Small-segment intensity modulated radiation therapy dosimetry with various ion detectors and Gafchromic EBT2 film

Transcription

1 University of Wollongong Research Online University of Wollongong Thesis Collection University of Wollongong Thesis Collections 2012 Small-segment intensity modulated radiation therapy dosimetry with various ion detectors and Gafchromic EBT2 film Yunfei Hu University of Wollongong Recommended Citation Hu, Yunfei, Small-segment intensity modulated radiation therapy dosimetry with various ion detectors and Gafchromic EBT2 film, Master of Engineering - Research thesis, Faculty of Engineering, University of Wollongong, Research Online is the open access institutional repository for the University of Wollongong. For further information contact the UOW Library: research-pubs@uow.edu.au

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3 Small-Segment Intensity Modulated Radiation Therapy Dosimetry with Various Ion Detectors and Gafchromic EBT2 Film By Yunfei Hu Submitted as part fulfillment of the requirements for the award of the degree of Master of Research from University of Wollongong Faculty of Engineering 1

4 Abstract Intensity modulated radiation therapy (IMRT) has been widely accepted in clinical applications. Compared to traditional 3D conformal radiation therapy (3D-CRT) techniques, IMRT can deliver the prescribed dose to the target while reducing the dose delivered to surrounding normal tissue. An IMRT treatment composes multiple field segments (typically 100 to 120); these segments usually involve the use of very small fields measuring as little as 1cm x 1cm. Unfortunately, the accuracy of the dosimetry of such small fields has been a concern to medical physicists because different ionisation dosimetry techniques may yield different results mainly due to different detector volumes. This causes a lack of confidence in IMRT quality assurance (QA) results. This thesis compares and assesses the performance of different dosimeters in small field measurements. The detectors include ionisation chambers of different volumes, a diode, and Gafchromic EBT2 film. Tests include verification of the basic properties of the detectors. Their performance is also compared with a radiation therapy planning system (RTPS). The focus of the project is the suitability of these detectors for IMRT QA. This research also develops and validates a simple and novel method of Gafchromic EBT2 film analysis for IMRT QA using a grey-scale value to calibrate the film as a viable option compared to the conventional colour channel analysis method. Results indicate that small-volume detectors have a higher spatial resolution, which is particularly useful in measuring high-dose gradients, such as the buildup region of PDDs or the penumbra of small-field profiles. Among all the detectors the Gafchromic EBT2 film has the highest spatial resolution, and its measurement is potentially the closest to the reading from a virtual zero-volume detector. However, these detectors usually have poor precision, as their noise is relatively high, especially the Gafchromic film. The poor precision will sometimes reduce the accuracy of the measurement. On the other hand, detectors with larger volumes have worse spatial resolution but a relatively more precise signal. When the detector is used to measure a high dose gradient region, its poor spatial resolution will certainly influence the measurement accuracy. This is usually called the volume-averaging effect. Other factors, such as the detector s non-tissue equivalence or lack of electron equilibrium in the field, may also affect the accuracy of the measurement. When scanned as a 16-bit grey-scale value (GSV) image, the Gafchromic EBT2 film can provide a very linear GSV-to-dose response in a dose range of 0 to 6Gy. This means that for a relative measurement, conversion to dose using a log or polynomial relation is not required and the GSV pixel value can be linearly related to dose in this dose range. Together with its other advantages, such as extremely high spatial resolution, planar geometry and water equivalence, the Gafchromic EBT2 film method is a suitable detector for small-segment IMRT QA. However, due to its poor 2

5 precision and previously reported batch calibration variation it is not recommended for absolute dose determination. Among the other detectors tested, the CC04 ionisation chamber was most suitable for small-segment IMRT dose determination because it provided the best balance between a reasonably high spatial resolution (to maintain a satisfactory accuracy) and sufficient volume for charge collection (to give relatively high precision). 3

6 Acknowledgements I would like to express my sincere gratitude to everyone who has helped in the writing of this thesis, with my largest appreciation going to: Professor Peter Metcalfe, my university supervisor in the later parts of the project, who has given me valuable advice about journal references relevant to the project and the use of proper scientific terms, and who has encouraged me to further investigate the Gaf TM film analysis method. Associate Professor Yang Wang, my clinics supervisor, whose extensive knowledge in medical physics has greatly expanded my view. Without his help the research would never have been successfully completed. Dr George Takacs, who contributed to project supervision in the early parts of the project. My former TEAP supervisor Mr Guangli Song, who helped me establish the fundamentals in medical physics. Associate Professor Martin Butson, who gave valuable advice on interpretation of the Gafchromic film experimental results. The physics group of Genesis Cancer Care Pty Ltd, who assisted me during the work so that I could finish writing the thesis. Laura E. Goodin, who provided editorial support. All others who have assisted. Your efforts to help me with this thesis will never be forgotten. 4

7 Contents Chapter 1: Introduction.12 Chapter 2: Literature Review Radiotherapy and Medical Linear Accelerator Multileaf Collimator DCRT and IMRT History of Small-Field Dosimetry Dosimeters Ionisation Chambers Semiconductor Diodes Radiochromic Films Thermoluminescene Dosimeters Diamond Detectors.35 Chapter 3: Materials and Methods General Experimental Design List of Equipment CC CC CC Diode Gafchromic EBT2 Film Electrometer Film Scanner RODOMS OmniPro-Accept Linear Accelerators (Linacs) Ion-Chamber Cylindrical Phantom Film Cylindrical Phantom Detailed Experiment Procedures Noise Measurements Gafchromic EBT2 Film Dose Response Calibration Output-Factor Measurements PDD Measurements

8 3.3.5 Off-Central-Axis Profile Measurements Clinical IMRT Case Comparison Error-Bar Determination Gamma-Function Assessment.68 Chapter 4: Results and Discussions Signal-to-Noise Ratio Measurements Calibration of Gafchromic EBT2 Film Using a Novel Method Output-Factor Measurements Percentage Depth Dose Measurements Off-Central-Axis Profile Measurements Clinical IMRT QA Comparison Planning Standard Output Ionisation Chamber Measurements (a) CC (b) CC (c) CC Gafchromic EBT2 Film Measurements. 106 Chapter 5: Conclusion 115 References

9 Abbreviations 3D-CRT Three-Dimensional Conformal Radiotherapy CAX Central Axis cgy CentiGray DPI Dots per Inch DTA Distance to Agreement G-T Gantry-Target Gy Gray H&N Head and Neck IMRT Intensity Modulated Radiotherapy ISP International Specialty Products L-R Left Right LINAC Linear Accelerator MM Metal-Insulator-Metal MLC Multileaf Collimator MU Monitor Unit nc NanoCoulombs NCRP National Council on Radiation Protection and Measurements OAR/OCR Off Central-axis Ratio PDD Percentage Depth Dose PMMA Poly(methyl methacrylate) POI Point of Interest PTV Planning Target Volume QA Quality Assurance RTPS Radiation Treatment Planning System SAD Source Axis Distance SSD Source Surface Distance TL Thermoluminescent TLD Thermoluminescent Detector TRS 398 Technical Reports Series No

10 List of Tables Table 4.1.1: Readings for 10 repeated measurements at both field sizes 67 Table 4.2.1: Dose response of Gafchromic EBT2 film at different channels.74 Table 4.2.2: Dose response of Gafchromic EBT2 film with different image adjustments 75 Table 4.2.3: Dose response of Gafchromic EBT2 film with different ranges...78 Table 4.3.1: Output factors measured by different ionisation chambers and diode.82 Table 4.3.2: Average output factors and standard deviations measured with Gafchromic film..83 Table 4.3.3: Output factors measured with different detectors 83 Table 4.5.1: d 80 and d 20 values of the 1cm x 1cm OCR curves collected by different dosimeters..94 Table 4.5.2: d 80 and d 20 values of the 10cm x 10cm OCR curves collected by different dosimeters..96 Table : Doses of interest point exported from XiO Table : IMRT point-dose comparison-cc Table : IMRT point-dose comparison-cc Table : IMRT point-dose comparison-cc Table : IMRT point-dose comparison summary..104 Table : IMRT point-dose comparison beam 4 excluded.105 Table : Point-dose comparison in relative mode.110 8

