the capabilities and 
limitations of the specific systems. Dose calculation models should serve, within the 
environment of a RTPS, to provide quick and accurate results for all relevant types of 
treatment field arrangements. RTPS require input data to their beam models for 
generating isodose curves and calculating monitor units. These data are generally 
obtained with a scanning water phantom system. The RTPS (Helax-TMS, 2003) at 
Johannesburg Hospital has two dose calculation algorithms available for external beam 
planning, the pencil beam (PB) and the collapsed cone (CC) algorithms. Dose 
calculations for patients on the Helax-TMS use electron density information from 
Computed Tomography (CT), based on the conversion from Hounsfield numbers (HN) , 
thus accounting for the attenuating properties of different tissues, as well as providing 
information on the material composition. Modern treatment planning techniques use 
highly advanced tools in order to conform the dose distribution to the target. The 
requirements for achieving accurate dose calculations have resulted in RTPS?s utilizing 
calculations based on convolution techniques to achieve high accuracy as cited by Weber 
and Nilson (Weber and Nilsson, 2002). Any algorithm needs to take into account several 
different dose components, e.g. primary dose, dose from charged particle contamination, 
head scatter dose, and phantom scatter dose. For simplicity and speed in dose 
 
 9 
calculations, models based on a PB kernel are frequently used for dose calculations. 
These kernels are generated by integrating Monte Carlo (MC) calculated point kernels 
(energy deposition in an infinite medium around a primary photon interaction point) over 
depth. The PB model has however been shown to have some limitations in heterogeneous 
media (Knoos et al, 1995). Lack of lateral scatter from heterogeneities, as well as 
disregard for the loss of electron equilibrium, may lead to specific errors in for example, 
the thoracic region. Other draw backs are due to the spatially invariant pencil kernel 
leading to an underestimation of the output for larger fields, whereas charged particle 
contamination models in some situations, lead to unwanted overestimations of dose in the 
build up region implying problems for in vivo dosimetry (Knoos et al, 1994). PB kernels 
have however, long found widespread application in electron beam dose calculation 
algorithms. The CC method is a specific approximation to the point kernel approach. The 
kernels are discretized in a number of directions, unevenly distributed in angle with a 
concentration in the forward direction where most of the photons are scattered (Ahnesjo, 
1997). For each direction, the kernel h is analytically described by an exponential over r2, 
r being the distance from the point of interaction, i.e. 
 
2( ) ( ) /ar brh r Ae Be r? ?= +  
 
Where A. a, B, and b are fitting parameters depending on the scattering angle. The dose 
deposited during the ray trace is distributed into the patient volume according to the 
specification in the point kernels. All energy emitted in a solid angle cone is assumed to 
be transported along the cone axis, hence the name collapsed cone. The two algorithms 
differ primarily in how they model radiation transport and calculate dose in 
heterogeneous media, with the CC algorithm better approximating dose to the medium 
(Ahnesjo, 1989). It has been shown that there is little or no clinically significant 
difference in the calculation of dose by the two algorithms when no large non-unit 
density heterogeneities are present and under full scatter conditions, such as in the 
irradiation of the pelvis (Aspradakis, 2003). However treatment planning calculations of 
the chest wall are complex due to missing tissue, the thinness of the chest wall and the 
 
 10 
presence of lung. The accuracy of calculated dose is in this case dependent on the type of 
algorithm employed (Aspradakis, 2006). 
 
2.3.1 Monitor Unit (MU) Calculation 
 
The calculation of the number of monitor units for the individual fields of a treatment 
plan is included in most modern RTPSs, however it is essential for the user of a RTPS to 
understand the principles of the MU calculation algorithm. Generally the MU calculation 
is closely linked to the calculation of the relative dose distribution. Due to the wide 
variation, both in monitor unit calculation algorithms and in clinical situations, it is 
difficult to indicate how much effort should be spent by each institution in testing 
algorithms for monitor unit calculations. Factors affecting these calculations and related 
quality assurance programmes have been discussed by Kutcher et al, as cited by Dutreix 
et al (Dutreix et al, 1997).  Monitor unit calculations must also be compared with 
point/absolute dose (e.g. Gy) calculations as significant discrepancies of up to 5% have 
been observed (Starkschall et al, 2000). The proposed method for calculating MUs or 
dose (D) in a wedged asymmetric field, assuming an isocentric set-up (SAD = 100 cm) 
and treatment unit calibration k = 1 cGy/MU at the prescription point for a 10 cm x 10 cm 
reference field size at a source to calibration distance SCD = 100 + to (to is the maximum 
depth of maximum dose (Khan, 2003 and Mihailidis, 2000)) is given by 
 
2
 (prescription)/      
* ( )* ( )* ( , )*( / ) * ( , )* ( , , )*c d d d w
 MU D q where
 q k Sc r Sp r TMR d r SCD SAD WF d r POAR x y d Cf
 =
 =
  
 
where Sc(rc) is the collimator scatter factor for collimator field size rc at central axis, 
Sp(rd ) is the phantom scatter factor for equivalent size rd at central axis, TMR (d, rd ) is 
the tissue-maximum ratio for the open field at central axis, WF (d, rd ) is the depth and 
field size dependent wedge factor at central axis, POARw is the depth-dependent primary 
wedged off-axis ratio defined in open asymmetric fields. POARw is the ratio of primary 
dose at the off-axis point of interest to the primary dose at the central axis at the same 
 
 11 
depth, for a symmetrically wide open field (Khan, 2004). Cf is the additional correction 
factor for attenuators such as block trays or compensators and the inverse square 
correction is performed from the calibration distance, SCD, to the treatment distance, 
SAD = SSD + to. The changes in the primary dose at a lateral distance from the central 
axis due to changes in the beam intensity are given by the factor POARw. However unlike 
open asymmetric fields, wedged asymmetric fields have a primary dose profile that is not 
created by the flattening filter alone. Thus, there is no radial symmetry of the POARw. 
POARw is then a function of the off-axis distance and depth, POARw(x, y, d). 
 
2.4 Film dosimetry 
 
Film dosimetry is based upon the calibration of the film response (Optical Density) at a 
specific depth in a phantom and then converting the film density to dose by using a 
calibration curve (Yeo et al, 1999). The mass energy absorption coefficient of typical x ?
 ray films increases as the photon energy decreases. Softening of the photon spectra with 
depth in phantoms results in an over response as a function of depth (Burch et al, 1997). 
Radiographic films are a valuable tool in 2D dosimetry. Film offers several advantages 
such as high spatial resolution, 2D information with one single irradiation, a permanent 
record of the integrated planar dose distribution and suitability for use in inhomogeneous 
phantoms. Typical radiographic film consists of a radiation sensitive emulsion coated on 
a transparent polyester base. The emulsion consists of silver halide crystals (typically 
95% silver bromide and 5% silver iodide suspended in gelatin as in the case of Kodak 
XV and XTL films). When the emulsion is exposed to ionizing radiation, ionization takes 
place in the silver halide film crystals that result in the formation of a latent image. The 
polyester film base is typically 0.2 mm thick and free from significant optical defects or 
impurities. The sensitive component, the emulsion, consists of gelatin and silver halide 
grains, which are typically 1 ? 3 ?m in dimension. There is a significant change that takes 
place in silver halide grains when the photographic emulsion is exposed to light, x rays or 
charged particles. Only adequately exposed grains will be developed and the remainder is 
left largely undeveloped, i.e. the small fraction that creates a low level darkening on the 
film referred to as ?fog?. The theory of how this happens has not yet been fully 
 
 12 
understood but Gurney and Mott have described a mechanism of latent image formation 
as shown in Figure 2.1 (Pai et al, 2007). 
 
 
Figure 2.1 Gurney and Mott model of latent image. AgBr remains in the ionic form (Ag+Br?) in 
the crystal of the grain .Radiation produces ionization of  Br? to Br+e. These electrons make the 
speck negatively charged. The Ag+ migrate to neutralize the speck and forms a lump of Ag 
(aggregate) on the speck. 
 
When the grain is ionized by radiation, the Br- ions are split into Br and electrons: 
 
( )Br Br e? ?? +  
 
The electrons liberated migrate towards the specks, making the speck negatively charged, 
which in turn, attracts Ag+ ions toward the speck, forming a latent image. Once the grains 
are developed, the grains comprising the latent image are converted to metallic silver, 
producing the dark regions on the film.  Film developing includes four steps: developing, 
fixing, washing and drying. The value of film opaqueness is quantified through the light 
transmission factor (T) and is measured by a quantity called optical density (OD). The 
OD is a value describing the darkness of a film and is measured with a device known as a 
densitometer. 
 
