IMPLEMENTATION OF IN VIVO DOSIMETRY FOR EXTERNAL PHOTON BEAM RADIOTHERAPY USING GAFCHROMIC EBT 2 FILM A THESIS PRESENTED TO THE:

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1 IMPLEMENTATION OF IN VIVO DOSIMETRY FOR EXTERNAL PHOTON BEAM RADIOTHERAPY USING GAFCHROMIC EBT 2 FILM A THESIS PRESENTED TO THE: DEPARTMENT OF MEDICAL PHYSICS, UNIVERSITY OF GHANA By DANIEL AKWEI ADDO; BSC (GHANA), 2004 IN PARTIAL FULFILLMENT OF THE REQUIREMENT FOR THE DEGREE OF MASTER OF PHILOSOPHY IN MEDICAL PHYSICS JUNE, 2014 i

2 DECLARATION I hereby declare that, except for references to other people s work, which have been duly cited, this thesis is the result of my own research work and that it has neither in part nor in whole been presented for any degree elsewhere. Daniel Akwei Addo (Student) Date:.. Prof. A.W.K. Kyere (Principal Supervisor) Date:. Dr. Elsie Effah Kaufmann (Co- Supervisor) Date:. Mr. Samuel Nii Tagoe (Co- Supervisor) Date:. ii

3 ABSTRACT In vivo dosimetry can be considered as a quality assurance tool that supplements port films, computational double checks and phantom studies. It is a process of monitoring the dose (skin, entrance, exit, mid plane and target) delivered to a patient during treatment and plays a role in relating the degree to which the delivered dose matches the prescribed dose. Dosimeters currently used for in vivo dosimetry are diodes, thermoluminescence dosimeters (TLDs), electronic portal imaging detectors (EPID) and radiochromic films. In this research, GafChromic EBT2 film (GAF2) from International Special Products (ISP), USA, was used as the in vivo dosimeter due to its intrinsic characteristics such as tissue equivalent nature (Z = 6.96), wide dose range (0.1cGy 8Gy), ease of handling, very small effective thickness( 0.29mm) and Two Dimensional (2D) dose distribution. The behaviors of both bare film and a specially constructed encapsulated film were considered in this research. The film was cut to 1cm 1 cm and exposed to known doses of radiation determined by ion chamber reading under reference conditions. The optical density (OD) was read with point densitometer and plotted against the corresponding dose to obtain the calibration curve. The response of the film was also determined for changes in field size, source to skin distance (SSD), gantry angle, trays and wedges. Correction factors were then determined for SSD, field size, gantry angle, wedges and tray since the clinical conditions may be different from the reference conditions. Also, skin dose as well as entrance dose calibration factor were determined for both bare and encapsulated GAF2. Knowledge of the surface film reading, calibration factor and correction factors were used in the determination of entrance dose, skin dose as well as percentage skin dose. From a preliminary % skin dose assessment using both bare film and encapsulated film at reference irradiation conditions, encapsulated film was observed to expose the skin to a higher dose (68%) than bare film (23.7%) due to the bulkiness of the encapsulation material and was subsequently eliminated from being used for entrance dose measurements. Furthermore, percentage skin dose determination with bare film was observed to be mainly affected by field size (45%). By varying one irradiation parameter iii

4 and keeping all others at reference conditions, the maximum percentage entrance dose difference observed was 2.6% for a dose prescription of 200 cgy to d max. This falls within the tolerance level of 5% set by accredited regulatory bodies. However, by varying all the irradiation parameters, the maximum percentage entrance dose difference was observed to be higher than the tolerance level. It is therefore recommended to carry out further studies on the use of GAF2 before being implemented as an in vivo dosimeter at KBTH by minimizing and being cognizant of the various levels of errors and uncertainties involved in the determination of the calibration curve, calibration factor and correction factors. iv

5 DEDICATION This work is dedicated to the Almighty God for helping me through this programme successfully and secondly to my mother Miss Christiana Ellen Aboagye for her endless support, encouragement, love, care and prayers. v

6 ACKNOWLEDGEMENTS My sincere thanks go to my supervisors Prof. A.W.K. Kyere, Head of Medical Physics Department, School of Nuclear and Allied Sciences (SNAS), University of Ghana (Legon), Dr. Elsie Effah Kaufmann, Senior Lecturer, Biomedical Engineering Department, University of Ghana (Legon) and Mr. Samuel Nii Tagoe. I would sincerely give thanks to Mr. Samuel Nii Adu Tagoe, the Medical Physicist of National Center for Radiotherapy and Nuclear Medicine (NCRNM) of Korle-Bu Teaching Hospital (KBTH), for his huge guidance, encouragements and support toward the experimental techniques set for this Thesis work. Special thanks go to all the medical physicists at NCRNM for the technical assistance offered. Lastly, am grateful to the National Centre for Radiotherapy and Nuclear Medicine for allowing me to use their facilities for this research work. vi

7 TABLE OF CONTENTS Page TITLE PAGE... i DECLARATION. ii ABSTRACT iii DEDICATION v ACKNOWLEDGEMENTS.... vi TABLE OF CONTENT.vii LIST OF FIGURES......xiii LIST OF TABLES....xv LIST OF ABBREVIATIONS.....xvi SYMBOLS xix CHAPTER 1: INTRODUCTION 1.1 Background Statement of Problem Objectives Relevance and justification Scope and Delimitation Organization of thesis Chapter 2: LITERATURE REVIEW 2.1 Overview of cancer in Ghana...7 vii

8 2.2 In vivo dosimetry for high energy photons Single beam Entrance dose Exit dose Surface dose Target dose Radiation dosimeter Properties of radiation dosimeters Categorization of in vivo dosimeters Characteristics of in vivo dosimeters Thermo Luminescence Dosimeter (TLD) systems Optically-stimulated luminescence (OSL) systems Semiconductor dosimeters Silicon diode dosimetry systems MOSFET dosimeter Gel dosimetry systems Radiochromic film Types of radiochromic films Optical density determination Storage and handling Temperature and humidity dependence viii

9 2.5.2 Water immersion Cutting Marking Energy dependence Post -irradiation optical density growth Scanning Phantoms Phantom selection for in vivo dosimetry Geometric phantom Anthropomorphic phantom Phantom materials Treatment Planning Systems (TPS) Eclipse Pinnacle XiO DICOM Standard Techniques for quantitative comparison of dose distribution Recommendations for using film for dosimetry..51 Chapter 3: MATERIALS AND METHODS 3.1 Materials Equinox 100 Cobalt-60 Teletherapy unit ix

10 3.1.2 GafChromic Films Mini Water Phantom Ionization Chamber Electrometer Densitometer Method Film Preparation Post exposure optical density growth determination GAF2 calibration curve determination Design of encapsulation cap Calibration factor determination of GAF2 for entrance dose measurements Correction factors determination of GAF Entrance dose calculations Skin dose determination Preliminary skin dose assessment for both bare film and encapsulated film.69 CHAPTER FOUR: RESULTS AND DISCUSSIONS 4.1 Evaluation of encapsulated GAF2 cap concepts Post-Irradiation Optical Density Growth Calibration curve determination for GAF Behavior of GAF2 with varying irradiation conditions Correction factor values...79 x

11 4.6 Phantom studies Entrance dose calibration factor determination Entrance dose determination Skin dose determination Evaluation, comparison and selection of better option of GAF2 for in vivo dosimetry Guidelines for the person performing the in vivo Recording of in vivo dosimetry Chapter 5: CONCLUSION AND RECOMMENDATION 5.1 Conclusion Recommendations Oncology and research centers Regulatory Authority Further Research Work. 88 REFERENCE...89 APPENDIX A: AutoCAD Design of encapsulation cap Option One.. 97 APPENDIX B: AutoCAD Design of encapsulation cap Option Two APPENDIX C: Post-exposure optical density growth APPENDIX D: Calibration curve APPENDIX E: Effect of field size on GAF2 response xi

12 APPENDIX F: Effect of SSD on GAF2 response. 112 APPENDIX G: Effect of gantry angle on GAF2 response APPENDIX H: Correction factors for bare film APPENDIX I: Phantom studies (Entrance dose measurements) APPENDIX J: Phantom studies (Skin dose Estimation). 127 APPENDIX K: Proposed IVD Microsoft Excel Chart APPENDIX L: Guidelines for the person carrying out in vivo dosimetry (taken from Barcelona and adapted for GAF2) xii

13 LIST OF FIGURES Figure Page Figure 2.1: Schematic representation of the different doses involved in in vivo dosimetry for a single beam Figure 2.2: Schematic representation of the different doses involved in in vivo dosimetry for 2 parallel opposed photon beams Figure 2.3: Effect of field size on diode response for photon energies of 18 MV X rays and 1.25 MV gamma rays using Scanditronix (EDP-20 and EDE).. 16 Figure 2.4 Configuration of GAF Figure 2.5: Densitometer with film for OD calculation Figure 2.6: Post exposure growth for MD-55-2 GafChromic film Figure 3.1: Equinox 100 Cobalt-60 Teletherapy unit Figure 3.2: 1 cm 1 cm pieces of GafChromic EBT2 film.. 56 Figure 3.3: Mini Water Phantom. 56 Figure 3.4: PTW Freiburg Ionization Chamber. 57 Figure 3.5: PTW UNIDOS electrometer Figure 3.6: PTW DensiX Film Densitometer...59 Figure 3.7: Storage medium for GAF Figure 3.8: Closed film storage medium Figure 3.4: GAF2 calibration factor determination for entrance dose measurements..64 Figure 4.1: Post-exposure optical density growth of GAF xiii

14 Figure 4.2: Calibration curve with standard error bars for GAF2. 73 Figure 4.3: A graph of normalized net optical density against field size of GAF2 with standard error bars...75 Figure 4.4: A graph of normalized net optical density against SSD of GAF2 response with standard error bars. 76 Figure 4.5: A graph of normalized net optical density against gantry angle of GAF2 response Figure 4.6: A graph of normalized net optical density against wedge angle at reference field size of 10 cm 10 cm of GAF2 response with standard error bars..78 xiv

15 LIST OF TABLES Table Page Table 2.1: Effect of photon energy and field size on the dmax determination...16 Table 4.1 Comparison of encapsulation cap (Option One and Option Two)...71 Table 4.2 Comparison of bare film and encapsulated film for in vivo dosimetry. 83 xv

16 LIST OF ABBREVIATIONS 2D, 3D, 4D 2 Dimensional, 3 Dimensional, 4 Dimensional AAPM CAD CEMA CT DICOM DTA FS GA GAF2 IAEA ICRP ICRU IGRT IMRT ISP IVD KATH KBTH KERMA American Association of Physicists in Medicine Computer Aided Design Converted Energy per Unit Mass Computed Tomography Digital Imaging and Communications in Medicine Distance To Agreement Field Size Gantry Angle GafChromic EBT2 Film International Atomic Energy Agency International Commission on Radiological Protection International Commission on Radiological Units Image Guided Radiotherapy Intensity Modulated Radiotherapy International Specialty Products In vivo dosimetry Komfo Anokye Teaching Hospital Korle-Bu Teaching Hospital Kinetic Energy Released per unit Mass xvi

17 LINAC MBS MLC MOSFET MRI MV NCRNM ND NMR OD OSL PDD PMMA PMT PTV QA TG TLDs TPS SD SGMC Linear Accelerator Model Based Segmentation Multi-Leaf Collimator Metal-Oxide Semiconductor Field Effect Transistor Magnetic Resonance Imaging Mega-Voltage National Centre for Radiotherapy Nuclear Medicine Normalized data Nuclear Magnetic Resonance Optical Density Optically Stimulated Luminescence Percentage Depth Dose Polymethyl Methacrylate Photo Multiplier Tube Planning Target Volume Quality Assurance Task Group Thermo-luminescent Dosimeters Treatment Planning System Standard deviation Sweden Ghana Medical Center xvii

18 SSD WHO Source Skin distance World Health Organization xviii

19 SYMBOLS λ D entrance D exit D max Gamma index Entrance Dose for Single Beam Radiotherapy Exit Dose for Single Beam Radiotherapy Dose at Maximum depth for Single Beam Radiotherapy D M1 and D M2 Maximum dose for parallel opposed beams direction 1&2 d M1 and d M2 d max D surface D target Gy I r Depth of maximum dose for parallel opposed beams Depth of Maximum dose for Single Beam Radiotherapy Surface Dose Target Dose for Single Beam Radiotherapy Gray Light Intensity Photon Backscatter Range xix