11 List of Figures Figure 2.2.1: Siemens 160MLC Multileaf collimator.16 Figure 2.2.2: Siemens Artiste 160MLC system illustration...17 Figure : Cylindrical IMRT phantom for ionisation chamber Figure : Cylindrical IMRT phantom for ionisation chamber, longitudinal view..44 Figure : Cylindrical IMRT phantom for ionisation chamber, saggital view..44 Figure : Cylindrical phantom and its chamber holders for different types of ionisation chambers..45 Figure : Cylindrical IMRT phantom for Gafchromic film.46 Figure : Cylindrical IMRT phantom for Gafchromic film, longitudinal view..46 Figure : Top and bottom half of the cylindrical IMRT phantom for Gafchromic film.. 46 Figure : Chamber-holder phantom slice for CC13 47 Figure : Alignment of the ionisation chamber to the iso-centre of the linac with the aid of the field-light crosshair 48 Figure : Alignment of the film stripe to the iso-centre of the linac with the aid of the field-light crosshair..49 Figure : Epson Scan software setup for Gafchromic EBT2 film scan Figure 4.1.1: Noise illustration at 1cm x 1cm field. 68 Figure 4.1.2: Noise illustration at 10cm x 10cm field...68 Figure 4.1.3: Magnified inter-umbra region of the OCR curves with different dosimeters under 10cm x 10cm field without the application of a smoothing function 70 Figure 4.2.1: 16-bit grey-scale Value image..71 9

12 Figure 4.2.2: 48-bit colour image.72 Figure 4.2.3: Red channel image...72 Figure 4.2.4: Green channel image 73 Figure 4.2.5: Blue channel image 73 Figure 4.2.6: Gafchromic film PV-dose response with different channels.74 Figure 4.2.7: Gafchromic film GSV PV-dose response with different image adjustments 76 Figure 4.2.8: Low-dose-range film from 0cGy to 55cGy in steps of 5cGy 77 Figure 4.2.9: Middle-dose-range film from 0cGy to 550cGy in steps of 50cGy 77 Figure : High-dose-range film from 0cGy to 1100cGy in steps of 100cGy. 78 Figure : Film dose range response in low-dose range..79 Figure : Film dose range response in middle-dose range 79 Figure : Film dose range response in high-dose range 80 Figure : Overall dose response of the Gafchromic film 81 Figure 4.3.1: JPEG image of film output factors, irradiated on 27/7/10.82 Figure 4.3.2: Output-factor curves measured with different dosimeters...84 Figure 4.3.3: Output-factor curves measured with different dosimeters normalised to 4cm x 4cm field..85 Figure 4.4.1: PDD curves measured with different dosimeters under 1cm x 1cm field..87 Figure 4.4.2: Build-up region of the PDD curves measured with different dosimeters under 1cm x 1cm field..88 Figure 4.4.3: Dose fall-off region of the PDD curves measured with different dosimeters under 1cm x 1cm field..90 Figure 4.4.4: PDD curves measured with different dosimeters under 10cm x 10cm 10

13 field..90 Figure 4.4.5: Build-up region of the PDD curves measured with different dosimeters under 10cm x 10cm field..91 Figure 4.4.6: Dose fall-off region of the PDD curves measured with different dosimeters under 10cm x 10cm field..92 Figure 4.5.1: OCR curves measured with different dosimeters under 1cm x 1cm field..93 Figure 4.5.2: Penumbra region of the OCR curves measured with different dosimeters 1cm x 1cm field..93 Figure 4.5.3: OCR curves measured with different dosimeters under 10cm x 10cm field..95 Figure 4.5.4: Penumbra region of the OCR curves measured with different dosimeters under 10cm x 10cm field..96 Figure 4.5.5: Virtual zero-volume detector profile extrapolation for 1cm x 1cm field...97 Figure 4.5.6: Virtual zero-volume detector profile extrapolation for 10cm x 10cm field..97 Figure : Virtual film from XiO Figure : IMRT Gafchromic film Figure : Reference Gafchromic film Figure : Screenshot of the relative dose analysis mode of RODOMS. 109 Figure : Screenshot of the gamma-value function results from RODOMS.111 Figure : Gamma-function results with the measurement film shifted to the left by 0.5mm..112 Fig : Gamma-function results with the measurement film shifted to the left by 1.0mm

14 Chapter 1: Introduction One important goal of quality radiation therapy (RT) is to accurately deliver the prescribed dose to the target volume, while minimising the radiation damage to the surrounding healthy tissues. Intensity modulated radiation therapy (IMRT), whose radiation fields are composed of multiple segments formed by multileaf collimators, was developed to achieve this goal. In an IMRT treatment, field shaping in 3D conformal radiotherapy ensures that the radiation fields are irregularly shaped but of uniform intensity; additionally, the intensity of the beam is modulated through the use of multileaf collimators (MLC) (Podgorsak, 2005). IMRT s many advantages in dose distribution and dose sparing have made it a mainstream radiotherapeutic technique. However, this technology includes the use of segments with small and irregular shapes. Consequently, the accuracy and the precision in the dosimetry with segment sizes less than 3cm x 3cm is very important for both commissioning data collection and clinical routine quality assurance (QA) for IMRT, and thus for patient s IMRT treatment delivery. In radiotherapy treatment planning system beam modeling, a treatment beam is modeled to match the measured beam data which are collected via ionisation measurements in a water phantom during commissioning. Ionisation measurements in IMRT data collection commonly use waterproof ionisation chambers with volumes from 0.01cc to 0.13cc in a water phantom to collect percentage depth dose and off-central-axis profile data at different depths. A 0.6cc Farmer-type chamber is often used to carry out output monitor unit calibration and output-factor measurements. For conventional radiation therapy planning systems (RTPS), beam data are usually collected for field sizes from 3cm x 3cm to 40cm x 40 cm. For RTPS that includes IMRT planning, data collection usually includes measurements with field sizes as small as 1cm x 1cm. Under such small fields the measured results from different dosimeters vary significantly due to different designs of the dosimeters (Low et al., 2003). Possible uncertainties in the small-beam measurements resulting from the limitations of existing dosimeters reduce confidence in both the commissioning results and the QA results for IMRT. To improve the accuracy and the precision in IMRT beam model data commissioning in treatment planning systems, and to develop an efficient and consistent clinical QA process for IMRT plan verification, small-beam data collected by different dosimeters should be tested and compared. The results should be evaluated in all aspects such as signal reliability, noise level, dose response, spatial resolution, and precision and accuracy in the measurement of percentage depth dose (PDD), profile and output factors. With the trend away from silver-halide-based dosimetry film, a relatively new 12

15 self-developing film, the Gafchomic radiochromic film, has become a useful alternative for measuring two-dimensional dose maps in IMRT fields. Gafchromic EBT2 film, the second generation of Gafchromic film, offers various options for its calibration. This thesis presents a simple and novel calibration method that is viable over a useful dose range. This method is compared with the colour-channel calibrations currently used by most Gafchromic film users. The new method is then successfully implemented in IMRT QA procedures. In summary, the main aims of this project are: To compare qualitatively the basic properties (such as metrics, noise and precision) of different-volume ionisation chambers, diodes, and Gafchromic films. To develop a suitable and simple calibration method for Gafchromic EBT2 film. To use ionisation chambers, diodes and Gafchromic films to collect RTPS-required data such as depth dose curves and output factors, and compare their performances. To select and implement these dosimeters for IMRT QA. 13

16 Chapter 2: Literature Review The literature review aims at providing the basic information relevant to this study, including the history of radiotherapy and the medical linear accelerator (linac), the development of IMRT delivery, progress in current dosimetry techniques, and important research outcomes in small-field dosimetry data collection. 2.1 Radiotherapy and the Medical Linear Accelerator Radiotherapy is the medical use of ionising radiation as part of cancer treatment to control malignant cells. Radiotherapy may be used for curative or adjuvant treatment. It is also commonly to combined with surgery, chemotherapy, hormone therapy or some mixture of the three. Most common cancer types can be treated with radiotherapy. The Australia Health Technology Advisory Committee states that overall, 52% of people with newly diagnosed cancer are likely to benefit from radiotherapy (Delaney et al., 2006). The precise treatment intent (curative, adjuvant, neoadjuvant, therapeutic or palliative) will depend on the tumour type, location, and stage, as well as the general health of the patient. In radiotherapy, to spare normal tissues, shaped radiation beams are aimed from several gantry angles to intersect at the tumour, providing a much larger absorbed dose in the tumour than in the surrounding healthy tissue. The linac is currently the most commonly used device for treating deep-seated tumours. The linac provides high-energy X-rays with penetrating characteristics to cover the target volume while minimising the dose to the organs at risk. The evolution of the linac was a direct result of radar-development work that culminated in the production of microwave generators in the form of magnetrons and klystrons (Metcalfe et al., 2007). The first clinical use of the linac was in September 1953 in England, but it only became the best choice for radiation therapy departments in the 1980s (Thwaites, 2006). Since then it has been continuously developed and refined, and is now able to create small irregular segmented fields via the use of MLCs. Some advances have also been made to compensate for patient motion and setup variation with the use of accurate image guidance. With the progress in linac and ancillary technology, radiotherapy treatment planning has more choices of techniques, such as IMRT, to improve dose delivery and target conformation than with older non-mlc machines, which permitted the targeting of only simple, large, open fields with limited custom blocking. The objective of these new techniques is to enable better clinical outcomes through a better conformity, but their complexity means they also require more comprehensive data collection as well as higher accuracy and precision in the dosimetry during the commissioning of a linac. This is the initiative of this thesis: to generate a proper protocol for the commissioning and quality assurance of a linac to ensure higher accuracy and precision without sacrificing the efficiency of the measurement, so that the advantages of new 14