10 10log ( ) log ( )oOD T I I= ? =  
 
 
 13 
Where Io is the incident light intensity measured in the absence of film and I is the 
intensity transmitted through the film perpendicular to the film plane. In dosimetry, the 
quantity of interest is the net optical density, defined as a measure of the difference 
between the unexposed processed film from the measured optical density. An extensive 
description of the relationship between optical density, grain size, and mechanism of 
photon interactions that relates to the optical density has been reported by Dainty and 
Shaw as cited by Pai et al (Pai et al, 2007). If 
 is the average area (cm2/grain) of the 
developed silver grain and if there are n developed grains/cm2 of film, then T can be 
written as  
 
10log ( ) 0.4343
 n
 n
 T e
 OD e n
 ?
 ? ?
 ?
 ?
 =
 = ? =
  
 
It is difficult to know the number of electrons needed to develop a grain, but if we assume 
that a single electron is responsible for developing one grain, then one can correlate the 
electron fluence (?) passing perpendicular to the film, to the optical density (Attix, 1986). 
If N is the number of silver bromide grains per unit area of the unexposed film, then n 
and OD can be written as: 
 
20.4343
 n N
 OD N
 ? ?
 ? ?
 =
 =
  
 
While the assumption is simplistic, it provides useful insights in OD film response. 
Because optical density is proportional to the number of silver grains per unit area and 
photon fluence, and because the photon or electron fluence is directly related to the 
radiation dose, the OD should be a function of dose. The relationship between dose and 
optical density is known as the sensitometric curve and is widely used in radiation 
dosimetry. Other characteristics of film can be plotted in various ways such as OD vs log 
(dose) known as the H & D curve (named after Hurter and Driffield). The H & D curve 
 
 14 
provides the characteristics of the film in three sections: toe, gradient and shoulder. In 
diagnostic radiology, this type of graph shows the optical density range that provides the 
optimal diagnostic information. Sensitometric curves are used in film dosimetry for 
calibration purposes. The film linearity is important for dosimetry because film should be 
exposed to doses within the linear region (Hale et al, 1994).  
 
The main challenges for using radiographic film as a megavoltage beam dosimeter is the 
dependence of OD on: (a) Photon energy, field size and depth in the phantom, (b) Film 
plane orientation with respect to the beam direction, (c)Emulsion differences amongst 
films of different batches, (d) Processing conditions (e)Method of densitometry and 
related artifacts (Pai et al, 2007).The implication being that the same optical density is 
not always associated with the same dose, making the conversion of OD to absolute dose 
difficult. In radiation oncology the dose versus OD presentation is used. The relationship 
between OD and dose depends strongly on the processing conditions, including developer 
temperature. OD increases as the developer temperature increases. Bogucki et al have 
shown that the optical density can be approximated as a function of temperature (Bogucki 
et al, 1997). Discrepancies in the magnitude of the effect of film orientation with respect 
to beam incidence have been reported. Burch et al observed no significant differences in 
optical densities due to film orientation (Burch et al, 1997) whilst on the other hand, 
Suchowerska et al showed that films exposed parallel to the beam axis had a measured 
over-response of about 15% at 25 cm depth in a phantom for a 10 cm x 10 cm field in a 6 
MV or 60Co beam (Suchowerska et al, 1999). Danclu et al found that the ratio of parallel 
to perpendicular response was almost unity except in the region around the dose 
maximum where the sensitivity for parallel response was about 4% lower for 60Co, 6 MV 
and 15 MV photon beams (Danclu et al, 2001). 
 
2.5 Thermoluminescent dosimetry  
 
Thermoluminescent dosimeters (TLD) are used in radiation therapy for dose 
measurement in total body irradiation, brachytherapy, stray radiation and verification of 
dose delivery (Yu and Luxton, 1999). TLDs are used to confirm or determine radiation 
 
 15 
dose to critical organs or to monitor special treatment. One of the commonly used TLD 
phosphors is TLD-100 (LiF), developed by the Harshaw Chemical Company. One merit 
of TLD dosimetry is that measurements can be carried out at different points placed 
inside solid phantoms simultaneously. With proper calibration, it is possible to obtain 
measured values with uncertainties within 3% (D?Angelo et al, 1999). The popularity of 
LiF-100 is due in part to its approximate tissue equivalence and low fading 
characteristics. A less desirable property however, is that it exhibits a non linear response 
with dose to certain types of radiation. The thermoluminescence (TL) output per unit 
dose is linear with dose up to 1 Gy, beyond which the response becomes supralinear 
(Folkard, 1989). TLDs of LiF in its purest form exhibit relatively little 
thermoluminescence, which is the phenomenon of photon emission subsequent to 
heating. The presence of impurities i.e., magnesium and titanium in TLD-100 LiF, 
appears to be necessary for radiation-induced TL. When a crystalline TLD is irradiated, a 
minute fraction of the absorbed energy is stored in the crystal lattice as shown in Figure 
2.2.  
 
 
Figure 2.2 A simplified energy diagram to illustrate the thermoluminescence process (Khan, 
2003). 
 
The energy stored in the irradiated TLD is recovered as visible light by placing it on a 
planchet heater in a commercially available TLD reader and heating the TLD in a light 
tight chamber. A photo multiplier tube (PMT) placed in the TLD reader detects the light 
emitted from the TLDs. The arrangement for measuring the TL output is shown in Figure 
2.3.  
 
 16 
 
Figure 2.3 Diagram showing apparatus for measuring thermoluminescence (Khan, 2003). 
 
The light output varies as the temperature is increased. The area under the glow curve, 
i.e. the total light output, is proportional to dose. Characteristics of TLD that are 
particularly important for radiation dosimetry include variability of the TL response with 
readout temperature, referred to as the glow curve. The variability of TL response relates 
to annealing procedures and the nonlinearity of the TL response. Zimmerman et al found 
that several prominent peaks in the TLD-100 glow curve decay at room temperature 
(Zimmerman et al, 1966). The peaks with longer half life are the most stable, and 
potentially the most suitable for radiation dosimetry. The peaks with shorter half life 
cause undesirable signals, which can be removed by various techniques of pre- and post- 
irradiation annealing. The presence of nitrogen gas in the planchet chamber has been 
observed to affect the precision of TLD reading, especially for radiation doses below 10 
cGy. Meigooni et al found that for radiation doses below 5 cGy, there was a large 
standard deviation in TL response (almost 100% of the signal) when TLDs were read 
without nitrogen gas flow in the TLD reader, whereas the standard deviation dropped to 
5% with nitrogen gas flow (Meigooni et al, 1995).  
 
Unfortunately, TLDs have low reproducibility and high batch to batch variability. In 
order to achieve the highest accuracy, the ?most stable? TLDs must be selected and 
calibrated identically. The calibration includes irradiation and dose reading, followed by 
annealing which involves heating to a certain temperature (e.g. 400oC) for an hour, 
cooling to room temperature and rechecking the sensitivity on the next day (Ertl et al, 
 
 17 
1996). Using an array of TLDs inserted into a phantom can provide point dose 
information that can be reconstructed into a two dimensional dose distribution. The 
necessary corrections can be established by plotting the TLD reading against dose. A 
useful empirical formula to correct for the supralinearity is 
 
( ) 12supralin 1 0.0278 0.000265C D D ?= + ?  
 
Where D is the dose given (not the TLD reading) and Csupralin is the factor by which the 
reading must be multiplied in order to correct for supralinearity. This formula is 
satisifactory up to 20 Gy. The supralinearity correction is not altered as the chips age 
(Mayles, 1993). The angular dependence of TLD is only a problem in so far as the 
orientation of the chips can affect the size of the cavity. This is more important for 
electrons and brachytherapy where the dose gradient may be high. It is however sensible 
to irradiate normal to the flat face of the chips. 
 
2.6 Computed Tomography  
 
CT images are produced with a highly filtered, high-kV x-ray beam of average energy 
approximately 75 keV. Allan Cormack realized that it was possible to observe radiation 
transmission profiles of an object and use that data to compute an image of that object 
(Cormack, 1963). The transmission of x-rays through an object is given by 
 
x
 oI I e
 ??
 =
  
 
Where I is the transmitted intensity, Io is the initial intensity, x is the thickness of the 
uniform object, and ?  is the linear attenuation coefficient. For non homogeneous media 
like the human body, the path along which the ray travels can be divided into a series of i 
elements (pixels) each with length (xi) and attenuation coefficient (? i) such that the x-ray 
transmission through the series of elements is then given by 
 
 
 18 
1 1( ....... )n nx x
 oI I e
 ? ?? +
 =
  
 
By moving the x-ray tube around the patient multiple transmissions can be obtained from 
different directions, resulting in many attenuation equations, which can be solved for ? i. 
These attenuation coefficients then expressed as CT numbers, 
 
[ ] 1000 ( )i water waterCT numbers ? ? ?= ?  
 
Where water? is the linear attenuation coefficient for water. Water and air yield CT 
numbers of 0 and -1000 respectively. On an image, each CT number is assigned to a 
shade of grey for display and the entire matrix of CT numbers results in a cross sectional 
image of the object. 
 