20 CHAPTER ONE INTRODUCTION 1.1 Background Radiotherapy has emerged to be one of the effective modalities for managing cancer treatment in the country. Radiotherapy is broadly divided into teletherapy where the source of ionizing radiations is at a distance from the patient body and brachytherapy where the source of ionizing radiation is in contact with the patient body. The ultimate goal of any radiotherapy procedure is to maximize the dose to the planning target volume (PTV) while minimizing the dose to the critical structures [1]. For external beam radiotherapy, the patient undergoes a planning computed tomography (CT) scan which is fed into the treatment planning computer. The treatment is simulated and various beam delivery parameters adjusted to achieve the desired distribution of radiation dose inside the patient. The treatment plan is then used to direct the radiation beam from the treatment machine to specific organs or parts of the body to treat the cancer. The treatment is planned to deliver the majority of the radiation to the PTV while minimizing the dose to critical structures in the patient. There are several types of external beam radiotherapy currently used worldwide. The commonly used methods include; fixed beam radiotherapy where there is no modulation in the beam delivered, intensity modulated radiotherapy (IMRT) which includes step and shoot as well as sliding window techniques and involves modulation of the beam using multi-leafs collimator (MLC) and compensators and image guided radiotherapy (IGRT) in which the treatment is guided by images such - 1 -

21 as at CT or Magnetic Resonance Imaging (MRI) [2]. Teletherapy equipment used in the management of cancer in the country are cobalt 60 unit (average energy of 1.25 MV) used in both Korle Bu Teaching Hospital (KBTH) and Komfo Anokye Teaching Hospital (KATH) as well as Linear accelerator (LINAC), energies of 6 MV and 10 MV used in Sweden Ghana Medical Centre (SGMC). The radiotherapy process is a complex multistep procedure which demands the control and assurance of quality in the overall treatment delivery. To be able to assure the delivery of quality, a number of tools are used. These include the use of port films to verify whether the treatment volume is covered and whether the healthy tissues are adequately spared, treatment planning systems (TPS) to plan virtually the number of beams, angle of incidence and modifiers required for the treatment delivery and phantom to simulate the human environment especially in the provision of scatter. Most quality assurance tools are used outside the domain of actual treatment delivery with the patient and makes detection and rectification of treatment errors difficult. Another method of quality control in external beam radiotherapy which complements step-by-step quality control rather than competes with it, is to judge the end-product, which is the treatment as delivered to the patient, and to try to trace back any observed deviations to the faulty step involved [3]. This approach is carried out using specially designed dosimeters placed on the patient s skin or in body cavities to measure the absorbed dose delivered in practice. This method is called in vivo dosimetry and its methodology using GafChromic Ebt2 film (GAF2) is the subject of this study. It is usually performed to detect errors in individual patients [4, 5], to - 2 -

22 detect errors in core procedures [6], to evaluate the quality of specific treatment techniques [7] or to evaluate the dose in situations in which the dose calculation is inaccurate or not possible [8]. The main aim of in vivo dosimetry is to compare the doses derived from the use of specially designed detectors placed on the skin with the theoretical values, as calculated by the Treatment Planning System (TPS). In vivo dosimetry usually involves determination of doses at points close to the skin. One point is close to the entrance, while the other is close to the exit surface of the beam. The corresponding doses are called entrance and exit doses, respectively. 1.2 Statement of Problem In the Radiotherapy Unit of Korle Bu Teaching Hospital (KBTH), Ghana there is no direct monitoring of physically delivered dose to patients. The department therefore relies on the accuracy of the simulation process, phantom studies and treatment planning system calculations. The problem that arises is whether the physically delivered dose to the target volume in the patient actually matches the prescribed dose by the oncologist and whether interventions to correct for any deviations outside the tolerance level and or action level are possible. Therefore the research undertaken used GAF2 to monitor and compare the dose physically delivered to the target volume to the theoretical value prescribed by the oncologist or TPS in order to provide timely and efficient interventions for the correction of underdose and overdose

23 1.3 Objectives The main objective of this study is to demonstrate the viability of both bare and encapsulated GAF2 as an in vivo dosimeter for the determination of entrance dose and skin dose from surface measurement. The specific objectives are: a. Determination of the effect of field size, gantry angle and SSD on GAF2 response and correction factors for each measuring parameter. b. Determination of the effect of beam modifiers such as trays and wedges on GAF2 response. c. Determination of the post-exposure optical density growth for GAF2. d. Determination of entrance dose calibration factor for GAF2. e. Determination of both skin dose and entrance dose for different irradiation conditions. 1.4 Relevance and Justification During external photon beam radiotherapy, clinical complications do arise from underdose of PTV and overdose of normal/ critical organs. These complications arise due to the introduction of errors along the various phases of the radiotherapy process. These errors could arise from patient setup, machine setup, treatment planning system algorithm, manual treatment time/ monitor unit calculations as well as transcription errors. American Association of Physicists in Medicine (AAPM) TG-40 therefore sanctioned that clinics should have access to TLD or other in vivo systems to meet - 4 -

24 the high accuracy in dose delivery expected from complex and conformal therapy techniques [9]. The application of in vivo dosimetry at NCRNM will help provide timely interventions for physically delivered doses outside the tolerance level and or action level and also improve quality in the overall treatment delivery for cancer patients in the country. 1.5 Scope and Delimitation The research focuses on the determination of entrance dose and skin dose for gamma rays from Equinox -100 Cobalt- 60 treatment machine using both bare and specially constructed encapsulation for GAF2. The exit dose determination which is highly dependent on the patient thickness was omitted due to the variability of errors, uncertainties and the long duration associated in the determination of this parameter. The entrance dose and skin dose were the sole in vivo dosimetric parameters considered in this research since it is less dependent on the patient thickness and is defined at depth of maximum dose (d max ) at the entrance side of the patient. The dose measurements were carried out on the central axis of the beam to make the film s positioning easily reproducible. 1.6 Organization of thesis The thesis is arranged in chronological order of five chapters. Chapter One is the introduction to the research work and it provides an idea of the problem to be addressed as well as the general overview of the current state of knowledge relevant to the study. Chapter Two reviews existing literature on similar works relevant to the - 5 -

25 research problem. Chapter Three focuses on the materials and methods used in the study. The results obtained are presented and discussed in Chapter Four. Conclusions of the study, recommendations and suggestions for further study are presented in Chapter Five

26 CHAPTER TWO LITERATURE REVIEW This chapter reviews the relevant literatures on cancer causes and pattern in Ghana, external photon beam categorizations used in cancer treatment, in vivo dosimetry, phantom studies as well as treatment planning systems overview. 2.1 Overview of cancer in Ghana Cancer is a broad group of diseases involving unregulated and uncontrolled cell growth and is known medically as malignant neoplasm. In cancer, cells divide and grow uncontrollably, forming malignant tumors, and invading parts of the body [10]. The cancer may be localized or spread to more distant parts of the body through the lymphatic system or bloodstream. Not all tumors are cancerous; benign tumors do not invade neighboring tissues and do not spread throughout the body. It has been estimated that over 100 different cancers are known to affect humans [11]. The causes of cancer are complex, diverse and only partially understood. Different factors have been predicted to increase the risk of cancer, including tobacco use, dietary factors, certain infections, exposure to radiation, lack of physical activity, obesity, and environmental pollutants [12]. These factors can directly damage genes or combine with existing genetic faults within cells to cause cancerous mutations and are referred to as oncogenic or carcinogenic agents [13]. Inherited genetic defects have been estimated to be the cause of approximately 5 10% of cancers [14]. Different studies have shown that many cancers could be prevented by not smoking, - 7 -

27 eating more vegetables, fruits and whole grains, eating less meat and refined carbohydrates, maintaining a healthy weight, exercising, minimizing sunlight exposure, and being vaccinated against some infectious diseases [12, 15]. Different modalities such as the presence of symptoms and signs, Screening tests or medical imaging make it possible to detect cancer. Once a possible cancer is detected, it is diagnosed by microscopic examination of a tissue sample. Cancer is usually treated with chemotherapy, radiation therapy and surgery. The chances of surviving the disease vary greatly by the type and location of the cancer and the extent of disease at the start of treatment. While cancer can affect people of all ages, and a few types of cancer are more common in children, the risk of developing cancer generally increases with age. In 2007, cancer caused about 13% of all human deaths worldwide (7.9 million). Rates are rising as more people live to an old age and as mass lifestyle changes occur in the developing world [16]. Studies on cancer patterns in Africa are woefully inadequate [17, 18] and populationbased epidemiological data on the occurrence of cancer in sub-saharan Africa, especially, are sparse. Until recently, cancers and other non-communicable diseases were thought to be unimportant public health problems in developing countries, like Ghana, because of the overwhelming high prevalence of communicable diseases. The previous studies have focused on cancer morbidity data from the few established cancer registries in Africa, with very few reporting on cancer mortality data

28 Malignant neoplasm accounted for 914 (2.6%) of all 34,598 admissions, and 141 (5.6%) of all 2,501 deaths at the Korle-Bu Teaching Hospital (KBTH) in the year 1996 [19]. The case fatality risk from malignancies was thus 15.4% [19]. A review of the cancer register in the Department of Child Health, KBTH, over a 40-month period revealed that malignancies accounted for 1.67% of all admissions, with lymphomas (mainly Burkett s lymphomas) being the commonest tumor (67%),followed by retinoblastoma (8.6%), leukemia (8.2%) and Wilm's tumor (7.8%) [19]. The above cancer morbidity trend in the Ghanaian pediatric population is similar to the picture seen in West Africa and other developing African countries [19 33] with the exception of Namibia [34]. Earlier reviews of cancer in Ghana have been mainly restricted to the study of single cancers, rather than the relative contributions of the various cancers to the disease burden because of the absence of a population-based registry. One of such reviews of autopsy material at the KBTH over the 10-year period from (two decades earlier than this study), with emphasis on carcinoma of the pancreas, revealed that pancreatic cancers accounted for 5.8% of all cancer deaths over that period [35]. There has been no recent report on a systematic study of cancer mortality pattern across all age groups in Ghana. The problems of collecting cancer mortality statistics in developing countries have been described [36 39]. Knowledge of cancer patterns in Africa have largely been based on hospital series collected by clinicians and pathologists [36 39]. Despite the problems associated with interpreting data from - 9 -

29 hospital-based series, they are an invaluable source of information on cancer patterns where incidence and mortality data are unavailable. 2.2 In vivo dosimetry for high energy photons One modality of cancer management and treatment in Ghana involves the use of external high energy ionizing photon beams. During any external beam radiotherapy process, different beam combinations are employed to optimize the dose distribution in the planning target volume (PTV) depending on the clinical condition and the goal of that particular treatment. The determination of entrance dose D entrance and the exit dose D exit are very crucial when using a single beam for the therapy procedure. These are defined at points at a certain distance from the patient s surface at the entrance and exit of the beam (referred to as the entrance and exit surface of the patient, respectively). This distance is equal to the depth of maximum build-up, d max when the beam traverses the patient or tissue equivalent phantom and is shown in figure 2.1. The distance d max is chosen as the reference for the determination of entrance and exit dose since dose determination at that point is more accurate and precise. This definition of D exit, symmetrical to D entrance with respect to the midline, is a simplification in the frequent use of two opposed beams, and is useful for the derivation of the target dose D target [40]. It is essential to check each beam contributing to the target dose individually, at least at the first treatment session, in order to identify the possible causes of errors and to correct them. In the particular case of GAF2 dosimetry, that implies the need to change the set of detectors after each irradiation beam

30 As shown on figure 2.1, the surface dose (D surface ) is defined at 0.07mm below the entrance surface, the entrance dose (D entrance ) at d max at the entrance side of the beam, the target dose (D target ) at the depth of dose specification, the exit dose (D exit ) at dmax from the exit surface on the beam axis. The definition of exit dose implies conditions of complete electron backscatter because d max is larger than the electron backscatter range (s) but smaller than the photon backscatter range (r). Once the quality of the irradiation delivered individually by each beam has been checked at the first treatment session, it is also possible to check the reproducibility of the treatment during the following sessions. D entrance D target D surface D exit 0.07 mm d max d max Figure 2.1: Schematic representation of the different doses involved in in vivo dosimetry for a single beam [40]. In order to save time, it is preferable to leave the same in vivo detectors on the patient for the full treatment session including all beams. As long as the entrance and exit

31 doses of each beam are not influenced by contributions from other beams, which are still often well approximated in the case of convergent beams, the definitions of D entrance and D exit are still valid for each beam individually. However this is not always the case, and certainly not for 2 parallel opposed beams as shown in figure 2.2 for which D entrance and D exit lose their meaning and should be replaced by the dose D M,1 and D M,2 at depths of maximum dose from the entrance side of each beam d M,1 and d M,2 respectively [40]. D M,1 D target D M, 2 D M,1 D target D M, 2 D surface, 1 D surface, 2 D surface, 1 D surface, mm 0.07 mm 0.07 mm 0.07 mm d M,1 d M,2 d M,1 d M,2 Figure 2.2: Schematic representation of the different doses involved in in vivo dosimetry for 2 parallel opposed photon beams [40]. From figure 2.2, the beams are equally weighted at the level of the target, which is either at midline with d1 =d2 (A) or in a non-central position with d1 < d2 (B). The entrance and exit doses shown in Figure 2.1 are replaced by the dose D M, 1 at the depth