17 treatment modalities such as IMRT will not be compromised due to the limitations of measurement. 2.2 Multi-Leaf Collimator 15

18 A multileaf collimator (MLC) for photon beams consists of a large number of leaves that can be driven automatically, independent of each other, to generate a field of any shape. The thickness of the leaves along the beam direction is sufficient to provide acceptably low beam transmission, usually less than 1% (Khan, 1992). Figure 2.2.1: Siemens 160MLC Multileaf Collimator (courtesy of Siemens Medical) The linac used for all the data collection for this thesis is the Siemens Artiste Solution, so it is worthwhile to discuss its MLC system in this section. The Siemens Artiste linac uses the 160 MLC system. It provides 160 leaves, 80 on each side. The thickness of each leaf projects to only 5mm at the isocentre. This is capable of providing accurate conformity to the actual target shape for homogeneous coverage. The leaf is made of tungsten, with a height of 95mm. Transmission through leaves is claimed to be less than 0.75%, and between leaves less than 1.5% (Siemens 160 MLC manufacturer s specifications). Unlike Varian and Elekta linacs, Siemens has chosen to mount their MLCs as a replacement of the lower-standard jaw system. This geometry gives an intermediate leaf-width dimension, relative to the other two manufacturers, for the same projected size at the iso-centre. This positioning creates a favourable geometry for the use of arcing trajectories so that the leading edge of each leaf follows beam divergence. The other MLCs do not employ this level of sophistication, instead using rounded leaf ends with a linear trajectory (Galvin et al., 1999). The Siemens MLC is currently configured as a lower-jaw replacement, where the lower jaws can be split into a set of leaves, rather than a tertiary device. As a result, the inter-play scatter effects with an extra jaw (collimator) are not an issue. On the other hand, when the MLC is used as a tertiary device just below the level of the upper and the lower jaws, such as in the configuration of the Varian linac, the presence of an extra jaw set compared to the Siemens configuration will introduce extra scatter in the linac head and reduce the collision-free zone (Boyer et al., 2001). Thus, small-field measurements on this Siemens linac MLC system represent 16

19 equivalent dosimetry results as small-segment measurements. Figure gives a simple illustration of the Siemens MLC system. Fig 2.2.2: Siemens Artiste 160 MLC system illustration (courtesy of Siemens Medical) 2.3 3DCRT and IMRT 17

20 Three-dimensional conformal radiation therapy (3DCRT) was developed in the early 1990s to improve dose conformation so as to better spare normal tissues and increase treatment quality. 3DCRT aims to exploit the potential biological improvements due to better spatial localisation of the high-dose volume (Webb 1998c). 3DCRT may be broadly divided into two classes of technique: those that employ geometric field-shaping alone and those that also modulate the intensity of fluence across the geometrically shaped field. IMRT is regarded as a sub-set of 3DCRT. The tenet is that by sparing more volume of OARs the dose to the planning target volume (PTV) can be escalated. Clinical trials show that an improved conformation correlates with an improved clinical outcome. For example, in prostate therapy, conformal blocking leads to less radiation-induced proctitis and bleeding (Webb, 2001c). 3DCRT increases the conformity of the prescription dose level to the defined target volume, especially when sensitive structures are adjacent to or near the high-dose regions. However, until the recent availability of intensity-modulated radiotherapy (IMRT), high dose conformity to all but very simple target volumes was not always achievable. This was especially true near concave surfaces of the target. If there is a sensitive structure within that concavity, not being able to keep high doses off this region may incur unacceptable complications, and will certainly limit the dose that can be delivered. Such situations include, for example, a tumour that is partially surrounding the spinal cord or brainstem. The situation is further exacerbated if previous irradiation has been given, in which case the availability of IMRT can mean the difference between radiotherapy being an option or not. The ability of IMRT to conform the dose around concave re-entrant structures is the key difference between IMRT and 3DCRT. Intensity-modulated radiation therapy, or IMRT, is an advanced form of 3DCRT. It must be mentioned that 3DCRT is conformal mostly in terms of the way beam apertures are shaped via a beam s-eye view (BEV) of the target, but even the use of a very large number of such conformal beams does not imply that the dose will conformally surround the target volume. The problem is that a BEV of the anatomy does not give any indication of the 3D shape of a structure, since all that is seen is a projection of the shape from each angle. Unlike 3DCRT, IMRT generates a complete 3D view of the object, as opposed to the reconstruction of a series of 2D projections, to create a true conformal dose distribution including convex targets for planning purposes. Its development and adoption into widespread clinical practice represents one of the most significant advances in radiotherapy treatment planning and delivery technology (Metcalfe et al., 2007). Although IMRT is a relatively new treatment technology, intensity-modulated distributions, starting with a simple block or a physical wedge to produce a gradient of intensity across a field, have been produced for many decades. Metal compensators 18

21 serve as another good example. This delivery technology can be deemed as the origin of IMRT, as it is also based on the concept of optimising intensity patterns with photons. The modern history of IMRT spans little more than a decade, going back to concepts first introduced by Brahme (1988) for inverse planning. Inverse planning creates the best set of IMBs (intensity-modulated beams) from the ideal dose prescription (or a statement of dose objectives) to provide a dose solution. Inverse planning dose not usually determine a method of treatment delivery, although constraints on the available delivery apparatus could be factored into these calculations. Despite its relatively short history of clinical application, IMRT is now widely used due to its many advantages over conventional 3DCRT. First, clinicians require concave dose distributions in approximately 30% of clinical cases (Webb, 2001c), and these cannot be achieved without IMRT. Inverse-planning techniques that determine IMRT distributions, which greatly ease the generation of an IMRT plan, have become commercially available for clinical use. Finally, the advent of image fusion of CT, MRI, SPECT and PET make possible a more accurate delineation of target and normal structures. IMRT techniques are usually delivered with photon beams. The selection of the photon-beam energy usually depends on the specific case to be treated. In most clinical situations, a beam-energy arrangement similar to that used for 3D-CRT is usually sufficient for IMRT to achieve a superior dose distribution (Hunt et al., 2002). However, Welsh et al. (2007) pointed out that despite their skin-sparing potential, high-energy photons (>10MV) are not recommended in IMRT treatment. The major concern is that some planning systems may not effectively model the beam narrowing at depth observed with high-energy photons and small radiation beamlets encountered in IMRT. This phenomenon is caused by lateral beam degradation due to penumbra widening; when very small fields are used, there is greater loss of electron equilibrium laterally with higher-energy photons (Welsh et al., 2007). The other concern about using high energies for IMRT is the neutrons present above 10MV (NCRP-151). Most IMRT beams also use multiple beam angles; with increasing beam directions and lower beam weights for each beam, it is understood that the extra penetration at high energies becomes less of an advantage. Because IMRT often employs high MUs, the extra MLC leaf leakage at high energies also needs to be accounted for. While none of these issues alone preclude the use of IMRT at higher X-ray beam energies, most centres have invested considerable time and effort into commissioning their IMRT beams on the 6MV X-ray beam energy (Health et al., 2004; Li et al., 2009). Currently at the author s clinic, 6MV photon is the most often used beam energy for IMRT treatment, while 10MV photon is sometimes used for the treatment of deep-seated tumours such as prostate cancer. It still remains a significant concern that an IMRT treatment s small, high-dose volumes with large low-dose areas can result in less target volume coverage and 19

22 more damage to healthy tissue than a traditional, less conformal approach. Randall (2006) has stated that the uncertainty in the instantaneous position of the tumour during treatment can lead to non-malignant volumes receiving the high dose intended for the tumour volume. Interfraction as well as intrafraction target movement relative to reference landmarks, coupled with setup errors and other inaccuracies, add to this uncertainty. The traditional standard approach has been to add margins to the target volume, usually at the expense of most of the potential benefits of the more precise treatment-delivery techniques. Recently, though, the development of image guidance, for example cone beam computed tomography (CBCT) has enabled visualisation of some critical anatomy just prior to patient treatment on a linac (Podgorsak, 2005). It is now becoming a common way to avoid geographical misses in RT. While its development is a key to the success of IMRT, and they are complementary, image guidance radiotherapy is not the focus of this thesis. On the other hand, the accuracy of treatment delivery can be improved by accurate commissioning of both the linac and the planning system. This can be helped by the accurate and precise measurements of the commissioning data collection, which demands the selection of proper dosimetry tools and guidance about their appropriate use. This is the aim of this thesis. 2.4 History of Small-Field Dosimetry 20