CT is used as an imaging device for diagnostic purposes. It provides a full (3-D) 
representation of the patient through a series of contiguous two dimensional (2-D) 
transaxial images. CT images consist of matrices of attenuation coefficients which are 
commonly used to construct a patient model in radiotherapy treatment planning. Part of 
the process is the conversion of CT numbers into electron density relative to water, for 
use in dose calculation. Routine measurement of CT number (and electron density) 
constancy is therefore a recommended part of the quality assurance programme in 
radiotherapy (Fraass et al, 1998). Accurate dose calculation not only requires accurate 
algorithm, but also accurate calibration of Hounsfield Unit (HU) for CT-based 
inhomogeneity corrections prior to dose calculation. Establishing a relationship between 
CT numbers and electron densities provides a simple method of correcting for 
inhomogeneous tissues. The possible errors in measured CT numbers need to be assessed 
in order to estimate the errors in calculated dose when the CT information is used directly 
(Papanikolaou et al, 2004). For instance, the CT number and thickness of bone cannot be 
measured accurately unless the thickness exceeds 3 mm. Below 3 mm there is a 
progressive underestimation of the CT number and an overestimation of thickness 
(Newman et al, 1998). 
 
 19 
2.7 The radiotherapy procedure 
 
 
Planning and delivering radiotherapy is a complex process which, for each individual 
patient, involves taking into consideration the definition of the target volume, decisions 
on total dosage, the fractionation schedule and the choice of appropriate radiation 
modality and method (Moller et al, 2003). There have been numerous steps in a 
radiotherapy process, which potentially causes a geometric discrepancy in patient/organ 
position between treatment planning and actual treatment delivery. In treatment 
simulation and planning, starting from patient positioning and ending at transferring the 
treatment plan to the radiotherapy machine, each step will introduce uncertainties in 
patient/organ geometry. On the other hand, most of the positioning steps will be repeated 
during the treatment delivery introducing the systematic and random variations in the 
patient/organ treatment geometry. The radiotherapy procedure includes: (a) Clinical 
investigation and type of treatment, (b) Radiotherapy Treatment Planning/ dose planning 
(c) Simulation (d) Irradiation (e) Follow up (f) Documentation and analysis 
 
2.7.1 Beam combinations  
 
Depending on the location of the tumor, several techniques and beam energies can be 
employed. The most common conventional techniques include a single beam, parallel 
opposed beams, a four field ?box?, rotational techniques, coplanar and non coplanar 
beams. 
 
2.7.2 Single photon beam 
 
Single photon beams are of limited use in the treatment of deep seated tumors beyond the 
depth of dose maximum. Single fields are often used for palliative treatments or for 
relatively superficial lesions (depth < 5-10 cm, depending on the beam energy). For 
deeper lesions, a combination of two or more photons beams is usually required to 
concentrate the dose in the target volume and spare the tissues surrounding the target as 
much as possible. 
 
 20 
2.7.3 Parallel opposed beams 
 
Parallel opposed beams overcome the difficulty of a decreasing dose gradient due to each 
individual beam. A decrease in the depth dose of one beam is partially compensated by 
an increased contribution from the opposing beam resulting in a uniform distribution 
within the tumor volume. 
 
2.7.4 Four field ?box? 
 
This is a technique of two opposing pairs of parallel opposed beams at right angles to 
each other, producing a relatively high dose ?box? shaped region. The region of highest 
dose occurs in the central portion of the volume that is irradiated by all four fields. This 
technique is commonly used for treatment of central organs in the pelvis.  
 
2.7.5 Treatment verification 
 
Treatment delivery encompasses a multi-faceted process to ensure that treatment delivery 
is implemented in accordance with the planned intention. Quality control helps to ensure 
that radiotherapy is administered safely and accurately and facilitates assessment of the 
reproducibility of treatment setups. Following localization and planning procedures, a 
simulator may be utilized to verify the treatment plan prior to its implementation on a 
treatment unit (e.g. verifying planned treatment fields, gantry angles, etc). This will 
confirm the appropriate coverage or avoidance of normal tissue structures. Various 
methods, materials and devices are conventionally used to try and reproduce patient 
position. These include sagittal and lateral lasers together with patient anatomical 
references e.g. supra sternal notch. Other aids may include immobilization devices such 
as breast boards, head and neck immobilization devices, etc. Installation and use of 
computerized record and verify systems (CRVS) help prevent treatment errors prior to 
treatment delivery. Data entered in the system before treatment delivery is available as a 
reference throughout the treatment course. CRVS verify a wide range of treatment 
 
 21 
parameters, including: (a) Patient name and identity number verification, (b) Mode of 
treatment, (c) Treatment field, (d) Cumulative doses, (e) Fraction number. 
 
Dose verification is as crucial as verifying the field placement position as it is the most 
obvious method of assessing the accuracy of patient?s treatment. Radiation detectors such 
as TLDs, semiconductor devices like silicon diodes and portal films are routinely 
employed in verification dosimetry (Cherry and Duxbury, 1998). Cylindrical phantoms of 
the appropriate radii are useful for simulating body and limb cross sections. However 
these homogeneous phantoms with regular shapes are not a good representation of a real 
patient and can give a misleading impression of the accuracy of computer calculations. 
The anthropomorphic phantom used in this work is affordable as compared to the 
commercially available humanoid phantoms. In certain clinical situations, in-vivo 
measurements can give unexpected results, even when the uncertainties in the 
measurements have been accounted for. Treatment planning algorithms should be 
verified by designing appropriate phantoms to test potential weaknesses. In-vivo 
measurements can be made but, because they introduce additional uncertainties, surface 
doses are not a useful test of dose calculation algorithms. This work presents a method of 
assessing the dose directly in a patient-like phantom as clinically this is restricted to 
placing TLDs on the surface of a patient. 
 
 
 
 
 
 
 
 
 
 
 
 
 
 22 
3.0 METHODOLOGY 
 
3.1 Phantom fabrication 
 
The pelvic anthropomorphic phantom that was used was fabricated at the Pretoria 
Academic Hospital, Department of Radiation Oncology (Kanduza, 2005). Perspex was 
used for the soft tissue, wax for the tumor volume and rectum, and plaster of paris for the 
left and right femurs of the pelvis based on mass attenuation coefficients, electron density 
and effective atomic number. External and internal contours of organs were fabricated 
based on the anthropomorphic phantom dimensions (Kanduza, 2005). The central part of 
the phantom accommodated circular inserts shown in Figures 3.1 and 3.2. Insert 1 was 
designed as the diagnostic anatomical insert and insert 2, the dosimetric verification 
insert. Insert 1 was used only for field design and placement. Insert 2 was designed to 
accommodate both TLDs and therapy verification film. These inserts were easy to 
exchange and could be repositioned in the phantom precisely. Figure 3.3 shows the 
position of film relative to the TLD positions in insert 2. 
 
 
 
 
 
Figure 3.1 Diagnostic Imaging insert 1 (Kanduza, 2005). 
 
 
 
 23 
   
 
 
Figure 3.2 Dosimetric insert 2 (Kanduza, 2005). 
 
 
 
 
 
 
Figure 3.3 Positions of film and TLDs in the Dosimetric insert (Kanduza, 2005). 
 
 
 
 
 
 
TLD 
positions 
 
 24 
3.2 TLD Procedure  
 
The Harshaw TLD System (Model 3500 Manual TLD Reader) was used for calibrating 
and reading the exposed TLDs. The Element Correction Coefficient (ECC) was used to 
compensate for the variations in sensitivity between dosimeters. This factor is a 
dimensionless parameter and it served to correct the sensitivity of each dosimeter in a 
batch to the average sensitivity of the batch. The TLDs were placed in a diagonal fashion, 
at dose points from the top left TLD position to bottom right and at the two opposite 
corners i.e bottom left and top right positions of the mid-axial insert. The same dose 
points were repeated throughout this experiment. Not all the TLD positions were used 
during the irradiations, which resulted in small air gaps. 
 
3.3 Calibration and irradiation process 
 
Lithium fluoride thermoluminescence dosimetry chips were used (TLD_100, Harshaw). 
A total of 40 TLDs were prepared for this work. All TLDs were handled with vacuum 
tweezers (to avoid scratches and contamination) and kept in dedicated trays. The 
radiation sources were 6 MV and 15 MV x-ray beam from a Siemens Primus M medical 
linear accelerator (Linac). The linac was calibrated isocentrically according to the TRS 
398 dosimetry protocol (IAEA TRS 398, 2001) to an isocentric output of 1.00 cGy/MU at 
the depth of maximum dose, in a water phantom, in the reference field size of 10 x 10 
cm2. A Poly-Methyl-Methacrylate (PMMA) phantom was used to irradiate a batch of 
TLDs at 5 cm water-equivalent depth. After exposure to 1 Gy, the TLDs were read to 
generate their respective ECCs. The TLD software was set to reject TLDs that had 
sensitivities of more than ?20% of the mean response. All 40 TLDs met this selection 
criterion. After readout, all TLDs were kept in a tray with identification for sorting 
purposes. A reader calibration factor (RCF) within the TLD reader software was used to 
convert the measured signal into dosimetric units.  
 
 
 
 
 25 
Direct measurement for treatment verification was therefore given by: 
 
( )    (  ) * ( /  ) *Dose Gy reader output reader units RCF Gy reader unit ECC=  
 
The percent deviation of all measurements was defined as: 
 
( ) % 1 *100
  
calculated doseDev Measured dose
 
 
 = ?   
 
where the calculated dose was derived from the planning system and the measured dose 
was the dose obtained from the phantom under radiation. 
 