32 of maximum dose d M, 1 at the entrance side of beam 1, and by the dose D M, 2 at d M, 2 at the entrance side of beam Single beam In most radiotherapy techniques, single beams are the predominant form of radiation employed in the treatment of malignant neoplasm. A single beam radiotherapy technique is usually employed in the palliative treatment of cancerous cells. For cases involving multiple fields (beams) such as breast cancer treatment, head and neck cancer treatment and prostate cancer treatment, there is the need to consider each beam separately and employ the needed in vivo dosimetry system. The doses usually considered in in vivo dosimetry are surface, entrance, exit and target dose and are defined at specified points within the patient Entrance dose When a medium is irradiated by a single photon beam, the area where the beam enters the medium is termed the entrance to the medium. The dose gradually increases from a low value at the surface up to a maximum value at a depth d max. The dose at this depth is termed as entrance dose, D entrance, and depends upon the energy of the incident photon, the collimator opening, the skin-source distance (SSD), the introduction of beam modifying devices and the distance separating them from the patient skin. As shown in figure 2.1, the increase in dose as a function of depth from surface to d max is the steepest just below the surface, gets less pronounced at larger depths and finally levels off at d max at exit side of beam

33 This means that the measurement of D entrance can be carried out with enough material in front and around the detector placed at skin level in order to be reproducible, avoid the determination of dose at very steep dose gradient, minimize the number of headscatter contaminating electrons and minimize the number of correction factors to be determined. Most of the detectors used for in vivo dosimetry have a sensitive part about 1 mm thick, or less, which means that, when they are on the skin, they integrate the dose in a region of very steep dose gradient [40]. This complicates greatly an accurate determination of the ratio between dose to the detector and D entrance. Moreover, when the bare detector is used it is subject to almost the full headscatter contaminating electron spectrum. The number of these electrons is known to increase as a function of the collimator opening and to decrease as a function of SSD, hence the need to introduce correction factors, whose determination is laborious, time consuming and tend to increase the overall treatment duration. In order to limit as much as possible the influence of headscatter electrons on D entrance, build up cap or layering the dimensions of which correspond to the dimensions necessary to ensure full build up for the smallest collimator opening in the absence of any accessory can be used (figure 2.3 and Table 2.1) [40]. Buildup caps are usually used with diodes, TLDs and MOSFET while layering is usually done with film dosimeters such as radiochromic films of which GAF2 is an example. The problem is that with the higher energies the thickness of the build-up cap or layering can be such (several centimeters of tissue-equivalent

34 material) that it might compromise patient comfort and lead to an underdosage of the treatment volume, combined with loss of skin sparing, in a large area. It should be noted that an in vivo dosimeter together with its build-up cap or layering films which are tissue equivalent best simulate the human environment and provides more reliable data. A way to reduce the build-up cap dimensions especially is to use a high density material, but in this case the build-up cap can change the response versus energy of the detector surrounded by it. All the precautions should be taken to avoid errors and uncertainties: e.g. calibration should be performed for the detector together with the build-up cap. As in this case the dimensions of the build-up cap are small, a possible way to decrease the disturbing effect mentioned above, is to introduce small variations in the daily position of the detector in order to smear out the dose perturbation. This is however only possible when the lateral dose distribution is sufficiently homogeneous, and may not be very convenient in practice [40]. As shown in figure 2.3, when the detector is covered with a full build-up thickness, a variation less than 1% is observed either for Cobalt-60 or for 18 MV X-rays, when increasing collimator opening from 5 cm x 5 cm to 30 cm x 30 cm. By contrast, when applying only 2 cm build-up for 18 MV X-rays, the increasing contribution of headscatter electrons causes an apparent sensitivity and a more pronounced increase as a function of collimator opening is observed [40]. The number of correction factors therefore plays a major role when using an in vivo

35 dosimeter with no or inadequate build-up cap material to correct for contaminating electrons. Figure 2.3: Effect of field size on diode response for photon energies of 18 MV X rays and 1.25 MV gamma rays using Scanditronix EDP-20 and EDE) [40]. Table 2.1: Effect of photon energy and field size on the d max determination [40]. Table 2.1 illustrates the dependence of the depth of maximum dose d max on the photon energy and field size. As photon energy increases from Cobalt-60 to 25 MV X-rays, d max also increases. However, for particular photon energy, increasing the

36 field size causes a reduction in d max. These data give an indication of the maximum thickness for the build-up cap to be used on the detector for in vivo dosimetry. In order to achieve most of the electronic equilibrium and yet allow for better and easier handling, it is proposed to decrease the thickness and the lateral dimensions of the cap [40]. Accurate measurements being only possible if the detector is not in too high a dose gradient hence there is the need to be cognizant of the active layer within the detector. The build-up thickness is a compromise between minimum field disturbance and minimum dose gradient over the detector. Moreover in case of incomplete build-up, some headscatter electrons will still reach the detector (figure 2.3). Attention should be paid in this case to the need of correction factors established on a phantom in the same irradiation conditions as for patient. They generally depend on the geometry of the irradiation (field size, SSD, etc) and the presence or absence of accessories on the treatment head. It is important to keep in mind that incomplete build-up on an entrance dose detector might lead to certain correction factors which are only determined by the influence of the headscatter electron contamination and not at all by the intrinsic characteristics of the detector type involved Exit dose A build-down region occurs at the exit side of the patient due to lack of backscatter radiation from the air behind the patient (figure 2.1). Photons as well as secondary electrons play a major role in the determination of backscatter radiation. While the lack of electron backscatter causes a build-down of the dose only in the latter few millimeters in front of the exit surface of the patient,

37 approximately from 1 mm for Cobalt-60 to 3 mm for 20 MV X-rays [40], the lack of photon backscatter influences a much deeper region and increases as a function of field size. This is because the range of electrons in a medium is limited due to its nature of interactions with atoms in the medium. As the stopping power of the medium increases the range of the electrons also decreases. Photons however have a longer range of interactions since they cannot be stopped but attenuated and depend on the attenuation coefficient of the medium. Moreover the extension of the region where an increase of the thickness of the backscatter layer results in a significant increase of the exit dose has been observed to decrease as a function of photon energy: for a 10 cm x 10 cm field, it extends to more than10 cm from the surface for Cobalt-60 gamma-rays but it is only about 1 cm for 20 MV X-rays [40]. Due to this two-component backscatter near the surface of the beam, the position at which the exit dose, D exit should be defined is much less obvious than for D entrance. Mainly for purpose of derivation of the target dose, D exit is taken at d max from the exit surface, e.g. symmetrically with the entrance dose with respect to the midline [40]. In the real patient, at d max from the exit surface, the electron backscatter is complete, but in most conditions only partial photon backscatter is achieved [40]. Mostly the in vivo measurement of the exit dose is performed concomitantly with that of the entrance dose. It is then important when positioning the exit detector, to avoid the shadowing effect of the entrance detector. Also the positioning of the detector at the exit of the patient poses a difficulty and increase the overall treatment time. Because of these problems

38 associated with the determination of exit dose most radiotherapy centers therefore prefer to base their in vivo dosimetry on the entrance dose measurements Surface dose This dose is defined at 0.07 mm under the surface of the skin [41] as shown in figure 2.1 and figure 2.2. To be able to determine surface dose, thin detectors must be used. When the detector used is very thin such as mono-coated photographic emulsions [40], it should be covered with about 0.07 mm of buildup material so that the active layer is at the specified depth. When thicker detectors are used, such as some thermoluminescence (TL) chips or thin layers of TL powder wrapped in envelopes made of paper, they have to be stuck on the skin without any build-up material. Correction factors have to be applied to their response when their effective point of measurement is not at 0.07 mm under their surface [41]. Surface dose is largely influenced by electron dose deposition especially from contaminating electrons from treatment head than from photon dose deposition and tends to increase with increasing field size and decreasing photon energy. Skin sparing is very important especially if the skin is not the primary treatment organ and the accurate determination of the surface dose helps to assess the degree of skin sparing Target dose The dose to the target volume especially the PTV is very crucial in any radiotherapy procedure since the goal of any form of radiotherapy is to deliver the

39 maximum dose to the PTV whilst minimizing the dose to critical structures. A very simplified approach consists of considering the target dose D target equal to the mean of D entrance and D exit. This method, which can be acceptable in some practical conditions, may induce errors of several percent in others especially in the determination of D exit. The target dose can therefore be determined from knowledge of entrance dose, D entrance, PDD values, output factors etc. 2.3 Radiation dosimeter Radiation dosimeter is a device, instrument or system that measures or evaluates, either directly or indirectly, the quantities exposure, kerma, absorbed dose or equivalent dose, or their time derivatives (rates) or related quantities of ionizing radiation. A dosimeter along with its reader is referred to as a dosimetry system. Measurement of a dosimetry quantity is the process of finding the value of the quantity experimentally using dosimetry systems. The result of measurement is the value of a dosimetry quantity expressed as the product of a numerical value and an appropriate unit. To function as a radiation dosimeter, the dosimeter must possess at least one physical effect that is a function of the measured dosimetry quantity and can be used for radiation dosimetry with proper calibration. In order to be useful, radiation dosimeters must exhibit several desirable characteristics. For example, in radiotherapy, the exact knowledge of both the absorbed dose to water at a specified point and its spatial distribution are of importance, as well as the possibility to derive the dose to an organ of interest in the patient. In this context, the desirable dosimeter properties will be characterized by accuracy and precision, linearity, dose or dose-rate

40 dependence, energy response, directional dependence and spatial resolution. Obviously, not all dosimeters can satisfy all characteristics, therefore, the choice of a radiation dosimeter and its reader must be made judiciously, taking into account the requirements of the measurement situation Properties of radiation dosimeters The ideal dosimeter has a number of features making it useful for dosimetry and some are summarized as follows: a. Accuracy: The most important feature of any dosimeter is its ability to correctly measure the dose. This is defined as the dosimeter accuracy. The accuracy may be limited by stochastic and systematic errors. Stochastic errors can be reduced by multiple measurements as they are the result of random variations, which may go in all directions and can thus be minimized with larger data set quantities. Systematic errors can result from items such as fogged film, electrometer leakage or repeated shifts in measurement in one direction. b. Precision: The reproducibility of the results from a measurement technique under similar conditions is defined as its precision. The definition excludes systematic errors and as such does not allow a conclusion to be drawn about the actual correctness of the measured result. It is, however, an important feature for measurements of consistency. Precision is usually defined to a level of 1 or 2 standard deviations of the fluctuations of the measurement around a mean [42]

41 c. Detection limit: The detection limit is a guide to the lowest detectable dose with a certain dosimeter type. Readings which include both fluctuations in the natural background and noise within the detector normally determine this level. d. Measurement range: The measurement range can be defined as the areas from the lowest usable reading to the highest usable reading. e. Dose response: The reading of the dosimeter should be linearly proportional to the given dose, i.e. the readings for an additional dose should be independent of the dose already registered. f. Dose rate response: An ideal detector would be independent of dose rate of delivery and this can be particularly important with the use of pulsed high-energy linear accelerators where pulses of high doses of radiation are delivered in short time periods. g. Energy dependence: For an ideal detector, there should be no difference in the dose response for different radiation qualities. An energy dependence of the dosimeter basically comes down to the fact that different doses can be delivered with the exact same radiation quality but in different materials or tissue types. The requirement of minimal change in dose response with radiation energy usually implies an effective atomic number of the dosimeter being close to that of the material under investigation. This is also an important fact for medical dosimetry where the required absorbed dose, which needs to be measured, is in human tissues of various kinds

42 h. Spatial resolution: An ideal dosimeter should be able to determine the dose in an infinitesimally small volume or the point dose. The location and size of this point should also be well defined in the measurement geometry. Practically, all dosimeters have a finite size and the measurement volume is limited by stochastic mechanisms of dose deposition in microscopic dimensions. Finite volumes can then affect dose measurements in regions of high-dose gradients, such as the penumbral regions of high-energy photon beams, which in turn provide inaccurate readings of delivered dose on a microscopic scale. i. Ease of handling: An ideal detector would be simple to use and physically sturdy enough for clinical and industrial use on a routine basis. There is no point having a dosimeter which is accurate but cannot be used in the situations required due to physical limitations of the measurement conditions. For example, a dosimeter which requires the temperature to be 100 C for measurement of dose in vivo on a patient s skin during radiotherapy treatment would be completely unsuitable. These above physical characteristics have led scientists to search for a radiation detector with a high spatial resolution, which does not require any special developmental procedure and gives a permanent absolute value of the absolute dose. The dosimeter must also have an acceptable accuracy and precision with a relative ease of handling and data analysis. Some of these features have been achieved with the introduction of radiochromic dosimeters such as GAF2. These dosimeters have a very high spatial resolution and relatively low-energy spectral sensitivity. They are relatively insensitive to visible light and thus offer a unique

43 ease of handling as they can be handled and prepared under normal room light. Radiochromic dosimeters undergo a color change directly and do not require chemical processing. The color change can vary considerably depending on the materials used. However, most radiochromic film dosimeters utilize materials which turn a blue color when exposed to radiation. The image formation in radiochromic products occurs as a dye forming or a polymerization process, in which energy is transferred from an energetic photon or particle to the receptive part of the leuko-dye or colorless photo monomer molecule, initiating color formation through chemical changes. At present, radiochromic media for dosimetry can be found in various forms including liquid solutions, gels, waveguides and films. Their dosimetry ranges also cover a wide range from doses as low as 0.1 up to 10 Gy [43] Categorization of in vivo dosimeters Over the past few years, different in vivo dosimeters have been introduced to monitor radiation dose physically delivered to the PTV. Based on their space of dose measurement, they have been categorized into the following: a. Point dosimeters: these are dosimeters that provide point readings. Examples are optically stimulated luminescence, diode, thermoluminescence dosimeter (TLD), metal oxide semiconductor field effect transistor (MOSFET)