23 As early as 1994, Beddar et al. (1994) reported on some research in dose perturbation caused by diodes in small field dosimetry. Although at that time detection systems such as ionisation chambers were already commercially available for photon-beam dosimetry, diodes were one of the most frequently used techniques to verify dose distributions for the small fields used in radiosurgery. Beddar et al. demonstrated the problems with using diodes for stereotactic field profile measurements as well as the effect of diode orientation on measured absorbed dose profiles in a water phantom, as absorbed dose measured with diodes in penumbral regions can exhibit perturbations. The perturbations came from the intrinsic detector material and the design of the diode, and could be significant for the small photon fields used in stereotactic radiosurgery. To quantify the perturbation, a p-type Si photon diode and a p-type Si electron diode were used for these measurements (Scanditronix). The sensitive volume of these detectors was specified to be about 0.2 to 0.3mm 3, with a diameter of 2.5mm and a thickness of 60μm. Measurements were performed using 1, 2, and 3cm-diameter circular field sizes as produced by their corresponding radiosurgery cone. The authors found that for each cone the absorbed-dose profiles were differed depending on the diode orientation and on the type of diode used. Moreover, the results showed that the determination of the position of the 90% isodose line also depended on diode orientation, which could result in a discrepancy of as much as 0.6mm for the 1.0cm stereotactic cone; this was quite significant for stereotactic radiosurgery. Similar discrepancies were reported for larger-diameter cone sizes as well. The authors decided that the anomalies were due to the non-water equivalence of the diode and the non-uniformity of the diode detector envelope surrounding the sensitive volume of the diode, which would raise some concern about using diodes for the measurement of penumbra for small, high-energy photon beams. They concluded that diodes could cause dose perturbations that may be significant for the small fields used with precision radiotherapy techniques. Distortion of profiles could rise when using diodes with their axes parallel to the direction of scanning and perpendicular to the central axis of the beam. Also, for small radiation fields the measurements of relative dose factors using diodes might be affected by the diodes response, although this uncertainty was not as significant as the spatial uncertainties. Lee et al. (2006) explored the application of radiochromic films in small-field dosimetry. They applied a high-resolution radiochromic film dosimetry method to verify a small field stereotactic radiosurgery (SRS) plan. They claimed that methods using micro-sized detectors might introduce a perturbation to the radiation field large enough to prevent an accurate result and render it impossible for the detectors resolve the steep dose gradients commonly encountered in SRS and IMRT plans. 21

24 Although silver-halide-type films might solve this problem, they were non-tissue-equivalent and had differential responses to low-energy radiation. They also required a film processor. Lee et al. preferred radiochromic films suitable for medical dosimetry work for their many desirable features: they were self-developing and daylight-handling, and had near tissue-equivalence. Another attractive property that made RCF a candidate for high-resolution 2D dose measurement was its intrinsically high spatial resolution, although this was unfortunately limited by the resolution of the reading densitometers. The authors were also concerned about the sensitivity of the RCF to the wavelength of the analysing light used in the densitometer. To solve this problem, the authors designed a novel microdensitometer with an improved film sensitivity and micrometer resolution, making the RCF a practical medium for high-resolution 2D dose measurement. Their results showed that the difference between the measured and the planned iso-centre doses for a 5mm-diameter collimator was about 2% (requirement within 5%), and the spatial deviation between the measured and planned isodose lines was less than 0.4mm. The 80% isodose line, normally taken as the prescription isodose line in SRS, had a maximum discrepancy of only 0.2mm. The authors concluded that with the aid of an advanced densitometer, the radiochromic films could provide acceptable dosimetric and geometric accuracies for small-field dosimetry. A more comprehensive comparison was performed by Pappas et al. (2008). The authors found that small photon beam measurements were problematic, mainly due to the presence of high dose gradients and the nonexistence of lateral electronic equilibrium in narrow photon beams. They asserted that the dosimeters used for such measurements should (i) be tissue-equivalent and not perturb the radiation beam, (ii) exhibit energy, dose rate, and directional independence of response, (iii) have a small sensitive volume and the ability to record high spatial resolution measurements, and (iv) overcome the positioning problems usually present in small-field dosimetry. However, no single detector fulfilled all the mentioned requirements; therefore they proposed to use and compare several types of dosimeters for small-field dosimetry. The aim of their work was to use the gamma-index concept to both quantitatively and qualitatively analyse and investigate the agreement between the profile measurements of 6MV circular SRS beams performed using a PTW PinPoint ionisation chamber, a PTW diamond detector, a novel silicon-diode array (DOSI), and a polymer-gel dosimeter. The authors found from their results that the PinPoint chamber would overestimate the penumbra width because of its finite size and the fact that the electron range in air was higher than that in water: the active volume of the PinPoint chamber was filled with air but the measurements were taken in water. Moreover, a number of investigators (McKerracher et al., 1999; Westermark et al., 2000; Martens et al., 2000; 22

25 Buccilioni et al., 2003; Pappas, 2001; Pappas et al., 2006) have reported that this kind of dosimeter broadens the penumbra and that in general even the ionisation chambers with small sensitive volume were not suitable for small-field profile measurements because (a) their finite size would introduce volume-averaging problems in high-gradient regions; and (b) the presence of air within the phantom would result in electron-transport alterations. The diamond detector was reported to provide accurate relative-dose measurements when used in small-field dosimetry due to their tissue equivalence, high spatial resolution, high sensitivity, and fast acquisition properties. However, diamond detectors showed great individuality with regard to their response depending on size, shape and purity of the crystal. They were also known to have a slight sub-linearity with dose rate (Haydarian, et al.). Moreover, possible directional dependence, possible asymmetric response and positioning difficulties could not be neglected for this type of detector, and were likely to reduce their measurement efficiency. The silicon diode array as a replacement for individual scanning diodes could provide high geometric resolution measurements capable of resolving steep dose gradients. However, they had the same drawbacks as the individual diodes, such as their response dependence on the spectral composition of the radiation beam, since they were non-tissue-equivalent materials. The authors concluded that the detector size and composition, together with positioning difficulties, would be the main problems related to small-field-profile measurements. Different measurement techniques had their own advantages and disadvantages, but none of them on their own could provide satisfactory accuracy and efficiency for small-field dosimetry. Laub et al. (2003) performed an experiment to investigate the volume effect of detectors in the dosimetry of small fields used in IMRT. They suggested that the delivery of small, multileaf collimated segments with at least one dimension smaller than 2cm was often required for IMRT to spare normal tissues; to calculate dose distributions and monitor units (MUs) for such small segments accurately, high-resolution absolute and relative dosimetry was of great importance. They also mentioned another problem: the exact position of multileaf collimators (MLCs) depended on the segment penumbra calculated in the IMRT planning engine, which occurred as a consequence of the step-and-shoot technique. They concluded that an inaccurate calculation of the penumbra could result in cold or hot spots between two adjacent segments. Consequently, the effect of detector size (volume effect) on the dosimetry of small fields, as well as the effect of steep-dose-gradient regions frequently encountered in IMRT, were critical for better accuracy. To investigate the effect, they measured the dose profiles and the output factors of the small field with 23

26 different detectors at three different photon energies: 06MV, 15MV and 18MV. In their experiment, different types of detectors were used. They first examined the dosimetry diode type (PTW Freiburg), which is a p-type Si diode. An over-response for this diode was expected for broad beams due to the relatively high atomic number of its material (silicon), but the authors thought that it was not critical for narrow-photon beams. Another disadvantage of single-diode detectors was the relatively large directional dependence of the detector response, which showed a variation of about 3%. However, it was still tested because its high spatial resolution was very advantageous in small-field dosimetry. Next, they examined the diamond detector type (PTW Freiburg). Its sensitive region is a low-impurity diamond disc with a thickness of 0.32mm and a volume of 3mm 3. Diamond detectors are known to show a slight sub-linearity of the current and dose rate. To compensate for this effect, a correction factor of approximately was applied to all measurements. This was only an estimate, and could not fully correct the sub-linearity. The third detector was the PinPoint ionisation chamber type (PTW Freiburg), or CC01, which is specifically designed for small-field dosimetry. The measuring volume of this ionisation chamber is only 0.015cm 3, providing a high spatial resolution for the measurements. However, the disadvantage of the chamber was that it over-responded to low-energy Compton scatters due to its aluminium electrode, as the authors had expected. The fourth detector examined was the 0.6cm 3 Farmer-type chamber.the TRS 398 protocol recommends this detector to measure the absorbed dose to water for a high-energy photon beam for field sizes less than 10cm x 10cm because of convenience in setup and relatively good signal reliability. However, the TRS 398 protocol also points out that when this chamber is used in measurements under non-reference conditions such as output-factor measurements, it is important to ensure the uniformity of the radiation fluence over the chamber cavity. For thimble chamber cavities that extend in the long axis, the variation in beam uniformity cannot be accounted for accurately. Hence large-volume thimble ionisation chambers are not recommended in small-field measurement (TRS 398, 2000). The film was also tested. At that time the Kodak EDR2 films were the commercially available dosimetry film, and were used in the experiment with a solid water phantom. Calibration data were measured for each film batch separately. The EDR2 film was used for the profile measurement, but not the output-factor measurement. In the profile measurement, because of the high resolution of the EDR2 film and the 24