For the phantom irradiations, 35 TLDs were placed in insert 2 at the points of interest, 
inserted into the Anthropomorphic phantom and exposed to 70 cGy delivered using either 
the AP/PA or the four field ?box? technique at 6 MV or 15 MV photon beam energy. 
Four TLDs were used as controls to monitor the TLD sensitivity following each 
annealing cycle. These TLDs were selected randomly from the batch of 40 TLDs and 
exposed to known doses of 0.5 Gy, 0.7 Gy, 1.0 Gy and 1.5 Gy before irradiation of the 
phantom. One TLD was not used. The variation of TLD output throughout the 
experiment was plotted as a function of the dose as shown in Figures 3.4 and 3.5 at 6 MV 
and 15 MV respectively. This provided a means of monitoring the sensitivity of the entire 
batch of TLDs after each annealing cycle and assessing the overall uncertainty. A typical 
cancer of the cervix protocol (used at Johannesburg Hospital) and a similar protocol 
using a 7 cm x 7 cm field size were used to develop the treatment plans (AP/PA and four 
field box). 
 
 26 
0.00
 0.20
 0.40
 0.60
 0.80
 1.00
 1.20
 1.40
 1.60
 1.80
 0.00 0.20 0.40 0.60 0.80 1.00 1.20 1.40 1.60 1.80
 Prescribed dose ,Gy
 TL
 D 
o
 u
 tp
 u
 t (G
 y)
   
Figure 3.4 Variation of TLD output (Gy) with dose (Gy) at 6 MV. The error bars represent the 
overall deviations in the measurements. 
 
 27 
0.00
 0.20
 0.40
 0.60
 0.80
 1.00
 1.20
 1.40
 1.60
 1.80
 0.00 0.20 0.40 0.60 0.80 1.00 1.20 1.40 1.60 1.80
 Prescribed dose, Gy
 TL
 D
  
o
 u
 tp
 u
 t, 
Gy
   
Figure 3.5 Variation of TLD output (Gy) with dose (Gy) at 15 MV. The error bars represent the 
overall deviation in the measurements. 
 
3.4 Film calibration 
 
The densitometer used had a calibration certificate traceable to the National Institute of 
Standards and Technology (NIST). Before use, the densitometer was calibrated using the 
calibration film that was provided with the densitometer. All films were irradiated, 
processed and measured at the Johannesburg Hospital Radiation Oncology Division for 
calibration purposes, the film was placed at the isocenter sandwiched in between PMMA 
sheets using the perpendicular geometry. The set up was similar to that used for 
calibrating the TLDs. For each of the photon beams, the films were irradiated to different 
 
 28 
doses covering the range from 0.1 Gy to 0.8 Gy and a sensitometric curve produced as 
shown in Figure 3.6.  
0
 0.5
 1
 1.5
 2
 2.5
 0 20 40 60 80 100
 Dose (cGy)
 N
 et
  
O
 pt
 ic
 al
  
de
 n
 si
 ty
 15 MV Photons
 6 MV Photons
  
 
Figure 3.6 Sensitometric curves at 6 MV and 15 MV for the perpendicular setup.. The 
uncertainties were very small and hence error bars are not visible. 
 
The films were irradiated over a single day, and at the end of the irradiation period, all 
films including one which was not irradiated, were processed under the same conditions 
using an automatic film processor operating at 34oC. The time between film exposure and 
processing was approximately 1 hour. It has been shown that a decrease in optical density 
of about 3% (for a dose of 0.45 Gy) and 1.8% (for a dose of 1.5 Gy) would be observed 
per month of delay between film irradiation and processing (Novotny et al, 1997). The 
film was used to obtain the 2D dose distributions. The films were exposed in their 
envelopes. For the films to fit into the phantom inserts, the envelopes were cut and made 
light tight using black insulation tape. A light proof test was conducted. The nuts on the 
phantom rods were then tightened to compress the film and to expel as much air as 
 
 29 
possible from the film envelopes. The isodoses recorded on film were generated at the 
Pretoria Academic Hospital using the Vidar RIT Dosimetry Pro film scanning system. 
 
3.5 Dose calculation on Helax-TMS  
 
The phantom was immobilized in a polyurethane, air equivalent cast and CT scanned in 
order to generate three dimensional images and to avoid treatment position uncertainties. 
Reference markers were used to enable repositioning during treatment delivery. A CT 
scan incorporating these markers was acquired with a 0.5 cm slice separation and this 
data set was transferred to Helax-TMS RTPS. The external contour (skin) and organs at 
risk were outlined. The phantom was then planned with 6 MV and 15 MV beams using 
the standard local technique for cancer of the cervix and a smaller field size of 7 cm x 7 
cm in terms of the isocentric placement and field borders. Dose calculations were carried 
out with both the PB and CC algorithms, both with and without inhomogeneity correction 
using the local cancer of the cervix protocol. The option of no inhomogeneity correction 
simply assigned all tissues in the patient geometry to unit mass density. This was done in 
order to compare the influence of tissue density on the dose calculations and to enable 
comparison to a manual 2D system of dose calculation. The isocentric planning technique 
was employed in this study. The plan normalization/prescription point was set to the 
isocenter. The treatment protocols obtained from the planning system are shown in Table 
3.1. The dose distributions in the central plane normalized to 100% at the isocentre are 
shown in Figures 3.7 to 3.16. The PB algorithm was used to obtain point doses because in 
this mode it was possible to obtain point doses at TLD positions which was not possible 
with the CC algorithm. 
 
 
 
 
 
 
 
 30 
Table 3.1 Treatment protocol showing the number of MUs needed to deliver 0.7 Gy to the 
isocenter for no inhomogeneity, manual calculation PB algorithm and CC algorithm. 
 
       
              
Technique             4 Field Box                 AP /PA 
             
Beam Number 1 2 3 4 1 2 
         
Beam Label Ant  Right Post Left Ant Post 
         
Beam Quality 15/6 15/6 15/6 15/6 15/6 15/6 
         
Radiation Type Photons Photons Photons Photons Photons Photons 
         
Gantry angle 0 270 180 90 0 180 
         
Treatment technique Iso Iso Iso Iso Iso Iso 
         
SSD (cm) 90.3 81.9 90.7 82.3 90.3 90.7 
         
Collimator Type Symmetric Asymmetric Symmetric Asymmetric Symmetric Symmetric 
         
Beam Width    [x] cm 16 12 16 12 16 16 
         
Beam Length   [y] cm 16 16 16 16 16 16 
         
          
No Inhomogeneities 
MUs to deliver         
 0.7 Gy to the         
isocenter (15MV/6 MV) 24.8/28.4 16.5/18.9 24.8/28.4 16.5/18.9 37.5/41.3 37.5/41.3 
          
          
PB MU's to deliver 
0.7Gy to the         
  isocenter (15MV/6 MV) 24.8/28.4 16.5/18.9 24.8/28.4 16.5/18.9 37.5/41.3 37.5/41.3 
          
          
CC MU's to deliver      
0.7 Gy to the          
 isocenter (15MV/6 MV) 25.2/28.7 16.8/19.1 25.2/29.7 16.8/19.1 37.9/41.8 37.9/41.8 
              
Manual MU calculations       
to deliver 0.7 Gy to the 
isocenter (15MV/6MV) 24.8/28.4 16.5/18.9 24.8/28.4 16.5/18.9 37.5/41.3 37.5/41.3 
 
 
 
 31 
 
Figure 3.7 Pencil Beam parallel opposed axial dose distribution at the level of the isocenter using 
6 MV beams. 
 
 
 
Figure 3.8 Collapsed Cone parallel opposed axial dose distribution at the level of the isocenter 
using 6 MV beams. 
 
 
 
 32 
 
Figure 3.9 Pencil Beam parallel opposed axial dose distribution at the level of the isocenter  using 
15 MV beams. 
  
 
 
Figure 3.10 Collapsed Cone parallel opposed axial dose distribution at the level of the isocenter 
using 15 MV beams. 
 
 
 
 33 
 
Figure 3.11 Pencil Beam four field ?box? axial dose distribution at the level of the isocenter and 
the position of the cut film using 6 MV beams. 
 
 
 
 
Figure 3.12 Collapsed Cone four field ?box? axial dose distribution at the level of the isocenter 
using 6 MV beams. 
 
 
 
 
Film shape and 
position 
 
 34 
 
 
 
Figure 3.13 Pencil Beam four field ?box? axial dose distributions at the level of the isocenter 
using 6 MV beams. 
 
 
 
Figure 3.14 Collapsed Cone four field ?box? axial dose distributions at the level of the isocenter 
using 15 MV beams. 
 
 
 35 
 
Figure 3.15 Pencil Beam parallel opposed axial dose distributions of a 7 cm x 7 cm field at the 
level of the isocenter using 6 MV beams.  
 
 
 
Figure 3.16 Pencil Beam four field ?box? axial dose distributions (7 cm x 7 cm box) at the level 
of the isocenter using 6 MV beams. 
 