44 b. 2D dosimeters: These are dosimeters that provide two dimensional (2D) dose readings and distribution. Examples are radiochromic films (e.g. GAF2), electronic portal imaging device (EPID), diodes array (Map check), etc c. 3D, 4D dosimeters: These are dosimeters that provide three dimensional (3D) or four dimensional (4D) dose readings and distribution. Examples are Gel, Delta4 (2 perpendicular diode arrays), Arc CHECK, 4D Monte Carlo simulation, EPID, etc Characteristics of in vivo dosimeters Different in-vivo dosimeters behave differently in an ionizing radiation field. Their different modes of interactions are due to their unique intrinsic and extrinsic characteristics. Cognizance should be made of these characteristics to optimize the use of these dosimeters for a particular dosimetry procedure Thermo Luminescence Dosimeter (TLD) systems a. When TLDs are exposed to ionizing radiations, electrons are excited and move to trapped states. Heat is then used to de-excite these trapped electrons and in the process release energy in the form of light. b. TLDs are available in various forms (e.g. rods, ribbon, powder, chips, etc.). c. Before they are used, annealing is carried out to erase any residual signal. Appropriate annealing cycles are then used

45 d. Fading of stored signal is one of the major problems of TLD systems and involves the loss in TL signal due to spontaneous emission of light at room temperature. Hence one factor to consider when choosing a TLD is one that has minimal fading with time. Typically, for LiF: Mg, Ti, the fading of the dosimetric peak does not exceed a few percent per year [45]. e. TL dose response is linear over a wide range of doses used in radiotherapy, although it increases in higher dose region exhibiting supra-linear behaviour before it saturates at even higher doses [44]. f. TL dosimeters are relative dosimeters and are calibrated against absolute dosimeters before they are used. Correction factors such as fading, energy correction, and dose- response non-linearity corrections must be determined to effectively derive the absorbed dose from the TL-reading. g. Typical applications of TLD in radiotherapy are: in vivo dosimetry on patients (either as a routine QA procedure or for dose monitoring in special cases, e.g., complicated geometries, dose to critical organs, total body irradiation, in brachytherapy, etc.), verification of treatment techniques in various phantoms (e.g., Rando phantom), dosimetry audits (such as the IAEA/WHO TLD postal dose audit programme) and comparisons among hospitals [44] Optically-stimulated luminescence (OSL) systems a. The principle of operation of Optically-stimulated luminescence (OSL) is similar to that of the TLD. However instead of the use of heat to release the trapped

46 energy, light is used. OSL offers a new technique for the evaluation of in vivo dosimetry in radiotherapy. One common example of OSL dosimeter currently in use is the optical fiber OSL system [46]. b. The optical fiber OSL dosimetry system consists of a laser, collimator, small (~1 mm 3 ) chip of carbon doped aluminum oxide (Al 2 O 3 : C) coupled with a long optical fiber, a beam-splitter and a Photo Multiplier Tube (PMT), electronics and software [47]. c. To produce the needed luminescence (blue light), the chip is excited with a laser light through an optical fiber and the resulting signal (luminescence) carried back in the same fiber is reflected 90 by a beam-splitter and measured in a PMT [48]. d. The optical fiber dosimeter exhibits a linear dose response with high sensitivity over the wide range of dose rates and doses used in radiotherapy. The OSL response is generally linear and independent of energy as well as the dose rate, although the angular response requires correction [49] Semiconductor dosimeters There are generally two types of semiconductor dosimeters; silicon diode dosimetry systems and MOSFET. These dosimeters are emerging as promising candidates for medical dosimetry because they provide direct and quick readings. Their main properties and characterizations are given below

47 Silicon diode dosimetry systems a. Silicon diode dosimeters are characterized as n-si or p-si dosimeters depending upon the base material. They are basically a p-n junction diode system and are produced by taking n-type or p-type silicon and counter-doping the surface to produce the opposite type material [50]. b. Both types of diodes are commercially available, but only the p-si type is suitable for radiotherapy dosimetry, since it is less affected by radiation damage and has a much smaller dark current. c. When a silicon diode dosimeter is exposed to radiation, electron-hole (e-h) pairs are produced in the body of the dosimeter including the depletion layer. The electric field generated due to the intrinsic potential causes charges (minority carriers) to be swept across the depletion region. In this way a current is generated in the reverse direction in the diode [51]. d. Diodes are generally operated without an external bias to reduce leakage current. They are used in the short circuit mode, since this mode exhibits a linear relationship between the measured charge and dose. e. Diodes are comparable to ionization chambers in that they both provide direct reading when connected to an electrometer. However, diodes are more sensitive and smaller in size compared to typical ionization chambers. As compared to ionization chambers, diodes are relative dosimeters and should not be used for

48 beam calibration, since their sensitivity changes with repeated use due to radiation damage [52]. f. Diodes play a useful role in phantom studies, e.g., small fields used in stereotactic radio-surgery or high dose gradient areas such as the penumbra region. They are used for both photon and electron beams dosimetry. For use with beam scanning devices in water phantoms, special packaging in the form of waterproof encapsulation is necessary. When used in electron beam dosimetry, diodes measure directly the dose distribution in contrast to the ionization measured by ionization chambers [53]. g. In most radiotherapy centers around the world diodes are the predominant form of in vivo dosimetry tool. They are widely used on patients or for bladder or rectum dose measurements. Diodes for in vivo dosimetry are provided with buildup encapsulation and are chosen, depending on the type and quality of the clinical beams. Mechanical damage to the fragile diode is limited by the inclusion of the buildup encapsulation [54]. h. Diodes used for in vivo dosimetry are reference dosimeters and are calibrated against absolute dosimetry system such as ionization chamber. Several correction factors are determined and included in the dose calculation. The sensitivity of diodes depends on their radiation history, so the calibration has to be repeated periodically

49 i. The dose response of diodes are influenced by temperature (particularly important for long treatments), dose rate (care should be taken for different source-skin distances), angular (directional) dependence and energy dependence even for small variation in the spectral composition of radiation beams (important for the measurement of entrance and exit doses) [55] MOSFET dosimeter a. The Metal-Oxide Semiconductor Field Effect Transistor (MOSFET), a miniature silicon transistor, seems to be emerging as a promising candidate for medical dosimetry [56]. b. MOSFETs like diodes provide direct and quick readings but are smaller in size when compared to diodes. Their smaller size means less perturbation in the dose distribution. c. A single dosimeter can cover the full energy range of photons and electrons, although the energy response should be examined, since it varies with radiation quality [56]. d. The life-span of MOSFETs is limited and their response is influenced by temperature changes. There exists a non-linearity of response with total absorbed dose hence regular sensitivity checks are required. The bias voltage during irradiation must be stable to avoid fluctuations in the MOSFETs sensitivity. Because their response drifts slightly after the irradiation the reading must be taken in a specified time after exposure

50 e. MOSFETs have been used in radiotherapy applications such as in vivo dosimetry, surface dose measurements, radio-surgery, and brachytherapy measurements Gel dosimetry systems a. Gel dosimetry systems are suitable for relative dosimetry and provide 3D dose distribution evaluations. Apart from being a dosimeter, they serve as a nearly tissue equivalent phantom and serves as a means of measuring the absorbed dose distributions. Depending on the clinical condition they can be shaped to any desirable form [57-58]. b. There are generally two main types of gel dosimeters namely Fricke gels based on the well established Fricke dosimetry and Polymer gels [59]. c. When Fricke gels are irradiated, Fe 2+ ions in ferrous sulfate solutions are oxidized to ferric ions Fe 3+. This causes a change in the paramagnetic properties of the gel and is measured using Nuclear Magnetic Resonance (NMR) or optical technique. This causes a 3D image of the dose distribution to be created. Before irradiation, Fe 2+ ions in ferrous sulfate solutions are dispersed throughout gelatin, agarose or polyvinyl alcohol matrix [60]. d. One of the major problems associated with the use of Fricke gel is blurred dose distribution which arises from continual post-irradiation diffusion of ions [61]. e. In polymer gel, monomers such as acryl amide are dispersed in a gelatin or agarose matrix. Upon radiation exposure, monomers undergo a polymerization

51 reaction resulting in a 3D polymer gel matrix which is a function of absorbed dose that can be evaluated using NMR, x-ray computer tomography, optical tomography, vibration spectroscopy or ultrasound [62-63]. f. There are generally two main groups of polymer gel namely Polyacrylamide gels and the new normoxic gels. The former are referred to as PAG gels and include BANG TM, while the latter include MAGIC gel which is insensitive to the presence of atmospheric oxygen [64-66]. Due to a large proportion of water, polymer gels are nearly water-equivalent and no energy corrections are required for photon and electron beams used in radiotherapy. g. The dose response between the absorbed dose at a point in the gel dosimeter and NMR relaxation rate is semi-linear. The dose map which is derived by computation and by proper calibration procedures is determined by mapping the relaxation rates using NMR scanner [67-68]. h. Artifacts such as image distortion are the major problems associated with the use of polymer gel and arise from post-irradiation effects such as continual polymerization, gelation and strengthening of the gel matrix [69]. i. The temperature at which the dosimeter is evaluated as well as the strength of the magnetic field has an influence on the dose response although no significant dose rate effects in polymer gels have been observed using NMR evaluation [70-71]

52 2.4 Radiochromic film Radiochromic films are emerging to be one of the effective modalities for in vivo dosimetry. They are used as relative dosimeters calibrated against absolute dosimeters such as ionization chamber. It is possible to achieve precision better than 3%, if proper care is taken in their calibration and with the environmental conditions [72]. They contain a special dye that polymerizes upon exposure to radiation. Radiochromic reactions can be defined as a direct coloration of a media by the absorption of radiation, which does not require any latent thermal, optical or chemical development or amplification as seen in radiographic film use. The radiochromic reaction is a solid state polymerization, whereby the film changes color proportionately to the radiation dose. When the film is exposed to ionizing radiation, electrons are released which initiate a series of reactions in the active monomer components in the coating of the film resulting in the formation of a deeper colored polymer product. The degree of transmission of light through the polymer helps to determine its optical density from which calibration protocols can be related to absorbed dose. Radiochromic film is self-developing, needs neither developer nor fixer. Since the radiochromic film is grainless as compared to most radiographic films, it has a very high resolution and can be used in high dose gradient regions for dosimetry, e.g., near brachytherapy sources, in measurement of dose distributions in stereotactic fields, etc [73-74]

53 2.4.1 Types of radiochromic films Over the past years, different radiochromic films have been developed for various uses in radiotherapy. The most commonly used radiochromic film is the GafChromic film produced by International Specialty Products, USA [75]. Examples of GafChromic films include EBT, MD-55-1 and MD These films are usually light blue in color before irradiation and progressively become bluer with increasing dose. GafChromic film called GAF2 has been introduced for in vivo dosimetry due to the discontinuation of EBT film. GAF2 is light yellow before irradiation but becomes green after irradiation. The green appearance is caused by the radiation interacting with the active component in the film which creates a blue colored polymer (i.e. blue inside a yellow film gives a green color). The configuration of GAF2 is shown in figure 2.4 below. The active part of the film has been reduced to a single layer about 30µm, 175µm polyester substrate, 50 µm polyester overlaminate, 25 µm adhesive layer and 5 µm topcoat. GAF2 contains a yellow marker dye that protects the active layer from exposure by ultraviolet and visible light and reduces the effect from these sources by several times [76] and the dye could cause minimal intra- and inter-sheet non-uniformity of the films

54 Figure 2.4 Configuration of GAF2 [76]. According to the manufacturer, the main atomic compositions of the overall composition of GAF2 are H: 40.9%, Li: 0.1%, C: 42.4%, O: 16.6%. The effective Z of GAF2 is 6.98 which make them approximate the human environment whose effective Z is 7.00 [77]. The dosimetry with GAF2 has a few advantages over the radiographic films, such as the ease of use, not requiring dark rooms, film cassettes or film processing; dose rate independence; better energy characteristics except for low energy X- rays (25 kv); insensitivity to ambient conditions (although excessive humidity should be avoided). GafChromic films are generally less sensitive than radiographic films and are useful at higher doses, although the dose-response nonlinearity should be corrected for in the upper dose region