27 scanner used, the authors assumed that the results measured by the film were the closest to the reality. However, since the water commissioning of the planning system was performed with a CC01 ionisation chamber, which had a limited spatial resolution, the insufficient spatial resolution of the commissioning detector led to local relative discrepancies of more than 10% between calculated cross-profiles and profiles measured with films. The results demonstrated that the commissioning of IMRT treatment planning tools with detectors that have a limited spatial resolution could lead to the introduction of systematic errors. The authors also mentioned that film measurements could still be problematic if the film used had a non-linearity response (which they did not investigate), and accurate film dosimetry for commissioning was very time-consuming. The deviation between absolute point-dose value as calculated by the planning system and dose values measured by the Farmer chamber and the CC01 chamber was interesting. It showed that measurements performed with a 0.6cm 3 Farmer chamber would lead to significant differences of more than 6%, while differences of less than 2% were found for dose values measured with a 0.015cm 3 PinPoint ionisation chamber. The authors suggested that the volume effect of the Farmer chamber could lead to inaccurate conclusions when comparing clinical verification dosimetry versus IMRT plans. The diode and the diamond detector both demonstrated directional dependence. When scanned in different directions, both detectors showed a response that differed by about 3%. The differences were bigger in the penumbra region. In the output-factor measurements, the authors found that for field sizes less than 3cm x 3cm, significant differences were found among different detectors. If measured with diode and diamond detectors, which were non-water-equivalent, the increase of the importance of secondary electrons in small fields would lead to an over-estimation of the output factors. On the other hand, if measured with the ionisation chambers, the increase of the lateral electron disequilibrium would give rise to an under-estimation of the output factors. The volume effect of the ionisation chambers would also reduce the accuracy. Therefore, the authors concluded that both the spatial resolution and water-equivalence of the detector were important for output-factor measurements in small fields. For ionisation chambers, the importance of the volume effect was small compared to the underestimation of the correct output factor that resulted from lateral electron disequilibrium. Laub et al. s experiment was a very helpful guide to this thesis, in that it compared the small field dosimetry measurements with most of the detectors commercially available at the time. From the measurements, the authors determined some advantages and disadvantages of each detector. However, there were still some issues that the article 25

28 did not address: 1) In the experiment, the authors only compared the profile measurements and the output-factor measurements; other critical measurements, such as the signal-to-noise ratio, the depth dose measurements and the dose-distribution measurements, were not mentioned. Including these measurements could yield a different conclusion. 2) As comprehensive as the comparison was, the report did not give clear guidance on which detector is most suitable for a certain situation. For a clinical protocol, this is very important. 3) The detectors used in the experiment were not the most up-to-date, as the experiment was performed eight years ago. Moreover, the Kodak EDR2 film has been replaced by the Gafchromic EBT2 and EBT3 film. Based on this literature review, the framework of the thesis was decided, and is described in Chapter Dosimeters 26

29 In this research, not all the currently available detectors suitable for small-field dosimetry were compared. This was due to the limitation of resources: only detectors readily available to the researcher s department were tested. The detectors tested were different-volume ionisation chambers; the p-type photon diode; and the Gafchromic EBT2 film. This precluded diamond detectors and thermo-luminescence dosimeters (TLDs). Although the use of diamond detectors in radiation therapy has been reported by others (Laub et al., 2003; Pappas et al., 2008; Kania et al., 1993), there were no diamond detectors available for this study. The TLD detectors were a potential option for small-field dosimetry due to their good spatial resolution and ability to provide multiple-point doses in a single exposure (Podgorsak, 2005). However, TLDs are time-consuming to employ as they require multiple readout and calibration procedures. For this reason this report excluded TLDs as a dosimeter, instead using the Gafchromic EBT2 film as a representative of high-spatial-resolution 2D detectors. This section briefly describes the working principles of the detectors. TLD and diamond, although not used in the thesis, are still described Ionisation chambers 27

30 Thimble-type ionisation chambers are currently the most frequently used dosimeters in medical radiation physics because of their ease of use (Metcalfe et al., 2007). The working principle of the thimble chamber was well explained by Khan (1994). The ideal ionisation chamber would be composed of a spherical or cylindrical volume of air with central air cavity. Suppose this sphere of air is irradiated uniformly with a photon beam. Also, suppose that the distance between the outer sphere and the inner cavity is equal to the maximum range of electrons generated in air. If the number of electrons entering the cavity is the same as that leaving the cavity, electronic equilibrium exists. Suppose also that the air surrounding the cavity can be measured. By knowing the volume or mass of air inside the cavity, the charge per unit mass or the beam exposure at the centre of the cavity can be calculated. The dose in the cavity is then estimated by multiplying the exposure Q m aaa by the energy expended in air per ion formed W aaa e, which is a constant of 33.97eV / pair: D cccccc = Q W aaa m aaa e The dose in the surrounding medium can be related to the dose in the cavity through the use of the Bragg-Gray cavity theory (Khan, 1992). According to the Bragg-Gray theory, the ionisation produced in a gas-filled cavity placed in a medium is related to the energy absorbed in the surrounding medium. When the cavity is small enough, that its introduction into the medium does not alter the number or distribution of the electrons that would exist in the medium without the cavity, the following Bragg-Gray relationship is satisfied: D mmm = D cccccc ( S ρ ) mmm,ccc where D med is the absorbed dose in the medium (in the absence of the cavity), and ( S ρ) mmm,ccc is the ratio of the average unrestricted mass-collision stopping power of the medium and the cavity. In practice, the thicknesses of the chamber wall are usually designed to be much thinner than the range of the secondary charged particles, so that the proportion of the cavity dose due to electrons generated in the phantom greatly exceeds the dose contribution from the chamber wall. In this case the phantom medium serves as the medium and the chamber is treated as a perturbation component (Podgorsak, 2005). A thimble chamber, based on the above principle, is built up by compressing the air wall into a solid shell. Although the thimble wall is solid, it is air-equivalent; in other words, its effective atomic number is close to that of air. In addition, the thickness of the thimble wall is such that the electronic equilibrium occurs inside the cavity; this requires that the wall thickness must be equal to or greater than the maximum range of the electrons liberated in the thimble wall. In practice, a thimble chamber is constructed with wall thicknesses of 1mm or less, supplemented with close-fitting 28

31 caps of Plexiglas or other plastic to bring the total wall thickness up to that needed for electronic equilibrium for the radiation in question Semiconductor diodes 29

32 For dosimetry of ionising radiation several types of semiconductor devices can be used. Compared to ionisation chambers, they typically have a small active volume, as the energy required to creating an ion pair is some 10 times smaller when a solid rather than a gaseous detector material is used (Metcalfe 2007). What s more, because solid material is about 1,000 times denser than air, ion pairs are more likely to occur. When the diode is irradiated, radiation produces electron-hole pairs in the body of the dosimeter, including the depletion layer. The charges produced in the body of the dosimeter, within the diffusion length, diffuse into the depleted region. They are swept across the depletion region under the action of the electric field due to the intrinsic potential. In this way a current is generated in the reverse direction in the diode; by measuring this current, the amount of radiation delivered can be predicted. A silicon-diode dosimeter is a p-n junction diode. The diodes are produced by taking n-type or p-type silicon and counter-doping the surface to produce the opposite type material. Only the p-type diode is suitable for radiotherapy dosimetry, as it is less affected by radiation damage and has a much smaller dark current (Podgorsak, 2005). Diode semiconductors can be operated in two different modes: with and without bias. In the first case, the resistance of the diode is taken as a measure of the intensity of the radiation. The current is measured as a function of the bias using an ampere meter. In the second mode no bias is required, and the diode works similarly to a solar cell. In this case the current generated by the diode is proportional to the radiation intensity (or dose rate), and consequently the charge collected from the diode can be taken as a measure of dose. For radiation applications the diode is usually used in this photovoltaic mode; the junction electrodes are left open and no external bias is applied. The diode biases itself in the forward direction, so that if no radiation is applied the net current is zero and an open-circuit voltage is developed. A number of diodes limitations should be noted when they are applied in radiation measurements. First, their conductivity and radiation response is temperature-dependent. Second, according to Van Dam et al. (1990), their readings display a dose-rate dependence. This can be particularly of concern in pulsed radiation produced by medical linear accelerators, as it may result in non-linear dose measurements. Third, the shape of the p-n junction and its geometry determines the directional response of semiconductors. Additionally there is a directional dependence of the dose reading; this is enhanced by the presence of the cable connection. Finally, the most subtle problem is the relatively high atomic number in the active zone, which can make the detector susceptible to slight variations in the spectral composition of radiation beams (Yin et al., 2004) Radiochromic Films 30