 
 
 
 
 
 36 
4.0 RESULTS 
 
The TLD results are given in appendix A. They show results of both the local cancer of 
the cervix protocol and when a smaller field size of 7 cm x 7 cm was employed. Figures 
4.1 to 4.4 show the comparison between the doses measured by the TLDs using the 
AP/PA and the four field ?box? technique to the doses calculated by the planning system 
at both 6 MV and 15 MV. On average, the measured doses were lower than those 
calculated by the planning system as is indicated by the line of agreement. Figures 4.2 
and 4.4 show the different dose gradients that emerged from using the 4 field technique 
indicating regions of low, intermediate and high dose gradient. 
 
 
0.40
 0.50
 0.60
 0.70
 0.80
 0.90
 1.00
 1.10
 0.40 0.45 0.50 0.55 0.60 0.65 0.70 0.75 0.80
 Calculated dose, Gy
 M
 e
 a
 s
 u
 re
 d 
do
 s
 e
 , 
Gy
  
Figure 4.1 The comparison between the dose measured by the TLD?s using the AP/PA technique 
to the dose calculated by the PB algorithm at 6MV.  
 
 
 37 
0.40
 0.50
 0.60
 0.70
 0.80
 0.90
 1.00
 1.10
 0.40 0.50 0.60 0.70 0.80
 Calculated dose, Gy
 M
 ea
 su
 re
 d 
do
 se
 , 
G
 y
  
Figure 4.2 A comparison between the dose measured by the TLD?s and the dose calculated by the 
PB algorithm using the four field ?box? technique at 6MV.  Three dose regions dose emerged, 
those of low, intermediate and high dose gradient. 
0.40
 0.50
 0.60
 0.70
 0.80
 0.90
 1.00
 1.10
 0.40 0.50 0.60 0.70 0.80
 Calculated dose, Gy
 M
 ea
 su
 re
 d 
do
 se
 , 
G
 y
  
Figure 4.3 A comparison of measured dose with that calculated by the planning system for the 
AP/PA technique at 15 MV using TLDs. The PB algorithm was used to calculate the doses. 
 
 
 38 
0.40
 0.50
 0.60
 0.70
 0.80
 0.90
 1.00
 1.10
 0.40 0.50 0.60 0.70 0.80
 Calculated dose, Gy
 M
 ea
 su
 re
 d 
do
 se
 , 
G
 y
  
Figure 4.4 A comparison between the dose measured by the TLD?s and the dose calculated by the 
PB algorithm using the four field ?box? technique at 15MV. Three dose regions emerged, 
those of low, intermediate and high dose gradient. 
 
The ratios of the dose measured by the TLDs to the calculated dose (when the local 
cancer of the cervix protocol was followed) are shown in Figures 4.5 - 4.8. The mean 
value of the difference between the calculated and measured doses for the AP/PA at 6 
MV and 15 MV was 0.94?0.02 and that of the four field technique at 6 MV and 15 MV 
was 0.91?0.05. Three dose regions emerged for the four field technique, those of low, 
intermediate and high dose gradient, figure 4.2 and 4.4. The four field technique for 6 
MV gave a dose deviation from -4.9% to -32.1% and at 15 MV from -1.5% to -20.6%. 
For 6 MV and 15 MV parallel opposed beams, percentage dose deviations from -2.7% to 
-11.8% and from  +0.2% to -11.5% were observed respectively. Tables A.1 to A.6 in the 
appendix give the individual TLD results for the techniques used in this study and also 
show the difference between the calculated and measured doses at 6 MV and 15 MV. 
Tables A.5 and A.6 also show optical density results measured at TLD positions and their 
corresponding doses when a smaller field size (7 cm x 7 cm) is used. Both the TLD and 
film dosimetry results obtained were lower than that predicted by the planning system. 
The results for the local cancer of the cervix protocol and a 7 cm x 7 cm field size 
placement are shown in appendix A, figures A1 and A2. The results for film (OD) when a 
 
 39 
local protocol was followed were not included in the results. Film was over exposed due 
to the large field size resulting in high OD values and dose points that could not give any 
meaningful OD readings. 
0.40
 0.50
 0.60
 0.70
 0.80
 0.90
 1.00
 1.10
 0.40 0.50 0.60 0.70 0.80
 Calculated dose, Gy
 Dm/Dc
  
Figure 4.5 Ratio of measured to the calculated doses vs calculated doses for the AP/PA at 6 MV 
using TLDs. The PB algorithm was used to calculate the doses. 
0.40
 0.50
 0.60
 0.70
 0.80
 0.90
 1.00
 1.10
 0.40 0.50 0.60 0.70 0.80
 Calculated dose, Gy
 Dm/Dc
  
Figure 4.6 Ratio of measured to the calculated doses vs calculated doses for four field ?box? at 6 
MV using TLDs. The PB algorithm was used to calculate doses. 
 
 40 
0.40
 0.50
 0.60
 0.70
 0.80
 0.90
 1.00
 1.10
 0.40 0.45 0.50 0.55 0.60 0.65 0.70 0.75
 Calculated dose, Gy
 Dm/Dc
  
Figure 4.7 Ratio of measured to the calculated doses vs calculated doses for AP/PA at 15 MV 
using TLDs. The PB algorithm was used to calculate doses. 
 
0.40
 0.50
 0.60
 0.70
 0.80
 0.90
 1.00
 1.10
 0.40 0.50 0.60 0.70 0.80
 Calculated dose, Gy
 Dm/Dc
  
Figure 4.8 Ratio of measured to the calculated doses vs calculated doses for four field ?box? at 15 
MV using TLDs. The PB Algorithm was used to calculate point doses 
 
 
 
 
 41 
5.0 DISCUSSION 
 
The treatment planning process consists of various steps, including the calculation of the 
relative isodose distribution and the treatment time for each individual field. A manual 
check of a complex dose distribution in inhomogeneous volumes is usually not 
considered to have the same accuracy as a RTPS calculation, since computer algorithms 
are sophisticated and able to handle the physics of the radiation interactions more 
accurately. Nisbet et al observed that for the same input data set, the CC algorithm with 
Helax-TMS version 6.1 improved the agreement between calculated and measured dose 
in the build up compared with version 5.1A, where significant discrepancies were 
present. The reason for the differences between the versions was not stated and this 
confirms that upgrades of commercial software should be checked prior to clinical use 
(Nisbet et al, 2004).  
 
This work attempted to validate the entire treatment chain using a pelvic 
anthropomorphic phantom, fabricated using locally available materials as human tissue 
substitutes. The planning system used had two dose calculation algorithms, the PB and 
the CC. Both algorithms were used for clinical treatment planning. The plans showed that 
there was little or no clinically significant difference in the calculation of dose by the two 
algorithms however, no large non-unit density heterogeneity scenarios were studied. A 
limitation of the RTPS was the lack of interaction when the CC algorithm was used, for 
this reason, point doses were only obtained for the PB algorithm. 
 
Thermoluminescent dosimetry is a relative dosimetry system and its accuracy depends on 
various parameters, such as the reproducibility of the TL reading, calibration techniques 
and efforts put into the measurement process. Most of the TLD results however were 
below the 10% dose deviation. For the 6 MV AP/PA field arrangement, 32 TLDs out of 
the total 35 TLDs exposed (91.4%) were below the 10% dose deviation, 40% (14/35) for 
the 6 MV four field arrangement, 91.4% (32/35) for the AP/PA field arrangement at 15 
MV and 57.1% (20/35) for the four field arrangement at 15 MV. This result indicated that 
the TLD results were better at 15 MV than at 6 MV. The readings were reproducible 
 
 42 
(standard deviations of less than 3) as shown in Tables A.1 ? A.6. The absorbed dose rate 
of the treatment unit was confirmed with a calibrated ionization chamber before each of 
the TLD irradiations. The main source of uncertainty in this experiment could have been 
due to loss of signal due to fading caused by the heating cycle and TLD sensitivity to 
light. For this reason detectors should be shielded from light during use by, for example, 
a black polyethylene wrapper which can also serve to keep the detectors clean. Another 
source of uncertainties could have been due to variation in the treatment unit output as 
shown in figures 3.4 and 3.5. The supralinearity correction was not applied to the TLD 
readings as the point doses of interest in this study were within the linear region of TLD 
sensitivity. Three dose regions emerged for the four field ?box? technique, those of low, 
intermediate and high dose gradient. TLDs placed in low dose gradients showed better 
agreement than those placed in high dose gradient regions. This was also observed by 
Kanduza (Kanduza, 2005).  
 