55 2.4.2 Optical density determination When radiochromic film is exposed to ionizing radiation, solid state polymerization reaction causes visible color change of the film. This coloration is due to an attenuation of some of the visible light coming through the developed film, resulting in a graying of its appearance. The reduction in light passing through the film is a measure of its blackness or optical density (OD). A pivotal assumption in film dosimetry is that the dose to the film is reflected in the resulting optical density of that film. The information from film dosimetry can be utilized in two ways, either qualitatively or quantitatively. Without a doubt, the film has no equal for communicating qualitative information about the dose distribution. However, accurate quantitative film dosimetry presents a few technical challenges. To determine the optical density of the film a densitometer of known light intensity (Io) is used. The transmission of light through the polymerized film is recorded in the densitometry system and used in the calculation of the optical density. This relationship can be expressed as follows: Optical density (OD) = log 10 (I0 / I) Transmission = I / I0 Where I0 is the light intensity with no film present and I is the light intensity after passing through the film. It should be noted that I0 / I has an exponential relationship to the dose making optical density appropriately linear with dose. The acceptance of this relationship has led to the wide use of the film as a dosimeter

56 A simplified operation of a densitometer is shown in figure 2.5. This consists of a light source and a detector on the other side. To ensure the proper operation of theses densitometers, quality checks should be carried out on both the light source and detector. There are generally two main types of densitometers; these are point densitometers and scanners. Point densitometers are appropriate for point measurements while scanners are used for 2D measurements. Film IO Figure 2.5: Densitometer with film for OD calculation [79] 2.5 Storage and handling The manner in which a film is stored and handled has an influence on its dosimetry quality. Errors and uncertainties are introduced in radiochromic film measurements due to physical damage caused during shipping or handling of the film. Mishandling of Films causes static charges to be produced on the outer layers of the films and causes dust to be deposited easily on the film. To minimize the effect this has on measurements, the film is wiped with lintless paper before use [78]

57 Films should not be handled with bare hands since skin oils can cause absorbance measurements to occur spontaneously. Based on these problems, it is recommended that tweezers and gloves be used when handling films for in vivo dosimetry [79]. Storage and handling conditions to be considered are temperature, humidity, water immersion, cutting, marking, energy dependence, post irradiation optical density growth and scanning. The effect of the factors/ conditions above on GAF2 is described below Temperature and humidity dependence Factors which may influence the response of radiochromic dosimeters are temperature, relative humidity and in some cases ambient light and gases. Variations of response with surroundings must be determined since environmental conditions during calibration may be different from practical use [80]. GAF2 has been designed to be handled in interior room light; however, it is recommended that the film be kept in the dark when not in use and exposure to sunlight should be avoided. The sensitivity of GAF2 to room light is approximately 10 times lower than EBT due to the incorporation of the yellow marker dye in the film [79, 81]. Storage of both EBT and GAF2 may be done at room temperature (20 25 C), but storage at refrigerator temperature or less helps to prolong the shelf life of the films. The shelf- life of the film is two years when stored at room ambient temperature. Brief exposures (e.g., < 1 min.) to temperatures up to 70 C, or more prolonged exposures (e.g. < 1 day) at temperatures of 50 C do not affect the response of EBT or GAF2 [82]

58 2.5.2 Water immersion The sandwich style of GAF2 makes it possible to take measurements directly in water without the need for waterproof encapsulation as evident in diodes and TLDs. When the film is immersed in water the inbuilt emulsion softens and turns milky color, however when dried, the clear color is mostly restored. The duration of film immersion should be controlled to limit the degree of water penetration between cut edges. Film immersion in water can be done up to an hour without significantly changing the properties of the film since the rate of diffusion is very slow and also because the active layer is protected by two layers of polyester. For this reason, it is advisable to avoid taking readings close to cut edges of films which have been irradiated in water since at these edges water can easily come in contact with the active layer [78, 82].Water penetration has been observed to reach a depth of 9 mm around cut edges and that as the immersion time increases so does the dose error [83] Cutting GAF2 configuration is such that different layers are brought together to form the whole. Care should therefore be taken when cutting the film to avoid separation of the layers of the film [79]. It is therefore recommended that scissors, guillotine, scalpel or sharp knife be used when cutting the film since their use produce good results and minimizes mechanical damage to the film [78, 82]. The cut edges of the film may be stressed and should be avoided for dosimetry analysis. It is also

59 recommended that the light analyzing beam be kept about 1.5 mm away from the cut edge [84] Marking During the use of EBT2 film for photon or electron beam dosimetry studies, they are usually cut into square pieces. Identification of individual and group films therefore becomes necessary for easy reference. Since the outer layers of EBT and GAF2 are polyester, they can be marked with a pen without damaging the active layer. If the marks interfere with scanning, or measurements, they can be removed with a soft rag or tissue moistened with a solvent which does not damage the polyester [84]. Fiducial markers on films also allows consistent placement of films relative to the radiation during irradiation. Fine- pointed felt-tip permanent markers are suitable for writing both the identifying numbers and for making fiducial markings on the film, which are necessary for reading the same film more than once and registering the multiple images together [79]. 2.6 Energy dependence In external beam radiotherapy, varying energies of electrons and photon beams from linear accelerators (LINAC) and Cobalt-60 units are employed to achieve a specific goal. The degree of energy dependence of a particular dosimeter can affect the dosimetry properties of that dosimeter when an unknown spectrum of radiation energies is present. To make GAF2 less dependent on energy, low effective atomic number elements are involved in its composition as compared to high density

60 materials such as those used in radiographic films making radiographic films highly dependent on energy [80]. In vivo dosimetry of GAF2 requires dose deposition (photoelectric absorption) within the dosimeter used in the measurement. Photoelectric absorption of the dosimeter which is a measure of its sensitivity increases as atomic number increases. To improve the sensitivity of the film, small amounts of sulphur, chlorine, potassium, and bromine are added [78]. These elements are added to boost the photoelectric absorption of KeV photons. GAF2 exhibits better energy dependence than earlier radiochromic films including EBT films. GAF2 has been shown to have low energy dependence with about 10% difference in response between 6 MeV and KeV photons [78]. 2.7 Post irradiation optical density growth The active component in GAF2 is a radiation-sensitive monomer. Upon irradiation, the active layer polymerizes to form a polymer colored dye. The reaction has an incubation period of at least 1 µs [84]. Furthermore, after irradiation has ceased polymerization proceeds causing a continuous increase in optical density of the film. This corresponds to an increased amount of formed polymer in the active layer. The rate of change in this optical density however diminishes rapidly with time and the optical absorption seems to asymptote to a constant value approximately 48 h after exposure has finished [84]. Figure 2.6 shows the post-irradiation growth of the optical density for MD-55-2 film, a type of GafChromic film, for various delivered absorbed

61 doses after exposure has ceased. As can be seen in the figure, a significant increase occurs over the first few hours and the film becomes relatively stable after approximately h. The manufacturers of GAF2 report that the density of the developmental film increases with time following exposure [84]. For all GafChromic dosimetry film, the density growth is approximately proportional to the log of the time after exposure [84]. Due to this relationship, films should not be scanned immediately after exposure because errors in the time of measurement could have a significant effect on dose accuracy. Figure 2.6: Post exposure growth for MD-55-2 GafChromic film [84] 2.8 Scanning GAF2 can be read with a film scanner or digitizer. The response of the film is enhanced if the spectral response of the scanner is matched to the absorbance of the film. GAF2 have their greatest response in the red color channel [78, 82]. Artifacts such as contrast bands are the major problem associated with the use of some scanners for assessing the optical density of radiochromic films. They are

62 referred to as Newton s rings or interference patterns and are produced across the top and bottom of the images [78, 82]. These contrast bands can be eliminated by scanning the films in a clear polyester sleeve or mask which holds the film off the glass. The use of the polyester sleeve will eliminate the interference patterns from the digitized image; however, a background correction is still necessary to achieve optimum results [82]. There are generally two modes for scanning films; these are reflection mode and transmission mode. Both modes are now currently available on most scanners. Scanning in the reflection mode has the benefit of not producing Newton s rings and it also results in more signals as the light passes through the active layer twice. In transmission mode, there is less noise and the lateral response effect is less [78]. According to the manufacturer s specifications, the best response is obtained when scanning in transmission mode [82]. Scanning in transmission mode reduces light scattering due to the white reflection backing of the scanner used for reflection. Due to the construction properties of GAF2, it is not symmetric. How the film is positioned on the scanner needs consideration to make results reproducible. The film should either be placed face up or face down consistently on the scanner. Calibration curves produced using both EBT and GAF2 scanned in landscape orientation in both reflection and transmission mode on an Epson 10000XL scanner indicated that face up or face down scanning affects the measurements when scanning in reflection mode however there were no statistically significant differences found when scanning in

63 transmission mode [82]. It has been demonstrated that potential errors of 17. 8% can be recorded by inverting the scanning side when performing gamma analysis on the head and neck IMRT plans (criterion of 3mm and 3%) [78]. Emphasis should therefore be placed on the importance of consistently placing the film either face up or face down when scanning. 2.9 Phantoms The use of phantoms for in vivo dosimetry is undeniably very important. Phantoms are meant to replace the patient in the radiotherapy process for activities such as determination of the output of the Cobalt-60 unit, output factors, calibration of dosimeters such as GAF2, provision of scatter, etc Phantom selection for in vivo dosimetry Verification processes for in vivo dosimeters vary significantly in their phantom requirements, with the appropriate phantom determined by the purpose of the measurement. Phantoms are typically constructed using either water or waterequivalent plastic. Open water phantoms can be used when the beam is perpendicular to the phantom surface, and where great flexibility in detector positioning is desired. With the proper procedures and design, water-equivalent plastic phantoms can support multiple detectors, radiographic film, GAF2 and rapid and efficient setup reproducibility. Such phantoms can also include the substitution or addition of heterogeneous materials. To conduct an overall

64 evaluation of an in vivo dosimetry system, anthropomorphic phantoms are useful in conjunction with other phantoms Geometric phantom The measurement of single point and planar doses are usually performed with simple geometric phantoms that can accommodate ionization chambers and films. Cubic phantoms, comprised of slabs, are easy and accurate to set up and allow for measurements at multiple depths. The slabs can be water equivalent or built with materials having relative electron densities representing specific anatomical tissues. To improve setup accuracy, the phantom is scribed with setup lines whose positions are accurately known with respect to the dosimeters. The use of fiducial marks on the film is often considered for registration of the film with respect to the phantom. Other geometric phantoms exist; these are rectangular phantom and cylindrical phantom. Rectangular phantoms are useful for measuring single field or composite dose distributions. Cylindrical phantoms have a convenient geometry for coplanar composite IMRT delivery verification, while allowing for multiple ionization chamber positions. There are also slab phantoms that are machined to accept commercial ionization chambers Anthropomorphic phantom Anthropomorphic phantoms are fabricated in the shape of a human and, consequently, they can be more difficult or at least more time consuming to accurately set up and align with respect to the Cobalt-60 unit. The preparation and

65 accurate placement of film can be more difficult than with geometric phantoms. The preparation of films involves cutting film to match the shape of the external phantom shape and sealing the phantom around the film with light-tight tape if light-sensitive films are used. In spite of the added difficulty, anthropomorphic phantoms have been effectively used for limited measurements to evaluate the process of patient treatment planning and delivery and to identify treatment planning or dose delivery problems that are not evident in simple homogeneous geometric phantoms. While anthropomorphic phantoms are good for assessing the overall fixed beam planning and delivery process, many commercial phantoms are composed of thick transverse slices, which limit the flexibility in film and point-dosimeter placement. Another problem is that causes of dose distribution discrepancies are difficult to isolate using an anthropomorphic phantom. Therefore, additional measurements using geometric phantoms may also be required to aid in the interpretation of any discrepancies between measurements and calculations Phantom materials Phantoms should be made of a water-equivalent or known electron-density material so that the treatment planning system can accurately calculate dose to the phantom. A large number of such phantoms of different shapes made of water equivalent materials are commercially available. When non- water-equivalent materials such as PMMA and polystyrene are used, validation of the dose

66 distribution calculation algorithm should be conducted before clinical use. Additional considerations are required when radiographic film is used. Optical and ultra violet light will expose the film, so the phantom must be light tight and internally opaque to prevent exposure by Cerenkov radiation [85]. The use of phantoms with lead (Pb) or other high atomic number materials is not recommended Treatment Planning Systems (TPS) Different radiotherapy treatment equipment requires specific treatment planning systems. Multiple treatment planning systems are currently available worldwide. Treatment planning systems plan virtually the actual radiotherapy treatment on the computer with the sole aim of ensuring the delivery of maximum radiation dose to the PTV whilst sparing critical structures as much as possible. For in vivo dosimetry procedures, the doses derived from the TPS calculations are compared with that recorded by the in vivo dosimeter such as GAF2. Most TPS enable both 2D and 3D virtual simulation of the treatment process with the dose distribution predominantly presented in the form of Dose Volume Histogram. Below is a short explanation of 3 of the commonly used planning systems Eclipse This is a treatment planning system developed by Varian Medical Systems. The Eclipse treatment planning system is a comprehensive solution that is open, integrated and easy to use. Its open architecture supports most treatment