33 The use of radiation-sensitive film was one of the earliest detection methods of X-rays (Trevert, 1896). Film dosimetry is currently undergoing a dramatic change, as conventional radiographic films are being replaced by new 2-D dosimetric media such as radiochromic films. The recent increased use of radiochromic film has led the American Association of Physicists in Medicine (AAPM) to assigning a task group to report on the issue. The most commonly used type of radiochromic film is Gafchromic film, which is a trademark of International Specialty Products (ISP) Technologies. Radiochromic film contains a dye that is polymerised upon exposure to radiation. The polymer absorbs light, and the transmission of light through the film can be measured with a suitable densitometer. Radiochromic film is self-developing, requiring neither developer nor fixer. Since radiochromic film is grainless, it has a very high resolution and can be used in high-dose-gradient regions for dosimetry (Podgorsak, 2005). The film used in this work is the Gafchromic EBT2 QD+ film. Initially, Gafchromic film was developed for dose monitoring in industrial radiation processing. These relatively insensitive films were suitable for relatively high-dose measurements (in excess of 50Gy to 2500Gy) and were used for several years for clinical dosimetry research under the name HD-810. Then a more sensitive Gafchromic film, the MD-55 model, was developed, with a thicker sensitive volume and a dose range from 10Gy to 100Gy. Although the sensitivity of the MD-55 model was relatively high, its dose response was reported to be non-uniform by 8 to 15% (Meigooni et al., 1996), limiting its use in clinical applications. Consequently, further radiochromic film developments led to the External Beam Therapy (EBT) Gafchromic film model (Lynch et al., 2004), initially designed to replace silver-halide radiographic film for intensity modulated radiotherapy QA procedures (Devic, 2011). The manufacturer claims that the new EBT2 has a number of advantages: it is self-developing; requires no processing; is largely energy-independent; is water-resistant; offers more stability in room light compared with earlier Gafchromic films and radiographic films; and has built-in uniformity enhancement. When using a radiochromic film, a number of points should still be considered, as it is a relatively new dosimetry technique: 1) Although radiochromic film is claimed to be self-developing, it takes at least two hours for the colour change to stabilise sufficiently for evaluation (Chu et al., 1990). Even with the most recent products, the manufacturer still recommends a two-hour developing time before analysis is conducted. 2) Excess humidity at temperature above 25 should be avoided. 3) Ultraviolet light may cause a colour change without exposure to ionising radiation; this has to be taken into account when light is used for evaluation (Butson et al., 2003a). 31

34 4) The measured optical density in all radiochromic films depends on the wavelength of the light used for the evaluation (Butson et al., 2003a). 5) Radiochromic film is a relative dosimeter, and dose standards must be included in the evaluation. 6) The orientation of the film scanning, including sides up and down must be kept consistent. 7) No dose rate or fractionation dependence has been reported (Butson et al 2003a). 8) Gafchromic film is not entirely tissue-equivalent. At photon energies around 25 kev its dose response is only about half that of water (Kron et al., 1998). 9) The edge of the film where it has been cut may have a different response to the rest of the film. This issue has been discussed in the film session of the EPSM 2010 meeting but to date there are no publications that address this issue Thermoluminescence Dosimeters (TLDs) TLDs have been important in the history of the small-field dosimetry. Compared with 32

35 other dosimetry techniques, thermoluminescence dosimetry is a relatively old technique, firstly used some 100 years ago. With its early invention TLD has until recently been almost the only available dosimeter for small fields. It has the advantage of high spatial resolution for measuring the high-dose gradients such as the dose build-up region and the beam penumbra, dose-response linearity and flexibility in shaping (Kron et al., 1993). TLDs consist of crystals that are in general non-conducting i.e., the conduction band in the band model of the crystal is nearly empty. At room temperature most of the electrons are confined to the valence band. If the crystal is irradiated with ionising radiation, some of the electrons gain enough energy to get into the conductivity band. Thermoluminescent (TL) materials contain a number of imperfections in the crystal that can trap electrons from the conduction band at an energy state that is between the conduction and valence bands. The energy gap between the conduction band and the trap is of the order of a few electron volts, and the number of electrons trapped is a function of the incident radiation intensity. Typically less than 1% of the electrons in the conduction band are trapped (Horowitz, 1984c). The probability that the electrons can gain enough energy to escape back to the conduction band, from which they can fall back to the lowest energy state in the valence band, depends on the depth of these traps and the temperature. On the other hand, positively charged holes created by the ionising radiation in the crystal structure behave in a similar way. With rising temperatures, the probability increases that electrons have enough energy to be raised to the conduction band and release energy in the form of light when they recombine with a positive hole in the valence band. As with diodes, impurities again play an important role. They create the centres where the transition of the electron from conduction to valence band results in the emission of light. The amount of the light emitted depends on the temperature and the number of electrons trapped, and therefore on the absorbed dose in the crystal. TLDs can be made from different materials. One of the most widely used TL materials is lithium fluoride doped with magnesium and titanium (LiF: Mg, Ti) to introduce impurity. TLDs have disadvantages yet to overcome. First, they have a very poor signal-to-noise ratio. A noise equal to 20% of the primary signal is not rare. Although increasing the sampling number can help reduce the noise, it greatly increases the experiment time. Second, the processing (annealing/calibration) procedures for TLDs are very time-consuming. Despite the fact that TLDs are reusable, they require a long processing time before the chip is ready for the next reading. Finally, the advantages of the TLDs can now be achieved easily with more-advanced dosimetric techniques such as diodes and Gafchromic film. Compared with Gafchromic film, TLDs cannot provide as high spatial resolution or a 2D dose map. One can argue that it is possible to decrease the dimensions of the TLD used, but the smaller the TLD is to 33

36 reduce volume averaging and increase spatial resolution, the less the received light signal, and thus the higher the noise, (Kron et al., 1993). Compared with diodes and ionisation chambers, TLDs, as integral dosimeters, cannot supply real-time signal and ease of setup. After careful consideration, the TLD was rejected for use in this thesis Diamond Detectors The diamond detector is usually designed as a metal-insulator-metal (MIM) structure. 34

37 In this structure, a high-sensitivity diamond is sandwiched between two metal electrodes as ohmic contacts. An external voltage provides an electric field across the device (Kania et al., 1993). When radiation interacts with the device, free charges, i.e., electrons and holes, are produced in the crystal. These generated mobile charges drift across the depleted region in the electric field and generate a current in the external circuit, which is proportional to the intensity of the radiation field. Diamond detectors have some advantages. The high resistivity of the intrinsic diamond reduces the amount of the thermally generated leakage. It is also radiation-hard and chemically inert. Diamond detectors can also be made small to provide reasonable spatial resolution. However, diamond s low atomic number is not ideal for measuring penetrating radiations such as MV photon beams, because the low Z implies a small reaction cross-section and poor detection efficiency. The large band gap in the diamond further reduces its sensitivity. This also means that the detector s sensitivity is energy-dependent. Although the diamond detector has been studied by many researches, currently there are not yet any commercially available and mature diamond detectors that have been clinically proven. Chapter 3: Materials and Methods 35

38 3.1 General Experimental Design The dosimeters used in this study were the CC01, CC04 and CC13 ionisation chambers, a silicon diode and the Gafchromic EBT2 film. The performances of different measurement techniques in small radiation fields were assessed in terms of: 1. Signal reliability and signal reproducibility, such as signal-to-noise ratio and dose-response linearity; 2. Characterisation in commissioning data collection, which includes curve measurement and point-dose measurement; and 3. Efficiency in planning dosimetry verification for the purpose of clinical routine QA. To assess the detectors three major properties, the following work was planned: 1) Signal-to-noise ratio measurements; 2) Investigation of the dose response of the Gafchromic EBT2 film; 3) Data collection, including measurements of output factors, percentage depth dose and off-central-axis ratio; and 4) IMRT dose verification, including verification of point dose and planar dose distribution. The measurements were performed on the Siemens Artieste linac, which provided both 6MV and 10MV photon beams. All measurements were performed with the 6MV photon beam only, since this photon energy is more often used in IMRT treatment for most anatomical sites, such as head and neck, chest wall and prostate, at the author s department. First, signal reliability and reproducibility was assessed in terms of signal-to-noise ratio. The signal reliability and reproducibility property was crucial to all ionisation measurement techniques, as good signal reliability and reproducibility would help assure both the precision and the confidence level of the results, while bad signal reliability and reproducibility would make the results scientifically useless. The effect is not negligible in both large-field measurements and small-field measurements, but is even greater in small-field measurements due to the relatively low signal magnitude. As a consequence, the first part of the work was to measure and compare the signal-to-noise ratios of the chosen detectors. Section gives details of the setup. Next, the dose response of the Gafchromic film was investigated. As discussed in section 2.4.3, the dose response of the Gafchromic film depended greatly on the scanning parameters such as the orientation and the colour region used. In this work, 36