The film calibration results obtained in this study gave sensitometric curves at 6 MV and 
15 MV that almost overlapped, indicating energy independence of the film. The OD 
measured on the film after exposure was higher than doses predicted by the planning 
system and those measured at TLD dose points when a local cancer of the cervix protocol 
was used. However, when a smaller field size was employed, the OD was comparable to 
the TLD readings suggesting that the higher readings were mainly due to large field size 
placement. Over response of film is generally attributed to the softening of the photon 
spectra with depth in the phantom and also to the fact that the OD depends on many 
factors (e.g. photon energy, film plane orientation and processing conditions) the 
implication being that OD is not always associated with the same dose, making the 
conversion of OD to absolute dose a very difficult task. The conical design of the 
dosimetry insert restricts the size of film that can be used from 8 cm x 8 cm to 14 cm x 14 
cm.  The 70% isodose was therefore the minimum level visible on the largest films when 
the local cancer of the cervix protocol is followed (appendix A, figure A.1). The 2D 
isodose distributions obtained were therefore not meaningfully comparable to the ones 
predicted by the planning system or to the 2D isodoses obtained by Kanduza (Kanduza, 
2005). Kanduza used field sizes of less than 12 x 12 cm for the dosimetry verification, 
 
 43 
however this is not typical of cancer of the cervix field size placement. When a smaller 
field size was used (figure 3.15 and 3.16), the film results obtained were comparable to 
the TLD results (Table A.5 ? A.6) even though both results were lower compared to 
doses predicted by the RTPS. Kanduza?s methodology included the four field ?box?, 
AP/PA and arc therapy techniques at 6 MV and 10 MV. The verification between 
measured and calculated dose using TLDs in this study did not totally agree with that 
obtained by Kanduza. Comparisons of point doses could not be done as Kanduza did not 
indicate how the TLDs were positioned in the dosimetry insert. The results obtained by 
Kanduza using the AP/PA at 6 MV are shown in appendix B. 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 44 
CONCLUSION AND RECOMMENDATIONS 
 
The treatment planning system accurately predicted doses in the anthropomorphic 
phantom to within 3% of the prescribed dose at selected dose points as can be seen from 
Tables A.1 to A.6. However, the generally stated goal of dose delivery to an accuracy of 
within 5% was not met at most dose points, especially where the TLDs were placed in 
regions of high dose gradient.  
 
The percentage dose deviation between dose measured by the TLDs and that calculated 
by the Helax planning system, was lower at 15 MV than at 6 MV. 
 
? The TLD results at 6 MV photon beam energy obtained in this study did not 
totally agree with those obtained by Kanduza (Kanduza, 2005). The phantom was 
not found to be useful for 2D isodose distribution with film when the local cancer 
of the cervix protocol was employed. This was because the cut film dimensions 
were smaller than the set field size when this protocol was followed. 
Discrepancies between the results obtained in this study and those obtained by 
Kanduza with film dosimetry imply that the phantom is not suitable for film 
dosimetry of all beam arrangements. Nonetheless, the phantom could still be 
useful for dosimetry audits of techniques where the set field sizes are smaller 
than the film insert dimensions (e.g. 8 cm x 8 cm to 14 cm x 14 cm).  
 
 
? The size of the TLD holes should be machined to accurately fit the dimensions of 
square TLD_100 chips to minimise systematic errors due to gaps and also 
prevent areas of non equilibrium. Alternatively, the holes could be machined to 
fit sachets of TLD powder.  
 
? The sensitometric curves at 6 MV and 15 MV almost overlapped, indicating that 
the energy dependence of the film at these beam energies was minimal. 
 
? The phantom is limited to film dosimetry of small field sizes. 
 
 45 
 
? The phantom could be modified to incorporate bigger cylindrical inserts that 
would allow for more typical cancer of the cervix field size placement.  
 
? A similar phantom that is affordable could be developed to accommodate an 
entire sheet of film in order to eliminate errors introduced by cutting the film. 
 
? Provisions for reference ionization chamber measurements could serve as a means 
of accurately confirming the absolute dose. 
 
? The decision to use either parallel or perpendicular irradiation for film calibration 
depends on the degree of energy dependence of the film used and the geometry 
that will be used for actual test measurements. The geometry used for the 
calibration of film should be the same as that used for actual measurement in the 
phantom. For this phantom, calibration of film should be carried out parallel. 
 
? The Helax treatment planning system should be checked if it accurately corrects 
for electron density. The electron-density is the most important material property 
that should be taken into account by photon dose calculation algorithms and not 
the mass-density. The mass-density method tends to overestimate dose, relative 
to electron-density prediction.   
 
 
 
 
 
 
 
 
 
 
 
 46 
REFERENCES  
 
Ahnesjo, A. (1989) Collapsed Cone convolution of radiant energy for photon dose 
calculation in heterogeneous media. Med. Phys. 16 577-592. 
 
Ahnesjo, A. (1997) Cone discretization for the collapsed cone algorithm, in proceeding of 
the int. conf on the use of computers in radiation therapy  XII ICCR Salt lake City, USA, 
edited by D. D Leavitt and G. Starkschall Medical Physics Publishing , Madison, WI  
114-116. 
 
Ahnesjo, A. (1994) Analytic modeling of photon scatter from flattening filters in photon 
therapy beans. Med. Phys. 21 1227-1235. 
 
Aspradakis, M.M., Morrison, R.H., Richmond, N.D. and Steele, A. (2003) Experimental 
verification of convolution/superposition photon dose calculation for radiotherapy 
treatment planning. Phys Med Biol. 48 2873-2893. 
 
Aspradakis, M. M., McCallum, H. M. and Wilson, N. (2006) Dosimetric and treatment 
planning considerations for radiotherapy of the chest wall. Br J Radiol. 79 828-836. 
 
Attix, F. H. (1986) Introduction to Radiological Physics and Radiation Dosimetry, Wiley, 
New York. 
 
Bogucki,T. M., Murphy, W.R., Baker, C.W., Piaza, S.S and Haus, A.G. (1997) Processor 
quality control in laser imaging systems, Med. Phys. 24 581-584. 
 
Brahme, A. (1984) Dosimetric precision requirements in radiation therapy. Act Radiol 
Oncol 23 379-391. 
 
Burch, S.E., Yeo. I. J. and Wang, C.K. (1997) A new approach to film dosimetry for high 
energy photon beams: lateral scatter filtering. Med. Phys. 24 773-783. 
 
 47 
Cherry, P. and Duxbury, A. (1998) Practical Radiotherapy Physics and Equipment. 
Greenwich Medical Media. McGraw-Hill, New York, 2nd edition. 
 
Cormack, A.M. (1963) Representation of a function by its line integrals, with some 
radiological application. J. App. Physics 34 2722-2727. 
 
Danclu, C., Proimos, B.S., Rosenwald, J and Mijnheer, B.J. (2001) Variation of 
sensitometric curves of radiographic films in high energy photon beams. Med. Phys. 28 
(6) 966-974. 
 
D?Angelo, L., Furetta, C., Giancola, S., Lannoli, D and Scaccol, A. (1999) Verification of 
treatment planning systems for therapeutic radiations using thermoluminescence 
dosimeters, Rad. Prot. Dos. 85 (4) 401-404. 
 
Dunscombe, P., McGhee, P. and Lederer, E. (1996) Anthropomorphic phantom 
measurements for the validation of a treatment planning system. Phys. Med. Biol. 41 
399?411. 
 
Dutreix, A., Blarngard, B.E., Bridier, A., Mijnheer, B., Shaw, J.E. and Svenson, H. 
(1997) Monitor unit calculation for high energy photon beams, ESTRO, booklet 3. 
 
Ertl, A., Hart1, R.F., Zehetmayer, M., Kitz, K., Griffitt, W.  (1996)  TLD array for 
precise dose measurements in stereotactic radiation techniques Phys. Med. Biol. 41(12) 
2679-2686. 
 
Fano, U. (1954) Note on the Bragg-gray cavity principle for measuring energy dissipation 
Radiat. Res. 1 237-240. 
 
Folkard, M., Roper, M.J. and Michael,B.D. (1989) Measurement and analysis of 
supralinearity in LiF TLD- 100 irradiated by 1.5 keV x-rays. Phys. Med. Biol. 34(6), 707-
 715. 
 
 48 
Fraass, B., Doppke, K., Hunt, M., Kutcher, G., Starkschall, G., Stem, R. and Van Dyke, J. 
(1998) American Association of Physicists in Medicine Radiation Therapy Committee 
Task group 53: Quality assurance for clinical radiotherapy treatment planning Med. Phys. 
25 1773?1829. 
 
Hale, J.I., Kerr, T. and Shragge, P.C. (1994) Calibration of film for accurate megavoltage 
photon dosimetry, Med. Dosimetry 19 43-46. 
 
Helax-TMS 6.1B Manual, (2003), Dose formalism and models in helax-tms 
 
Herrick, S. J. and Newman, D. F. (1999) Verification of a photon beam algorithm in a 3D 
radiation therapy treatment planning system. Med. Dosimetry 24 (3) 179-182.  
 
ICRU Report 24, (1976) Determination of absorbed dose in a patient irradiated by beams 
of X or gamma rays, Washington DC. 
 
ICRU Report 42, (1987) Use of computers in external radiotherapy procedures with high 
energy photons and electron, Washington D.C. 
 
ICRU report 59, (1998) Clinical proton dosimetry part 1, beam production, beam delivery 
and measurement of absorbed dose. International commission on radiation and 
measurement . Bethusda M.D. 
 
International Atomic Energy Agency, (2001) Absorbed dose determination in external 
beam therapy: An international code of practice for dosimetry based on absorbed dose to 
water. IAEA Technical Reports Series 398, Vienna. 
 
Johns, H. E. and Cunningham, J. R. (1983) The physics of radiology, Fourth edition, 219-
 220, Thomas, Illinois. 
 
 
 49 
Kanduza, M. (2005) Design and evaluation of a dosimetry system to verify the 
radiotherapy procedure from treatment planning to treatment delivery, MSc thesis, 
Pretoria University. 
 