67 modalities and works across numerous linear accelerator platforms. It enables the planning of sophisticated treatments, including 3D conformal, Intensity-modulated radiation therapy (IMRT), electron, proton and brachytherapy with a full palette of powerful tools and robust functionality [86]. Its open (DICOM RT) architecture allows connection to multiple imaging devices needed for treatment planning Pinnacle Pinnacle is the treatment planning system produced by Philips. It consists of fully integrated photon, electron, stereotaxy, brachytherapy, simulation, image fusion, and IMRT options which allow treatment planning tasks to be performed from a single platform. Pinnacle s Collapsed Cone Convolution Superposition algorithm provides an accurate, true 3D dose calculation algorithm. Model Based Segmentation (MBS) software designed by Philips includes an anatomical library of 3D patient organ structure models, which reduce the time oncologists spend manually drawing contours. IGRT workflow enhanced with MBS software allows clinicians to propagate organs to alternate 4D datasets to determine the extent of tumor movement within the patient [87] XiO XiO is a treatment planning system produced by Computerized Medical Systems Software Inc recently acquired by Elekta. XiO is a comprehensive 3D IMRT treatment planning platform that combines the latest tools and more robust dose calculation algorithm. XiO supports a range of treatment modalities, including 2D,

68 3D, MLC-based IMRT, solid compensator-based IMRT and brachytherapy. In addition, dynamic conformal arc therapy and stereotactic delivery are supported. XiO offers advanced dose calculation algorithms, including Clarkson, Full Fourier Transform Convolution, Multi-Grid Superposition, Fast Superposition and Electron Monte Carlo. Clinicians can therefore choose the algorithm that is most appropriate for each plan [88] DICOM Standard DICOM stands for digital imaging and communications in medicine and serves as a standard which standardizes the format in which medical imaging information is transferred. Different planning system use language peculiar to that planning system and makes communication between different planning systems difficult hence the need to introduce DICOM standard which provides a common language for the different TPS. The main aims of DICOM standard as given in a report by the National Electrical Manufacturers Association in 2003 are to: a. Create a standard for the way devices interact using a specific computer language. This includes how devices are expected to react to commands and associated data, not just the information being moved between devices. b. Address the semantics of file services, file formats and information directories necessary for off-line communications

69 c. Facilitate operation in a networked environment. d. Accommodate the introduction of new services, thus facilitating support for future medical imaging applications Techniques for quantitative comparison of dose distribution In order to ensure thorough patient-specific quality assurance radiotherapy, there is the need to adequately compare the dose distribution between measured (by in vivo dosimetry) and calculated data points by the TPS. Qualitative and quantitative evaluation techniques have been derived for dose distribution comparison. The physicist first determines a set of irradiation conditions for which the treatment planning system is to be evaluated. Measured (often planar) dose distributions are obtained for these geometries, and the corresponding isodose distributions are subsequently displayed or printed. Qualitative evaluation of the treatment planning system calculation is made by superimposing the isodose distributions, either using software tools or by hand using printed isodose distributions and a light box [89]. This evaluation highlights areas of significant disagreement, but a more quantitative assessment may be needed for final system approval. Quantitative evaluation methods directly compare the measured and calculated dose distribution values. Various studies have described the quality assurance procedures of treatment planning systems and subdivided the dose distribution comparisons into

70 regions of high and low dose gradients [90-94], each with a different acceptance criterion. In low dose gradient regions (e.g. in the main radiation field), the doses are compared directly using the dose difference technique, with an acceptance tolerance placed on the difference between the measured and calculated doses (that is ± 3%) [98]. In high dose gradient regions (e.g. near the penumbra region), the dose difference technique fails and therefore a new technique known as the distance to agreement (DTA) is used [96]. The acceptance tolerance placed on the DTA technique is 3 mm. Because both the dose difference and DTA if used singly have their strengths and weaknesses a new technique known as gamma (λ) dose distribution comparison is usually employed [97]. This new technique combines both the dose difference and DTA technique and makes them complementary to each other. Before computing λ, the dose and distance scales of the two distributions, referred to as evaluated and reference, are renormalized by dose and distance criteria, respectively. The renormalization allows the dose distribution comparison to be conducted simultaneously along dose and distance axes. The λ quantity, calculated independently for each reference point, is the minimum distance in the renormalized multidimensional space between the evaluated distribution and the reference point. In typical clinical use, the fraction of points that exceed 3% and 3 mm can be extensive, so typically 5% and 2 3 mm are used in clinical evaluations [98] Recommended procedure for using radiochromic film dosimetry The AAPM Radiation Therapy Committee Task Group No. 55 (1998) summarized the recommended procedure for radiochromic film dosimetry as follows:

71 a. When selecting a scanning densitometer, the signal to noise ratio of the scanning equipment should be kept in mind. An 8 bit densitometer only provides 256 shades of grey, which may not be enough to give good images at low dose levels. b. The characteristics and/ or limitations of the scanner should be considered. c. The maximum optical density for which the densitometry system will provide a reading should be known. This should be verified and acceptable for intended use. d. The sensitive emulsion layer(s) of radiochromic films absorbs strongly in the red wavelength (about 660 nm) region. Thus, the densitometer response is optimized at the wavelengths. It is desirable to have the wavelength of the light source in the densitometer between 600 and 670 nm, which are the wavelengths of the two main absorption peaks, thereby maximizing the signal obtained from the system. e. It must be kept in mind that the measured optical density is determined by the absorption spectrum of the sensitive emulsion layer(s) of the film as well as the spectrum of the readout (scanner) light source. Thus the use of broad band light source, such as white light source, may not yield the desired contrast levels especially when used with an 8 bit pixel depth. f. Prior to use, films should be visually inspected and handled with care. g. Films should always be kept in a dry and dark environment at the temperature and humidity at which they will be utilized for clinical purposes

72 h. Since the radiochromic films are sensitive to fluorescent light and to sunlight, they should be read and handled in normal incandescent light. i. The lot number and model number of the film should be noted. This will allow the user to verify any variation in the manufacturing of the film. j. Film orientation and alignment should be noted to minimize polarization effects. k. Since the film response changes with time, especially during the first 24 hours after irradiation, the exposure time and readout time for all the films should be documented and if necessary, appropriate correction factor for instabilities should be applied. The recommendation is to read the film at least 24 hours (preferably 48 hours) after the exposure. l. Film uniformity should be examined. If necessary, the double exposure technique should be considered to improve film uniformity. m. Radiochromic film should be calibrated in a large well-characterized uniform radiation field. n. The dose response curve and film sensitivity should be obtained in the dose range and conditions of interest

73 CHAPTER THREE MATERIALS AND METHODS This chapter describes the experimental methods used in this study. It was carried-out at the Radiotherapy unit of the National Centre of Radiotherapy and Nuclear Medicine (NCRNM), Korle-Bu Teaching Hospital (KBTH). 3.1 Materials The materials used in this study were Equinox-100 Cobalt-60 teletherapy unit, GafChromic EBT2 films, Mini water phantom, Ionization chamber, Electrometer, Densitometer, Polymethyl methacrylate (PMMA), Specially constructed PMMA encapsulation cap, Thermometer, Barometer and Prowess 4.60 Treatment Planning System (TPS) Equinox-100 Cobalt-60 Teletherapy unit The Theratron Equinox is the newest member of the Theratron line and one of the most advanced Isotope based teletherapy system available on the market as shown in figure 3.1[105]. The machine offers asymmetric jaws and motorized wedge. Completely integrated with an Avanza couch, the Equinox provides state of the art controls with modern proven technology. The machine can facilitate advanced treatment techniques and reduces treatment times with numerous time saving features. The machine is controlled by an advanced software control system which monitors treatment parameters

74 continuously. In addition, the machine can be optionally configured to communicate with all generally available record and verification systems. Figure 3.1: Equinox-100 Cobalt-60 Teletherapy unit GafChromic Films GafChromic EBT2 film with Lot# A manufactured by international specialty Product was used. The film can measure doses from 1cGy to 10Gy. The film was handled in interior room light and kept in the dark case when not in use. Exposure to sunlight was avoided since the film may darken. It was exposed, measured and stored at room ambient temperature of 25 C.The films were interleaved with a tissue paper which provided a homogeneous environment around individual pieces of film. Each sheet has a size of 8 cm 10cm, which was cut to 1cm 1cm for calibration as shown in figure

75 Figure 3.2: 1 cm 1 cm pieces of GafChromic EBT2 film Mini Water Phantom A mini water phantom of dimension; was used by the oncology centre where the research was carried out to measure beam output of their Equinox-100 Cobalt-60 teletherapy unit. It is made of Perspex (PMMA) and has a field size of 10 cm x 10 cm inscribed on one of its surfaces. At one of the sides is a hole provided by the manufacturer to accommodate 0.6 cc farmer type ionization chamber. On one of the surfaces is an opening used for filling the phantom with water for the beam output measurement. Figure 3.3 is a diagram of the mini water phantom used in this research showing reference field size condition. 10 cm x 10 cm Figure 3.3: Mini Water Phantom

76 3.1.4 Ionization Chamber The Ionization Chamber used for this study was a cylindrical farmer chamber type of model PTW30001, manufactured by PTW Freiburg with Serial number of 1510, and was used with the mini water phantom for the beam output measurement. It is cylindrical in shape with a 0.6 cm 3 volume. The chamber has calibration traceability to the International Atomic Energy Agency secondary standard laboratory. The chamber has a calibration factor, N D,W of 5.17 determined with chamber bias voltage of V at temperature of 20 C and pressure of kp for humidity not exceeding 70 %. The ionization chamber was used to establish dosimetry protocol for the GafChromic film. Figure 3.4 is a diagram of the ionization chamber used in this research. Figure 3.4: PTW Freiburg Ionization Chamber Electrometer The electrometer used was PTW UNIDOS model with Serial Number T Connected to the ionization chamber, it was used to quantify the charges detected

77 by the ionization in nanocoulomb/minute (nc/min) to help in the evaluation of absorbed dose. The electrometer was calibrated together with PTW farmer type ionization chamber. Figure 3.5 is a diagram of the electrometer used in this research. Figure 3.5: PTW UNIDOS electrometer Densitometer The densitometer used in this study was manufactured by PTW FREIBURG (Germany) with serial number T called DensiX Film Densitometer. DensiX is a manually operated film densitometer for measuring the optical density of processed X-ray or radiochromic films, exposed by a sensitometer (light exposure of a step wedge) or by X-ray or Gamma-ray equipment. The length of the measuring arm is 20 cm, which makes it possible to measure the optical density even in the middle of 35 cm x 43 cm large size films. Figure 3.6 is a diagram of the densitometer used in the research

78 Figure 3.6: PTW DensiX Film Densitometer 3.2 Method The calibrated output (dose rate) of the Cobalt-60 unit at the time of irradiation was always determined. The calibrated output was determined at reference conditions of field size (FS) of 10cm 10cm, source to skin distance (SSD) of 100cm, gantry angle (GA) of 0. Beam modifiers such as wedges, trays and blocks were omitted during the determination of the calibrated output of the treatment machine since they do not allow accurate determination of the dose rate from the machine. 3.3 Film Preparation The films were handled according to accredited international protocols [81]. Films were cut to 1cm 1cm using scissors and the corresponding faces of the films noted with markers as shown in figure 3.2. The faces of the cut films were noted with markers to ensure consistency in side orientation exposures. This step was necessary due to the asymmetric structure of the film. Cut edges and marked faces were omitted from film densitometry due to the large margin of errors introduced. Cutting of the films was done with precaution to minimize surface scratches and surface oiling from

79 fingerprints. The films were then kept in a dark airtight box to prevent unnecessary exposure to light (e.g. ultra violet light) and moisture as shown in figure 3.7 and figure 3.8. The optical densities of the films before irradiation was always measured to correct for background radiations and recorded as OD1. The point densitometer was always warmed for 5 minutes before being used to take readings to help produce reproducible results. Figure 3.7: Storage medium for GAF2 Figure 3.8: Closed film storage medium Cut and stored films were then used for the calibration and correction factors determination, skin dose as well as entrance dose determination experimentations. Post exposure optical density growth determination Four sets of films were used. Each set consisted of three films and were irradiated under the same conditions to minimize systematic errors. The four sets each received radiation doses of 50 cgy, 200 cgy, 400 cgy and 800 cgy respectively. The goal of the study was to monitor the post exposure optical density growth of the irradiated film. Conditions of field size, source to surface distance, gantry angle could be

80 arbitrary however, for a particular set the same irradiation conditions were used. The irradiated film signal was then read with a point densitometer at intervals of 1, 60, 120, 180, minutes. This was meant to provide the optimum time to determine the optical densities of an exposed film since literature has shown that the optical density of an irradiated film increases with time [102]. 3.4 GAF2 calibration curve determination The calibration process was carried out using the IAEA TRS398 protocol. The EBT films used for calibration were cut in 1.0 cm 1.0 cm size, and sandwiched between solid water slabs of dimension 30 cm 30 cm 25 cm perpendicular to the beam direction. Ten sets of films with each set comprising three films were irradiated to different durations under reference conditions (field size = 10cm 10cm, Source Surface Distance = 100cm, Gantry angle = 0, depth = 5.0 cm) to give doses of 50cGy, 100cGy, 150 cgy, 200 cgy, 250 cgy, 300 cgy, 350 cgy, 400 cgy, 450 cgy and 800 cgy. The set of films were then scanned using a point densitometer and their mean optical density recorded as OD2. The net optical density (NOD) obtained by subtracting the optical density of background radiation (OD1) from the optical density of irradiated films (OD2) was then determined. A graph of dose against net optical density was then plotted using Microsoft Excel to obtain the calibration curve and calibration equation. The calibration curve helps to interpolate any film signal (optical density) to absorbed dose. The process of carrying out this is called calibration of the film. Note:

81 NOD = OD2 OD1. (1) All doses were previously determined by Farmer ionization chamber dosimetry at identical conditions. Temperature and pressure corrections were also determined. 3.5 Design of encapsulation cap In this research, Perspex (PMMA) was used for the construction of the encapsulation cap and needed a thickness of 0.5 cm above the GAF2 to ensure electronic equilibrium for precise and accurate dose measurements. In the design of the encapsulated cap, two concepts were considered and evaluated. In silico designs of the cap were created using Autodesk Mechanical Desktop. 2D and 3D dimensions as well as assembly designs and constraints of the cap were done using the same software. After the best concept had been chosen, it was taken to a machine shop to be constructed. 3.6 Calibration factor determination of GAF2 for entrance dose measurements GAF2 was calibrated to measure entrance dose. Calibration factor was determined for both bare film and encapsulated film. To determine the calibration factor, the film was first placed in the chosen encapsulated cap, and then positioned on the surface of a suitable calibration water phantom at reference conditions of field size of 10 cm 10 cm, SSD of 100 cm, d max of 0.5cm, and gantry angle of 0. The ion chamber was placed inside the phantom on the central axis, at reference depth as shown in figure

82 3.4. According to the definition of entrance dose, this should be the depth of maximum dose (d max ). The ion chamber is thus probing the depth dose curve at its maximum, and not at its subsequent fall- off. During measurements the electrometer is set in the integral mode to measure charges within 60 seconds intervals. The electrometer has a polarity switch at the back, which allows the user to change the bias voltage polarity of the connected chamber. In the front of the electrometer are; a display panel which shows charge or current reading, measurement mode, duration of measurement and option menus; there are bottoms to enable the user to enter into library of chambers, select reading range and change bias voltage of chambers as well as measuring intervals. To determine the calibration factor, equation 2 was used. Fcal = (2) Where: Fcal is calibration factor Ric is the reading of the ion chamber Rf is the reading of GAF2-63 -

83 The setup used for the determination of the calibration factor is shown in figure 3.4. SSD = 100cm GAF2 Ion chamber d max =0.5 Solid Field size= 10 cm 10 cm Figure 3.4: GAF2 calibration factor determination for entrance dose measurements. The ionization chamber is positioned at the reference depth in the phantom and the film at the entrance surface in the reference geometry. 3.7 Correction factors determination of GAF2 The response of the films was evaluated for varying field size, SSD, gantry angle wedge effects as well as tray effects. When carrying out the effect of any parameter on the film response, all other factors were kept constant as well as the irradiation time. The side of the field sizes was varied from 4cm to 24cm, the SSD from 75cm to 120 cm, while the gantry angle was varied from 0 to 90 and -90 at 5 intervals. The response of the ionization chamber was also evaluated for field size, SSD, gantry

84 angle wedge effects as well as tray effects. A normalized data of field size effects, SSD effects and gantry angle effects were determined using reference conditions of field size of 10 cm 10 cm; SSD of 100 cm and gantry angle of 0. Also, normalized data of the effects of wedges and trays were determined for field sizes ranging from 4cm 4cm to 18cm 18cm. Absence of wedges and trays were used as reference conditions. Correction factors accounting for the variations in response were determined as the ratio of the reading of an ionization chamber and the reading of the film for clinical irradiation situation normalized to the same ratio for the reference situation. The equations for the determination of correction factors are in equation 3 and 4. CF i = CF SSD, CF FS, CF ANGLE, CF WEDGE, CF TRAY... (3) CF =... (4) Where: CF is correction factor Ric is the reading of the ion chamber Rf is the reading of GAF2 The reference conditions are SSD of 100cm, Field size of 10 cm 10 cm. The variation in response due to different beam incident angles was measured for different gantry angles and normalized to the response measured when the central beam axis and the symmetry axis coincide. Field size correction factors were measured for square fields ranging from 4 cm 4 cm to 24 cm 24 cm, at the reference SSD of

85 100cm. SSD correction factors were measured for SSDs from 75 cm to 120 cm, at the reference field of 10 cm 10 cm. The SSD correction factors and field size correction factors were assumed to be independent. Wedge correction factors were measured at reference SSD, for square fields of side length ranging from 4 cm to 18 cm; beyond which the treatment machine gave error messages. The ratio of the signal of the ionization chamber to the film signal was normalized to the same ratio for the open beam (with the same field size). Tray correction factors depend mainly on field size. They were determined by repeating all measurements carried out for the field size correction factors, and normalizing the data to the reference situation of an open beam with the appropriate field size. 3.8 Entrance dose calculations After the calibration, correction factors as well as GAF2 reading have been determined, they are multiplied together to obtain the entrance dose as shown equation 5: D entrance = D s F cal, entrance ΠCF. (5) Where: D entry represents the entrance dose and is the dose at dmax D s represent GAF2 dose reading F cal, entrance represents entrance dose calibration factor ΠCF represents the product of correction factors used during clinical applications

86 To obtain the dose at any depth, D depth, D entry was multiplied with percentage depth dose (PDD) for Cobalt-60 energy. 3.9 Skin dose determination For all entrance dose measurements carried out using bare GAF2, the skin dose was also determined in order to estimate the amount of radiation being deposited in the skin tissue at a depth of 0.07 mm. The skin dose associated with radiotherapy is often of interest for clinical evaluation or examining the risk of late effects. Since the depth of active layer within the film is not at a depth of 0.07 mm, a skin dose calibration factor was necessary. To do this, an extrapolation technique was used. A single GAF2 was first irradiated for one minute. Depth of active layer for one film was 0.08 mm. Four films were then irradiated and the bottom film s measurement recorded. Depth of active layer of bottom film with three films on top was now 0.935mm. Optical density at depth of 0.07 mm within film was then calculated from knowledge of the optical density at a depth of 0.08 mm and mm. Since GAF2 is asymmetrical, attention was paid to the side of irradiation to minimize error propagation. The calculations of skin dose calibration factor, skin dose and percentage skin dose are shown as: Dskin = Df Fcal, skin... (6) Where Fcal, skin =

87 To determine the dose at depth of 0.07mm within the film, the following relation was used: 0.08 mm 0.03 OD mm 0.08 OD 0.07 mm X OD Therefore: X = OD Therefore: Fcal, skin = Dose to skin then becomes: Dskin = Df 0.98 % skin dose =... (7) Therefore % skin dose =. (8) But Fcal, entrance for bare film equals Therefore % skin dose = 100 =... (9) From equation 9, the percentage skin dose was found to be dependent mainly on irradiation field parameters such as field size, SSD, gantry angle, as well as the presence of beam modifiers such as trays and wedges

88 3.10 Preliminary skin dose assessment for both bare film and encapsulated film At reference irradiation conditions of field size of 10 cm 10 cm, SSD of 100cm and gantry angle of 0, the impact of both the bare film and the encapsulation cap on skin dose and % skin dose were then assessed and calculated for a prescribed dose of 200 cgy to d max at reference irradiation conditions as follows: a) Skin dose due to bare film on skin Assumption: the bare film was assumed to flash out with the surface of the skin such that its surface and that of the skin were the same. D skin = D f F cal, skin ; D f = 0.05 OD = cgy; F cal, skin = 0.98 Therefore: D skin = = cgy % skin dose = 23.7% b) Skin dose due to encapsulation cap on skin D skin = D f F cal, skin ; D f = 0.14 OD = cgy; Fcal, skin = 0.98 Therefore: Dskin = = cgy % skin dose = = 68.64%

89 CHAPTER FOUR RESULTS AND DISCUSSIONS In this chapter, the results from the study are presented and discussed. The study considered the design of an encapsulation cap for bare film for some dosimetry procedures, calibration factor determination as well as correction factors determination for both bare film and encapsulated film. The measured results were then compared with TPS or calculated data using the dose difference technique. 4.1 Evaluation of encapsulated GAF2 cap concepts After the two concepts for the encapsulation cap had been prototyped as shown in Appendix A, they were evaluated based on designer and Medical Physicist criteria. The criteria for comparison and selection of the best concept included cost of manufacture, aesthetics, total number of parts, duration for insertion of film into cap, duration for removal of film from cap, ease of manufacturing, weight and safety. All criteria were scored on a scale of 1 to 3, with 1 being worst, 2 being ok and 3 being best. Table 4.1 shows the scores for each concept when assessed on each design criteria

90 Table 4.1 Comparison of OPTION ONE and OPTION TWO OPTION OPTION ONE OPTION TWO CRITERIA Cost of manufacturing 1 3 Duration for insertion of 1 3 film into cap Number of parts 1 3 duration for removal of 2 3 film from cap ease of manufacturing 1 3 Durability 2 3 Safety 1 3 Aesthetics 3 2 TOTAL SCORE From the table, it was observed that the overall cost of manufacturing for Option Two was lower than Option One, hence more affordable. Also the duration for insertion and removal of film from Option Two was observed to be less than Option One, hence a reduction in the overall treatment time. The material used for the construction of the cap was Perspex (PMMA) since its effective atomic number closely matches that of soft tissues. The approximate thickness of the portion of the cap was 0.5 cm which corresponds to the depth of maximum dose d max. Furthermore, option one had a larger number of assembly parts increasing the overall assembly time and cost. Based on the above criteria, Option Two was chosen and used for post irradiation, calibration factors, correction factors, film behavior and response characteristics experiments

91 4.2 Post-Irradiation Optical Density Growth As shown in figure 4.1, it was observed that after exposing the film to ionization radiation, the optical density begin to increase steeply and gradually become asymptotic to the horizontal axis at longer hours after irradiation. Figure 4.1: Post-exposure optical density growth of GAF2 It was also observed that exposing the films to higher radiation doses increases the optical density and also increases the time needed for the response to become constant. While low doses of radiation (50 cgy) causes optical density growth which levels off quickly, high doses (200 cgy, 400 cgy, 800 cgy) of radiation were observed to produce optical density growth which took some time before remaining constant. From the results, it was also observed that after twenty four hours (1440 minutes) of exposure the optical densities of all the exposed doses of radiation were

92 observed to be constant. The optimum time to take measurements after exposing the film was concluded to be twenty four hours. 4.3 Calibration curve determination for GAF2 The absorbed doses obtained from the ionization chamber were plotted against the corresponding optical densities from the irradiated films obtained from densitometer readings. For the same batch number of films and irradiation conditions, depending on the type of densitometer used, the behaviour of the calibration curve in terms of optical density measurements may vary. Hence, the calibration curve needs to be reevaluated when a different densitometer or scanner is being used. The calibration curve with standard error bars obtained for the film is shown in figure 4.2. Figure 4.2: Calibration curve with standard error bars for GAF2-73 -

93 Figure 4.2 illustrates that as the net optical density (NOD) of GAF2 increases, its corresponding absorbed dose also increases. A polynomial of degree four represented the line of best fit that relates the net optical density to absorbed dose. The correlation coefficient was estimated to be The equation of the polynomial was: ; Where x represents the optical density and y the absorbed dose. The calibration curve was used to convert optical density measurements to absorbed dose in gray. 4.4 Behavior of GAF2 with varying irradiation conditions Both bare films and encapsulated films were used for the in vivo dosimetry studies. The bare films were used for both entrance dose and surface dose measurements while the encapsulated films were used for entrance dose measurements. The behaviour of both the bare films and encapsulated films to varying irradiation conditions are illustrated from figure 4.3 to figure 4.7. From figure 4.3, as the field size increases, the normalized mean optical density (ND) for bare films were observed to increase because as field size increases the contribution of secondary head scatter electrons also increases and since there is no medium above the bare film to stop them, most are trapped in the active layer in the film thereby increasing the resulting absorbed dose

94 Figure 4.3: A graph of normalized net optical density against field size of GAF2 with standard error bars For the encapsulated film, as the field size increases from 4 cm 4 cm to 6 cm 6 cm, the normalized optical density (ND) was observed to increase steeply and gradually became constant from 16 cm 16 cm to 26 cm 26 cm. The increase in ND in the encapsulated film for smaller films resulted from an increase in number of secondary electrons generated within the encapsulated material which increase with field size. Electrons generated outside the encapsulated film such as head scatter electrons are considered unwanted and must be stopped from reaching the film while electrons generated within the encapsulated film setup (encapsulation material and film) is of utmost importance and should actually be those that contribute to absorbed dose in the film. The inclusion of encapsulation material is necessary to stop head scatter electrons since the range of scattered electrons in soft tissues for Co-60 energy at reference conditions is approximately 0.5 cm. The effects of field size and head scatter electrons on bare films are more pronounced than encapsulated films. For