39 the Gafchromic films were irradiated with different doses, and then processed in terms of grey-scale value (GSV) using digital imaging processing techniques, rather than conventional optical density readings. This GSV corresponded to the change in colour of the Gafchromic film, which changed upon irradiation. After development, the irradiated films were scanned using a pre-set scanning setup at a fixed position of the scanner. The films were then saved digitally as JPEG files in the computer. To process the film data, the JPEG files were imported into the RODOMS software. The software gave the GSV value of each film, which corresponded to a certain dose/mu value. Based on the results from all the irradiated films, a calibration chart of GSV to dose was built up, which illustrated the dose response of the Gafchromic film when processed in GSV. If the dose response of the Gafchromic film was not linear, a correction factor would need to be applied to achieve linearity. The dose responses of the other detectors, such as ionisation chambers and diodes, were not investigated, because they were already well known from profound relevant studies. Section gives the detailed setup of this part of the work. Third, the commissioning data for small fields were collected using different measurement techniques. The X-ray MV commissioning usually required two parts of measurements: the profile data collection, comprised of the output-factor measurement, the percentage depth dose (PDD) curve collection and the off-central-axis ratio (OCR) profile curve collection; and the absolute-dose verification. The output factor was measured first. The dose output was monitored by the ionisation chambers in the head of a linac. Feedback circuits from these chambers had a calibration adjustment at the console control that could be adjusted to match a calibrated dose rate. The units recorded at the control console from the linac ionisation chambers were called monitor units (MUs). However, 1MU equaled a different dose at different field sizes, since the monitor ionisation chambers were located above the final beam collimation system. Thus, the wider the collimators were set, the more dose was delivered per MU, due to an increase in head and phantom scatter. To compensate for this effect, output factor ratios were applied to the reference field in the planning dose calculation: output factors were measured and input into the planning system directly during the commissioning; this required high accuracy and high precision for all field sizes. The field factor Scp, which combined the collimator scatter factor Sc and the phantom scatter factor Sp, was measured as the output factor for different fields. The measurement was performed in water (or water-equivalent materials) rather than in air, since Sc was not to be measured. Results from different detectors were compared and any observed differences were discussed. Section gives the detailed stup of this part of the work. The next stage was to measure the percentage depth dose (PDD): for computer treatment planning calculations, a reference dose is usually calculated at the depth of 37

40 maximum dose dmax, where the percentage relative dose is normalised to Dmax. The dose within the body is then considered in terms of percentage depth dose, %D, which is the ratio, expressed as a percentage, of the absorbed dose at any point to the dose at Dmax. Consequently, in the planning system, if the value of the absolute absorbed dose at a point, for example Dmax, is well known, and the percentage depth dose is fully collected, the absorbed dose for any other central-axis point can be calculated. This central axis percentage dose is often plotted against depth, as this indicates the penetration properties of the beam. This is commonly referred to as a PDD curve (Metcalfe et al., 2007). The PDD depends on four parameters: depth in phantom, field size, source-surface distance (SSD) and photon-beam energy. In this study, data collection was always performed in a water tank. However, the accuracy and the precision of the measured PDD curves for small-field beams may vary between different dosimeters due to different responses, characteristics and volume limitations, as was discussed in Chapter 2. On one hand, detectors with smaller volume would likely yield a higher spatial resolution, and consequently a more accurate PDD curve measurement. On the other hand, smaller volumes also give a lower signal-to-noise ratio and thus introduce low precision into the measurement. A compromise between high spatial resolution and low noise is usually the choice. In this study, the PDD curves were measured by different detectors, then analysed and the differences discussed. Section gives the detailed setup information of this part. Complete commissioning data collection also required the measurement of the OCR. Like the PDD, the OCR helps calculate the dose at points other than the reference point, but it is the lateral profile of the beam, giving the dose at an off-axis position relative to that on the central axis. It is generally assumed that these profiles are axially symmetric and only depended on the distance from the axis. The two most important regions of a megavoltage X-ray beam dose profile are the inter-umbral region, where the beam is not affected by the collimators, and the penumbral region, where the beam is affected by the field-defining collimators. The penumbral region is characterised by a sharp dose gradient where the primary photons are shielded by the collimators. However, there is a finite width of a few millimeters in the penumbra region due to the finite size of the virtual X-ray source and lateral electron disequilibrium (Metcalfe et al., 2007). It has historically been challenging to measure the penumbra region of the OCR curve, especially for small fields, since at this region the dose falls off rapidly in a very short distance. This leads to a sharp dose gradient, where the limited spatial resolution of the detectors is very likely to cause loss of the sharpness and details of the region. This is commonly referred to as the volume averaging of the detector, as the measured penumbra width measured by any detector is essentially broader than the penumbral width measured by a virtual zero-point detector. The issue is more severe for small field sizes, since for small beams with field sizes < 5cm x 5cm, the OCR 38

41 profile is mainly composed of the penumbra region; thus the accurate and precise measurement of the penumbra is particularly critical, especially for IMRT treatments, in which small segments are frequently encountered. As a consequence, the last part of small-field data collection in this study was to measure the OCR profile with different detectors under both small fields and regular-sized fields. Section gives the detailed setup information of this part. The absolute point-dose measurement was not compared among different detectors, since the IAEA Technical Reports Series No. 398, Absorbed Dose Determination in External Beam Radiotherapy, explained this clearly enough, and has become the standard protocol for most hospitals medical physics departments, including the author s department. The last session of the work was to conduct an IMRT QA case to confirm the results from previous sections and compare the efficiencies. There are currently two common ways to conduct IMRT quality assurance. The first method is to measure a single-point or multi-point dose of the IMRT case in a three-dimensional phantom and compare the result to the accumulated planning dose. This point dose is usually measured by small-volume ionisation chambers or diodes. In this case, the volume of the detector is very crucial to the accuracy and precision of the QA result, as while the reference point dose output is from a very small voxel in the planning system, the actual size of the space where the signal is measured is directly related to the volume of the detector but never a point. As a result, various detectors were used to measure the point dose of an IMRT case and the readings compared to the detectors volumes. The second method is to verify the IMRT s two-dimensional dose distribution in a certain plane and compare the calculated distribution with the actual measurements by a 2D detector, such as Gafchromic film, or a diode array exposed in the same plane. One way to evaluate the variation between the calculation and measurement could use the gamma method (Low et al., 1998), which recognises that the requirements for dosimetric accuracy are highest in regions of low dose gradient and the requirements for geometric accuracy were highest in regions of high dose gradient (Williams, 2003). Only Gafchromic EBT2 film was used to measure the dose distribution, but its performance was assessed. At the time this study was being conducted, neither a diode nor an ionisation chamber array were readily available. Section 3.2 lists the equipment used in the work; section 3.3 gives the detailed setup information; and section 3.4 and 3.5 give the calculation method of the error bars and the gamma function, respectively. 3.2 List of Equipment 39

42 3.2.1 Scanditronix-Wellhofer CC01 ionisation chamber The chamber has an active volume of 0.01cm 3, a radius of 1.0mm and a total length of 3.6mm, and its reference point is at 2.3mm from the distal end of the chamber thimble. The sensitivity of the chamber is claimed to be approximately 1.10 x 10-9 C/Gy. The CC01 ionisation chamber is an extremely small-volume waterproof chamber with very high spatial resolution for measurement in both water and solid-slab phantoms for small-field beam profile and output-factor data collection Scanditronix-Wellhofer CC04 ionisation chamber The chamber has an active volume of 0.04cm 3, a radius of 2.0mm and a total length of 3.6mm, and its reference point is at 2.3mm from the distal end of the chamber thimble. The sensitivity of the chamber is claimed to be approximately 1.10 x 10-9 C/Gy. The CC04 ionisation chamber is a small-volume waterproof chamber designed for measurement in both water and solid-slab phantoms for small-field beam profile and output-factor data collection Scanditronix-Wellhofer CC13 ionisation chamber The chamber has an active volume of 0.13 cm 3, a radius of 3.0mm and a total length of 5.8mm, and its reference point is at 3.5mm from the distal end of the chamber thimble. The sensitivity of the chamber is claimed to be 3.8 x 10-9 C/Gy. The CC13 ionisation chamber is a standard waterproof chamber designed for routine medium and large field-size beam curve data collection and output-factor measurement in both water and solid-slab phantoms IBA Dosimetry Diode, Type PFD-3G This is the third generation of psi semiconductors, with a high uniform spatial resolution in the beam plane and precise definition of the measurement depth, and is claimed to be independent of bias, pressure and moisture that is, very stable under different external conditions. It has a side length of 2.5mm, thickness of 0.5mm, and radius of 2.0mm. It is specifically designed for curve data collection in blue-water-tank scanning systems, and is being used increasingly often in small-field measurement because of its very small dimensions Gafchromic EBT2 QD+ Film This is the new generation of self-developing dosimetry film produced by ISP. The dimension of the film is 8 inch x 10 inch. The film is claimed to have a dose range of 1 centigray (cgy) to 10Gray (Gy), <10% energy dependence, tissue equivalent and stability under room conditions Fluke Therapy Dosimeter, SN99918 (Electrometer) This is designed for integrated dose measurement. Its leakage current is 40