Karzmark, C.J. (1993) Medical Linear Accelerators, New York: McGraw Hill Inc 
 
Khan, F.M. (2003) The Physics of Radiation Therapy 3rd ed. Philadelphia USA 
Lippincott Williams and Wilkins. 
 
Knoos, T., Ceberg, C., Weber, L. and Nilson. P. (1994) ?The dosimetric verification of a 
pencil beam based treatment planning system. Phys. Med. Biol. 39 1609-1628. 
 
Knoos, T. Ceberg, C., Weber, L. and Nilson. P. (1995) Limitations of a pencil beam 
approach to photon dose calculation in lung tissue.  Phys. Med. Biol. 40 1411-1420. 
 
Leszczynski, K.W. and Dunscombe, P.B. (2000) Independent corroboration of monitor 
unit calculations performed by 3D computerized planning system. J. App Clin Med Phys, 
1, (4) 234-242. 
 
Mayles, W. P.M., Heisig, S. and Mayles, H.M.O. (1993) Treatment verification and in 
vivo dosimetry. Radiotherapy Physics in Practice, Oxford University Press, Oxford. 
 
Meigooni, S. A., Mishra, V., Panth, H. and Williamson, J. (1995) Instrumentation and 
dosimeter size artifacts in quantitative thermoluminescence dosimetry of low dose fields, 
Med. Phys. 22 555-561. 
 
Metcalfe, P., Kron, T. and Hoban, P. (1997) The Physics of Radiotherapy X-Rays from 
Linear Accelerators. Madison, WI: Medical Physics Publishing.  
 
 
 50 
Mijnheer, B.J., Battermann, J.J. and Wambersie, A. (1987) What degree of accuracy is 
required and can be achieved in photon and neutron therapy? Radiother. Oncol. 8 237?
 252. 
 
Mihailidis, D.N., Alkhatib,H., Gibbons, J.P. (2000). Use of primary off-axis ratios for 
dose and monitor unit calculation in wedged photon beams with asymmetric collimation. 
2 1088-1091. 
 
Moller, T.R., Ceberg, J., Eimham, N., Lindholm, C. and Nylen, U. (2003) Radiotherapy 
techniques in current use in Sweden. Act Oncologica  42, (5/6) 376-386. 
 
Newman, D.L, Dougherty, G.A.l., Obaid, A.  and Hajrasy, H.A.l. (1998) Limitations of 
CT in assessing cortical thickness and density, Phys. Med. Biol. 43 619?626. 
 
Nilsson, M. and Knoos,T. (1992) ?Application of the Fano theorem in inhomogeneous 
media using a convolution algorithm,? Phys. Med. Biol. 37 69?83. 
 
Nisbet, A., Beange, I., Vollmar, H., Irvine, C., Morgan, A and Thwaities D.I. (2004) 
Dosimetric verification of a commercial collapsed cone algorithm in simulated clinical 
situations. Radiotherapy and oncology 73 79-88 
 
Novotny, J., Gomola, I., Izewska, J. and Huyskens, D. (1997) External audit of photon 
beams by mailed film dosimetry: feasilibility study. Phys. Med. Biol. 42 1277-1288 
 
O?Connor, J.E. (1957) The variation of scattered x-rays with density in an irradiated body 
Phys. Med Biol 1 352-369 
 
Pai, S., Das, I.J., Dempsey, J.F., Lam, K.L., LoSasso, J.T., Olch, J.A., Palta, J.R., 
Reinstein, L.E., Ritt, D and Wilcox, E.E. (2007) Radiographic film for mega voltage 
beam dosimetry, Med Phys 34 (6) 2228-2258 
 
 
 51 
Papinikalaou, N., Battista, J.J., Boyer, A.L,. Kappas, C., Klein, E., Mackie, T.R., Sharpe, 
M. and Van Dyk, J (2004) Tissue inhomogeneity correction for megavoltage photon 
beams. Report of Task Group No 65 of the Radiation Therapy Committee of the 
American Association of Physicists in Medicine. Medical Physics Publishing. 
 
Starkschall, G., Steadham, R.E., Wells, N.H., O?Neil, L., Miller, L.A and Rosen, I.I. 
(2000) On the need for monitor unit calculations as part of a beam commissioning 
methodology for a radiation treatment planning system. J. Appl. Clin. Med.Phys. 1 86-94 
 
Suchowerska, N., Hoban, P., Davison, A. and Metcalfe, P. (1999) ?Perturbation of 
radiotherapy beams by radiographic film: measurements and Monte Carlo simulations,? 
Phys. Med. Biol. 44 1755?1765  
 
Swinnen, A. (2005) Quality assurance in radiotherapy: Development and validation of a 
mailed dosimetry procedure for external audits using a multipurpose phantom and in vivo 
dosimetry. PhD Thesis Katholieke University Leuven 
 
Weber, L. and Nilsson, P. (2002) Verification of dose calculations with a clinical 
treatment planning system based on a point kernel dose engine. J.Appl Clin Med Phys, 3 
73-87 
 
Yeo, I.J., Wang, C.K. and Bursch, S.E. (1999). A new approach to film dosimetry for 
high energy photon beams using organic plastic scintillators. Phys. Med. Biol. 44 3055-
 3069. 
 
Yu, C. and Luxton, G. (1999) TLD dose measurement: A simplified accurate technique 
for the dose range from 0.5 cGy to 1000 cGy. Med Phys. 26 (6) 1010-1016 
 
Zimmerman, D. W., Rhyner, C.R. and Cameron, J.R. (1966) Thermal annealing effects 
on the thermoluminescence of LiF, Health physics 12 525-531 
 
 
 52 
APPENDIX A 
 
Table A.1 Individual TLD results for 6 MV AP/PA compared to point doses predicted by the 
planning system when the local cancer of the cervix protocol is used. 
 
       
TLD Average  Expected dose  % Dose  
Identification measured dose  (cGy) deviation 
  (cGy)  (PB)   
        
A5 67.8 ? 0.4 73.5 8.5 
A6 68.6 ? 0.2 73.5 7.2 
A7 70.1 ? 1.3 72.1 2.9 
A8 68.4 ? 2.5 72.1 5.5 
A9 66.6 ? 0.4 70.7 6.2 
B1 62.6 ? 0.6 70.0 11.8 
B2 69.3 ? 0.4 73.5 6.1 
B3 68.5 ? 2.3 74.2 8.3 
B4 71.6 ? 2.5 73.5 2.7 
B5 66.7 ? 1.9 72.1 8.2 
B6 66.9 ? 0.2 72.1 7.9 
B7 65.9 ? 0.5 71.4 8.4 
B8 63.8 ? 1.8 70.7 10.8 
B9 69.5 ? 3.0 73.5 5.8 
C1 67.8 ? 0.4 74.2 9.4 
C2 69.1 ? 2.3 73.5 6.4 
C3 66.1 ? 0.8 72.1 9.1 
C4 65.8 ? 1.1 72.1 9.6 
C5 64.1 ? 0.1 71.4 11.4 
C6 65.4 ? 0.4 70.7 8.1 
C7 69.9 ? 0.1 74.2 6.2 
C8 72.0 ? 1.5 74.9 4.1 
C9 69.4 ? 1.2 72.8 5.0 
D1 67.6 ? 0.6 72.1 6.7 
D2 68.7 ? 0.5 72.1 5.0 
D3 68.5 ? 0.8 72.8 6.4 
D4 68.5 ? 0.7 71.4 4.2 
D5 69.7 ? 0.4 74.9 7.5 
D6 70.2 ? 1.1 74.2 5.8 
D7 67.3 ? 0.1 72.8 8.2 
D8 66.0 ? 1.0 70.7 7.1 
D9 66.7 ? 2.4 71.4 7.0 
E1 67.4 ? 1.8 72.1 7.0 
E2 65.6 ? 2.0 71.4 8.8 
E3 68.3 ? 0.2 74.2 8.7 
 
 
 
 53 
Table A.2 Individual TLD results for 6 MV Four Field ?box? compared to point doses predicted 
by the planning system when the local cancer of the cervix protocol is used. 
 