95 encapsulated films, the effects of field size and head scatter electrons can be ignored since it is negligible. As shown in figure 4.4, the normalized mean optical densities (ND) for encapsulated film were generally higher than those for bare film. Figure 4.4: A graph of normalized net optical density against SSD of GAF2 response with standard error bars As SSD increases, the dose rate of the Equinox 100 Co-60 machine decreases hence a reduction in the absorbed dose and optical density of the film with increasing SSD. The influence of head scatter electrons on encapsulated film is very negligible. Both encapsulated and bare films were severely influenced by SSD, hence the need for correction factors. SSD is therefore a very important parameter to consider when using GAF2 for in vivo dosimetry of external photon beam radiotherapy

96 From figure 4.5, the use of encapsulated GAF2 for in vivo dosimetry can be considered independent of gantry angle since the normalized mean optical density was observed to be unity for almost all the gantry angles. However, bare films behaviour was observed to be highly dependent on gantry angle with a maximum ND recorded at gantry angle 80 and -80. Figure 4.5: A graph of normalized net optical density against gantry angle of GAF2 response. In general, the bare film s ND was also observed to increase gradually from gantry angle 40 to 80 but decreased gradually from gantry angle 80 to 90. The irregularities in the behaviour of the bare film to gantry angles can be attributed to the irregularities in the number of contaminating head scatter electrons that interacted with the active layer in the film. Also, as the angle of incidence increases in general, the numbers of contaminating electrons that interacted with the active layer within the film also increased, hence the increase in net optical density of the film. These

97 irregularities were not seen in the behaviour of encapsulated films to gantry angles because the encapsulation material prevented the contaminating electrons from interacting with the film within. From figure 4.6, at reference field size of 10 cm 10 cm, as wedge angle increases, normalized net optical density was observed to decrease. Figure 4.6: A graph of normalized net optical density against wedge angle at reference field size of 10 cm 10 cm of GAF2 response with standard error bars. This is because as wedge angle increases, thickness of attenuating material increases, decreasing the amount of radiation leaving the wedge, hence the amount of radiation absorbed by the film decreases. The optical density of the encapsulated film was also observed to be higher than that of bare film because the film in the encapsulation material was at a depth of maximum dose

98 4.5 Correction factor values During actual treatment delivery, irradiation parameters may be different from those used in calibrating the film, hence the need to determine correction factors for these varying conditions. Correction factors determination considered in this research included that for SSD, field size, gantry angle, full tray and wedge. Table of values for the various correction factors for bare film is illustrated in Appendix H. The behaviors of encapsulated films were found to be less dependent on irradiation parameters than bare films. This means the need for more correction factors when using bare film for in vivo dosimetry. 4.6 Phantom studies Phantom studies carried out included entrance dose calibration factor determination as well as skin/ surface dose estimation. Analysis of the results obtained from the phantom studies are shown below Entrance dose calibration factor determination Entrance dose calibration factor for the use of GAF2 for in vivo dosimetry was done weekly for bare film due to the slightly independent nature of GAF2 to marginal dose rate differences as compared to ionization chamber readings. The Ionization Chamber was therefore always necessary to monitor the absolute dose at d max. A ratio of Ion Chamber reading at d max to film reading at surface of phantom, D surface was then determined to obtain the calibration factor. The reading

99 of the Ionization chamber was in centigray (cgy).the reading of the film was however converted from optical density to dose (cgy) so as to obtain a constant for the calibration factor. a) Bare film calibration factor determined on 31/05/2014 = = b) Encapsulated film calibration factor determined on 31/05/2014 Determination of the calibration factor was done to relate surface dose determination to dose at d max inside the phantom Entrance dose determination Entrance dose determination was carried out with bare film. Entrance dose provides an idea of the dose to d max within the patient or tissue equivalent phantom from surface dose measurements and is given by equation 8. D entrance = D surface F cal, entrance ΠCF (10) At reference irradiation conditions, calculated TPS doses were compared with GAF2 measured doses as shown in Appendix I. The maximum deviation was 2.6% and the minimum deviation was 0.1% which falls within the tolerance level of 5%. However, after varying all the irradiation parameters, the maximum percentage entrance dose difference was observed to be higher than the tolerance

100 level. It is therefore recommended to carry out further studies on the use and behavior of GAF2 before implementing it as an in vivo dosimeter by minimizing and being cognizant of the various levels of errors and uncertainties involved in the determination of the calibration curve, calibration factor and correction factors. To minimize these deviations the following should be undertaken; use highly sensitive 2D scanners, keep notice of side of film irradiation, monitor temperature and humidity changes and take readings after 48 hours of film irradiation Skin dose determination Graphs of how the respective irradiation field parameters affect percentage skin dose are shown in Appendix I. From the graphs, percentage skin dose was observed to increase with increasing field size as well as decreasing SSD though the effect of varying field size on % skin dose was observed to be of paramount importance. Also from the graphs, it was observed that as beam obliquity increases in general, the percentage skin dose also increases with the highest percentage skin dose of 40.2% recorded at gantry angle of 80. The effect of tray on % skin dose was also observed to increase slightly at field size increases above reference irradiation condition. Wedge angle effect on % skin dose was observed to be more pronounced at 60.The highest percentage skin dose of 45%for all the irradiation parameters was recorded for open beam geometry at field size of 24 cm 24 cm. As field size increases, the number of head scatter contaminating

101 electrons reaching the film also increases, hence the increase in percentage skin dose. From the calculations of % skin dose due to either the use of bare film or encapsulated film for in vivo dosimetry, it was observed that the dose to skin due to the use of encapsulated film was higher than the upper limit of normal % skin dose values. For cases where the skin is not the primary organ of treatment, this will mean a loss of skin sparing effect when using encapsulated film for in vivo dosimetry. For radiation therapy procedures for which skin sparing is of utmost importance the use of encapsulated film should not be considered. However, to minimize the effects of loss of skin sparing when using encapsulated film for in vivo dosimetry, the surface area of the encapsulation cap material should be made smaller than that used in this research. 4.7 Evaluation, comparison and selection of better option of GAF2 for in vivo dosimetry For purposes of this research, after analyzing the behavior of both bare film and encapsulated film, they were carefully assessed and the better option chosen for further analysis and experimentation, that is for skin dose, entrance dose as well as target dose determination. The criteria used for assessing both film configurations were skin sparing, degree of scatter from dosimeter, ease of placement on patient s skin, beam perturbation, probability to lead to under-dose of target volume, level of acceptability for use in radiotherapy centers, number and complexity of correction factors needed as well as cost of use. The various criteria used were ranked high or

102 low and positive (+) or negative (-); positive meaning good while negative means bad. The results of the evaluation process are shown in Table 4.2. Table 4.2 Comparison of bare film and encapsulated film for in vivo dosimetry CRITERIA OPTION BARE FILM ENCAPSULATED FILM Skin sparing High (+) Low (-) Degree of scatter from Low (+) High (-) dosimeter Ease of placement of High (+) Low (-) patient s skin Beam perturbation Low (+) High (-) Probability to lead to Low (+) High (-) under dose of target volume Level of acceptability High (+) Low (-) for use in radiotherapy centers number and complexity High (-) Low (+) of correction factors needed cost of use Low (+) High (-) TOTAL SCORE Seven (7) positives One (1) positive From table 4.2, it was observed that bare film had more positives than encapsulated film for use as an in vivo dosimeter. Despite encapsulated film having more negatives, some Medical physicists agreed to opt for that if the physical dimension can be worked on especially the surface area and thickness. Bare film was then chosen as the better option and used throughout this research to determine skin dose as well as entrance dose

103 4.8 Guidelines on in vivo dosimetry Measured dose with GAF2 might be different from the prescribed dose by the oncologist. It is therefore pertinent for the radiotherapy unit undertaking this mode of in vivo dosimetry protocol to define departmental guidelines and/ or procedures describing the immediate actions to be undertaken when the measured dose is out of tolerance and/ or action levels. The choice of tolerance/ action levels is very important since they will in practice determine the number of errors detected and will influence the associated workload to implement or maintain in vivo entrance dose measurements at the departmental level. If a too broad tolerance window is chosen, some causes of erroneous treatment delivery may not be detected (for instance a wedge 30 instead of a wedge 15, presence of tray etc.). If the tolerance window is too small, a large number of measurements will have to be repeated (due to e.g. inherent statistical fluctuations or a too critical positioning of GAF2 in e.g. wedged beams). For most radiotherapy units, the tolerance level is set at 5% and is often assumed to be the same as action level. Any deviations greater than 5% must therefore be investigated. Guideline flowchart including actions undertaken at different levels is shown in appendix L. Other questions regarding staffing and management of personnel should also be clarified such as: 1) Who is the contact person for measurements out of tolerance or action levels? 2) If a second measurement is requested should it be performed in the presence of the physicist?

104 4.9 Recording of in vivo dosimetry For purposes of this research, a special in vivo dosimetry chart was designed with Microsoft excel and is to be made available at radiotherapy centers interested in using GAF2 as an in vivo dosimeter. The chart was designed to include sufficient information such as the date of measurement, the type of field, the anatomical location and so on. The chart was programmed with appropriate formulas to automatically generate all the values of the correction factors as well as entrance dose and surface dose values. A sample of the chart is shown in appendix K. Guideline flowchart including actions undertaken at different levels is shown in appendix L, figure L

105 CHAPTER FIVE Conclusion and recommendation This chapter presents the conclusion drawn from the results and the recommendations to stakeholder institutions on how to improve their overall quality in external photon beam radiotherapy in the near future through the use GAF2 as an in vivo dosimeter. 5.1 Conclusion The response of both bare film and encapsulated film was determined for varying irradiation conditions such as field size, SSD, gantry angle, open field as well as wedged field. The response of the encapsulated film was observed to be nearly independent of irradiation conditions while bare film showed strong dependence on irradiation conditions. For both bare film and encapsulated film, SSD was also observed to be the main irradiation parameter that significantly affected their behaviour. From the entrance doses calculated and measured in tissue equivalent solid slabs, the percentage dose deviation was found to be in the acceptable range of 5% when considering reference irradiation conditions as well as variation of only one of the irradiation parameters. The lowest percentage dose difference was found to be 0.1% while the highest percentage dose difference was found to 2.6%. However, by varying all the irradiation parameters for open field, the maximum percentage entrance dose difference of 11.2% was observed to be higher than the tolerance level of 5%

106 The percentage skin dose measured with bare film at reference conditions was 23.7% while that measured with encapsulated cap on top of the skin was 68.64%. Percentage skin dose was also observed to increase significantly with increasing field size although a slight increase was also observed with decreasing SSD. At large gantry angles (70, 75, 80, etc), the percentage skin dose was also observed to increase significantly. The inclusion of beam modifiers such as trays and wedges also resulted in a marginal increase in percentage skin dose. The percentage skin dose was however observed to be significantly higher at wedge angle of 60 due to the increase in beam obliquity. 5.2 Recommendations The following are the recommendations made to the various stakeholders at the National Centre for Radiotherapy and Nuclear Medicine of the Korle-Bu Teaching Hospital (KBTH) Oncology and research centers 2D scanners which are very sensitive to minute changes in optical densities should be made available so as to ensure faster 2D dose assessment and monitoring as well as minimize the overall error margins involved in the use of GAF2 for in vivo dosimetry

107 5.2.2 Regulatory Authority It is recommended that regulators dealing with radiotherapy facilities through their random visits employ this method of in vivo dosimetry to ensure that patients are not under-treated or over-treated during radiation treatment delivery Further Research Work It is therefore recommended to carry out further studies on the use of GAF2 before being implemented as an in vivo dosimeter at NCRNM, KBTH, by minimizing and being cognizant of the various levels of errors and uncertainties involved in the determination of the calibration curve, calibration factor and correction factors. This will go a long way to improve the overall treatment quality delivery during external beam radiotherapy

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116 APPENDIX A AutoCAD Design of encapsulation cap Option One Figure A1: 2D dimension of buildup material above film * All measurements are in centimeters Figure A2: Buildup cap above film

117 Figure A3: 2D dimension of buildup cap below film Figure A4: 3D buildup cap below film

118 Figure A5: 3D assembly drawing of buildup cap with film slot Figure A6: 3D assembly drawing of buildup cap with film being inserted

119 Figure A8: 3D assembly drawing of fully inserted film in buildup cap Figure A9: 2D dimension of film handler *This is a device designed to aid the insertion and removal of the film from the cap

120 Figure A10: 3D assembly drawing of film handler Figure A11: 3D assembly drawing showing film being pulled with film handler

121 Figure A12: Closer view of film handler, film, buildup cap assembly Figure A13: Physically constructed Option One

122 APPENDIX B Design of encapsulation cap OPTION TWO Figure B1: 2D dimension of buildup cap Figure B2: 2D dimension of buildup cap locker

123 Figure B3: 3D Part design of buildup cap assembly Figure B4: 3D Assembly drawing showing locker in position

124 Figure B5: 3D Assembly drawing showing GAF2 in position Figure B6: 3D Assembly drawing showing locker in position

125 Figure B7: 3D Assembly drawing showing fully encapsulated GAF2 Figure B8: Physically constructed Option Two