        
TLD  Average Expected dose % Dose 
Identification measured dose (cGy) deviation 
  (cGy)  (PB)   
        
A5 45.6 ? 1.3 53.2 16.9 
A6 46.4 ? 1.3 53.2 14.9 
A7 46.9 ? 0.7 53.9 14.9 
A8 47.7 ? 0.7 63 32.1 
A9 64.0 ?3 .5 72.8 13.8 
B1 60.8 ? 0.9 73.5 21.0 
B2 45.6 ? 0.3 53.2 16.7 
B3 48.0 ? 0.8 52.5 9.5 
B4 47.8 ? 0.5 53.2 11.4 
B5 47.9 ? 0.6 53.2 11.1 
B6 49.5 ? 1.1 61.6 24.6 
B7 63.5 ? 1.2 70.7 11.4 
B8 62.6 ? 0.4 70.7 13.0 
B9 45.9 ? 0.1 52.5 14.4 
C1 47.1 ? 0.4 53.2 13.0 
C2 46.0 ? 1.8 53.2 15.8 
C3 49.5 ? 0.4 53.2 7.5 
C4 48.5 ? 0.4 60.2 24.3 
C5 62.1 ? 0.4 70.7 13.9 
C6 63.6 ? 1.0 70.7 11.2 
C7 49.2 ? 0.9 52.5 6.8 
C8 49.9 ? 0.6 53.2 6.7 
C9 50.7 ? 1.3 53.2 4.9 
D1 49.7 ? 0.1 53.2 7.0 
D2 50.5 ? 0.1 61.6 22.1 
D3 66.8 ? 1.1 72.1 7.9 
D4 67.5 ? 0.8 71.4 5.8 
D5 50.0 ? 0.6 53.2 6.5 
D6 50.3 ? 0.1 53.9 7.2 
D7 48.9 ? 0.3 53.2 8.8 
D8 50.0 ? 0.1 61.6 23.2 
D9 65.4 ? 0.1 71.4 9.2 
E1 64.6 ? 2.7 71.4 10.5 
E2 67.3 ? 1.3 71.4 6.2 
E3 50.6 ? 0.4 53.9 6.6 
 
 
 
 
 54 
Table A.3 Individual TLD results for 15 MV AP/PA compared to point doses predicted by the 
planning system when the local cancer of the cervix protocol is used. 
  
 
TLD Average Expected Dose % Dose  
Identification measured dose (cGy) deviation 
  (cGy) (PB)   
      
A06 64.9 ? 0.4 71.54 9.3 
A07 63.8 ? 0.9 72.24 11.8 
A08 66.7 ? 0.3 71.82 7.1 
A09 65.5 ? 0.9 71.26 8.2 
B01 65.9 ? 0.1 70.21 6.1 
B02 65.9 ? 0.8 69.93 5.8 
B03 65.7 ? 0.5 71.54 8.2 
B04 67.6 ? 0.2 72.24 6.5 
B05 66.0 ? 0.4 72.59 9.1 
B06 65.1 ? 0.1 71.75 9.3 
B07 66.3 ? 0.6 71.82 7.8 
B08 65.5 ? 0.5 70.91 7.7 
B09 66.4 ? 1.1 70.56 6.0 
C01 66.8 ? 1.1 72.1 7.4 
C02 68.0 ? 1.2 72.59 6.4 
C03 66.7 ? 0.4 73.01 8.6 
C04 63.6 ? 1.1 71.82 11.5 
C05 66.7 ? 0.2 71.89 7.3 
C06 67.3 ? 1.1 71.4 5.7 
C07 68.0 ? 0.6 70.91 4.2 
C08 70.1 ? 1.1 72.59 3.4 
C09 69.5 ? 0.9 73.01 4.9 
D01 66.7 ? 1.7 72.31 7.8 
D02 66.3 ? 0.3 71.12 6.8 
D03 65.0 ? 0.1 71.68 9.4 
D04 70.2 ? 1.0 72.17 2.7 
D05 68.5 ? 0.8 71.54 4.3 
D06 68.4 ? 0.3 72.17 5.2 
D07 70.8 ? 1.8 72.94 2.9 
D08 65.9 ? 2.8 72.31 8.9 
D09 63.5 ? 2.1 70.7 10.3 
E01 71.2 ? 1.1 71.05 0.2 
E02 71.6 ? 1.2 72.03 0.7 
E03 72.9 ? 1.2 71.54 1.8 
E04 72.7 ? 0.4 73.01 0.4 
 
 
 
 
 55 
Table A.4 Individual TLD results for 15 MV four field ?box? compared to point doses predicted 
by the planning system when the local cancer of the cervix protocol is used. 
 
 
TLD  Average  Expected dose  
 
% Dose 
Identification measured dose (cGy) deviation 
   (cGy) (PB)   
      
A06 43.30 ? 2.8 49.21 12.0 
A07 43.85 ? 0.8 50.26 12.8 
A08 46.75 ? 1.5 51.10 8.5 
A09 48.85 ? 0.6 61.53 20.6 
B01 62.85 ? 0.1 71.68 12.3 
B02 65.80 ? 1.1 72.17 8.8 
B03 44.55 ? 0.1 49.70 10.4 
B04 44.50 ? 0.3 49.07 9.3 
B05 45.05 ? 0.4 50.19 10.2 
B06 45.70 ? 0.4 51.17 10.7 
B07 51.40 62.09 17.2 
B08 64.95 ? 0.9 70.21 7.5 
B09 67.60 ? 1.1 70.84 4.6 
C01 43.80 ? 0.3 49.07 10.7 
C02 44.50 ? 0.3 49.07 9.3 
C03 45.55 ? 0.6 50.19 9.2 
C04 45.40 ? 0.8 50.82 10.7 
F01 49.55 ? 1.2 59.99 17.4 
C06 65.35 ? 0.9 70.28 7.0 
C07 67.55 ? 0.1 70.56 4.3 
C08 46.95 ? 0.4 49.00 4.2 
C09 46.95 ? 0.4 49.91 5.9 
D01 44.40 ? 0.7 49.98 11.2 
D02 45.75 ? 0.1 50.47 9.4 
D03 48.75 ? 1.3 59.01 17.4 
D04 70.10 ? 1.4 71.19 1.5 
D05 67.75 ? 1.3 71.26 4.9 
D06 44.75 ? 0.8 49.00 8.7 
D07 48.10 ? 1.0 50.40 4.6 
D08 46.00 ? 1.0 51.38 10.5 
D09 47.90 ? 1.8 57.68 17.0 
E01 69.00 ? 0.4 70.28 1.8 
E02 69.70 ? 1.0 71.54 2.6 
E03 70.25 ? 0.2 71.40 1.6 
E04 49.15 ? 1.5 50.54 2.8 
 
 
 
 
 56 
Table A.5 Individual TLD and film results for 6 MV AP/PA treatment plan based on figure 4.9 
compared to point doses predicted by the planning system.  
 
 
      
  
  
TLD Average ) Expected  % Dose  Optical  Corresponding 
Identification (cGy) Dose Deviation Density Dose  on film 
    (cGy) (TLD)   (cGy) 
        
 
 
A20 63.3?0,2 74.5 17.8 2.57 64 
A50 68.5?0.3 74.3 8.5 2.64 67 
A60 64.8?0.1 74.2 14.5 2.69 68 
A70 68.5?0.7 73.5 7.3 2.59 66 
A80 64.8?0.1 73.5 13.5 2.69 68 
A90 63.9?0.3 73.5 15.0 2.72 70 
B10 68.5?0.4 72.5 5.6 2.59 66 
B20 62.7?0.8 72.7 16.1 2.71 70 
B30 63.2?0.9 73.5 16.3 2.72 70 
B40 61.7?1.4 71.7 16.3 2.58 66 
B50 65.3?1.3 72.0 10.2 2.7 69 
B70 64.5?0.3 71.9 11.5 2.76 71 
C20 62.6?0.2 71.1 13.7 2.61 67 
C30 65.7?0.7 71.5 8.9 2.69 69 
C50 64.5?1.3 71.9 11.5 2.76 71 
      
  
 
  
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 57 
Table A.6 Individual TLD and film results for 6 MV four field ?box? treatment plan based on 
figure 4.10 compared to point doses predicted by the planning system. 
 
      
  
  
TLD Average ) Expected  % Dose  Optical Corresponding  
Identification (cGy) Dose Deviation Density Dose (film) 
    (cGy) (TLD)    (cGy) 
        
 
  
D10 47.4?0.6 46.2 -2.5 2.03 46 
D20 41.0?2.8 46.2 12.7 2.02 46 
D30 41.2?0.3 46.2 12.1 1.93 42 
D40 43.7?0.8 46.2 5.8 2.02 44 
D60 42.8?0.6 46.2 7.9 2.02 44 
D70 48.1?0.2 46.2 -3.9 1.97 45 
D80 44.5?1.6 49.0 10.2 2.09 47 
E10 45.4?0.6 48.3 6.5 2.06 46 
E20 45.0?0.8 49.7 10.4 2.05 46 
E30 62.8?0.1 69.3 10.4 2.62 66 
E40 68.7?0.2 69.3 0.9 2.62 66 
E50 65.4?0.1 70.7 8.2 2.57 63 
E60 71.3?0.4 71.4 0.2 2.63 67 
E80 67.5?0.2 71.4 5.9 2.65 67 
F10 64.3 71.4 11.0 2.55 63 
      
  
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 58 
 
 
Figure A.1 2-D isodose distribution at 15 MV for the four field box technique using a 
local cancer of the cervix protocol. Film was cut and positioned as shown in figure 4.5 
 
 
 
 
 
 
 
 
 
 
 
 
 
Figure A.2 Shows an exposed film (using a 7 cm x 7 cm field size) at 6 MV for the four 
field box technique, showing the box shape. Film was cut and positioned in figure 4.10.  
 
 
 59 
APPENDIX B 
 
Table B.1This table shows Kanduza?s TLD results at 6 MV using the AP/PA technique.  
A field size of 10 cm x 10 cm was used (Kanduza, 2005). 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 60 
 
Calculated         Measured    % difference