43 negligible from tests and the instrument s sensitivity is high enough for the requirements of this experiment Epson Expression 10000XL Film Scanner with Film Transparency Adaptor This is an A3 scanner, designed for graphic-arts and business applications, with a resolution of 2,400 dots per inch (dpi) resolution, 3.8 optical density and network capability. It has become widely used in hospitals in recent years as a scanner for the processing of Gafchromic films due to its high resolution and stability. It is also recommended by the film manufacturer RODOMS (Radiation Oncology Dosimetry Management System) This is in-house (Genesis Cancer Care) software designed and programmed by Dr Yang Wang for use in radiation oncology and is authorised for use in this experiment. Although it is multi-functioned, the software was mainly used in this experiment to analyse and record the grey-scale value (GSV) of the film. The film-dosimetry function of the software includes scanner-signal calibration according to irradiated dose response; radiation-field check; reconstruction from exported planning-dose map to virtual film; and, most importantly, composite IMRT dose comparison and analysis with gamma-value function OmniPro-Accept This is a workflow-oriented water-phantom software designed for the commissioning and QA of the linac. It is produced by iba-dosimetry, and according to the manufacturer, the program can provide both step-by-step and continuous scanning at variable scan speeds. The version used in this experiment is the OmniPro-Accept 6.5A, and is used mainly for processing the beam-curve data (such as curve plot and ASCII data-point export) Linear Accelerator The linac used in these experiments is the Siemens Artiste at St Vincent s Clinic. This linac is capable of producing photon energies of 6MV and 10MV, and electron energies of 6MeV, 9MeV, 12MeV, 15MeV, 18MeV and 21MeV Ionisation-Chamber Cylindrical Phantom When testing clinical IMRT cases in this study, a specially designed coaxial cylindrical phantom is used to simulate the patient s geometry as closely as possible. Figure to show the relevant photos. 41

44 Figure : Cylindrical IMRT phantom for ionisation chamber Figure : Cylindrical IMRT phantom for ionisation chamber, longitudinal view 42

45 Figure : Cylindrical IMRT phantom for ionisation chamber, saggital view Figure : Cylindrical phantom and its chamber holders for different types of ionisation chambers The coaxial ionisation-chamber cylindrical phantom is made up of three parts: the supporting feet, the outer cylindrical phantom and the inner cylindrical chamber holder. The diameter for the outer phantom is 24.0cm; the diameter for the inner 43

46 phantom is 4.0cm; the length is 20.0cm. This phantom is designed and made locally with poly(methyl methacrylate) (PMMA) by Professor Yang Wang Film Cylindrical Phantom In a similar manner a cylindrical phantom designed for film measurement was used in this study (Figures to ). The phantom is composed of the supporting feet, the upper semi-circle half-phantom and the lower semi-circle half-phantom. The film is placed in between the two semicircles, which is the central plane of the whole cylinder. Figure : Cylindrical IMRT phantom for Gafchromic film 44

47 Figure : Cylindrical IMRT phantom for Gafchromic film, longitudinal view Figure : Top and bottom half of the Cylindrical IMRT phantom for Gafchromic film The diameter for the whole phantom is 24.0cm and the length is 30.0cm. This phantom is designed and made locally with PMMA by Professor Yang Wang. 45

48 3.3 Detailed Experiment Procedures Signal-to Noise-Ratio Measurements The signal-to-noise ratio experiment was performed first. The detectors tested were: CC01, CC04, CC13, p-type diode and Gafchromic EBT2 film. Before setting up the phantoms, the gantry angle, the collimator angle and the couch angle of the linac were set to 0 and the field size was set to a 10cm x 10cm square. First, 10cm white solid water slabs were placed on the couch as the backscatter material. For the ionisation chambers and the diode, a respective chamber-holder phantom slice was used (Figure ). The holder slice was designed such that once the chamber was fully inserted, the centre of the chamber was at the centre of the phantom, indicated by a cross. This holder phantom was placed on top of the backscatter materials, with its centre aligned with the optical crosshairs (Figure ). The ionisation chambers or the silicon diode were then inserted into the chamber holder after the alignment. Figure : Chamber-holder phantom slice for CC13 46

49 Figure : Alignment of the centre of the ion chamber/diode to the isocentre of the linac with the aid of the field-light crosshair Aligning the Gafchromic film was easier, as no holder phantom was required. To compensate for the influence of the variations between different films in this experiment, only one piece of film was used. This film was cut into 12 2cm x 20.2cm strips; the centre of the strip was then placed onto the top of the backscatter phantoms, with their geometrical centre aligned at the isocentre of the linac with the aid of the field-light crosshair (Figure ). 47

50 Figure : Alignment of the film stripe to the iso-centre of the linac with the aid of the field-light crosshair After alignment, a 10cm white solid water phantom slab was added to the top of the dosimeters as the build-up material. The front pointer was then employed to set a 100cm SSD to the top surface of the build-up phantom. The build-up material was set to 10cm depth rather than dmax depth (1.5cm for 6MV and 2.5cm for 10MV) because in the PDD curve, a 10cm depth was far beyond the build-up region of the beam; this ensured that electron equilibrium was fully achieved, and that the secondary scatters (scattered photons and electrons produced by photon interactions in the phantom) at this depth were already negligible. Thus it could be concluded that the fluctuation of the signals at this depth mainly came from the detector itself. After the dosimeters were set up, the ionisation chambers and the diode were connected to the electrometer via an extension. For the ionisation chambers, a -300V polarising voltage was applied; for the diode, no high voltage was applied. Then the diode and the ionisation chambers were warmed up, but not the film. Next, all the dosimeters were irradiated by 6MV X-rays with 200MU at a 10cm x 10cm field followed by 1cm x 1cm field. The measurement was repeated 10 times at each field size for each dosimeter to see the signal fluctuation. For the ionisation chambers and the diode, the electrometer reading was taken directly as the final result, but for the Gafchromic film, processing was compulsory before accurate results could be obtained from the film. Handling the exposed Gafchromic film posed considerable challenges, as 48

51 according to different literature reviews and the manufacturer s manual, improper handling of the exposed film would introduce significant errors to the measurement. As a result, the processing method for the Gafchromic film was carefully considered according to its properties. The processing method used here was applied to all the films used in the other parts in this experiment so that consistency was kept and no systematic error was likely to be introduced during the processing. First, the exposed films were kept in darkness for at least two hours for full development, as suggested by the manufacturer. Then the films were scanned with an Epson Expression 10000XL Photo flatbed scanner, which was recommended by the manufacturer for film digitisation. The film was placed in the centre of the scanning bed with the aid of a positioning frame so that every piece of film was scanned at the same position of the scanning bed. The film orientation was so that it was always scanned in the landscape orientation, the film s marking slit in the top right corner, and the beam-entry side of the film facing downwards the scanning bed. This scan setup was kept strictly consistent through the whole experiment to minimise systematic errors. To scan the film, the Epson Scan software was used. In this software, the user could choose the type (colour images of different bit dimensions or grey-scale images) and spatial resolution of the image to be scanned. The author could also adjst the contrast and the brightness of the film. All these parameters would potentially influence the calibration curve of the Gafchromic EBT2 film; thus their effects on the film processing were carefully studied, and it was found that among different combinations of setups, the following could provide the best sensitivity and linear calibration curve for the digitised Gafchromic EBT2 film (Figure ). The setup was applied to every piece of film used in the experiment. 49

52 Figure : Epson Scan software setup for Gafchromic EBT2 film scan In the software, the document type was set as film and the film type was positive film. The image type was 16-bit grey scale rather than the conventional 48-bit colour. The resolution was fixed at 100dpi to provide adequate spatial resolution (each dot had a size of 0.254mm). A higher spatial resolution could be used when required, but this would greatly increase the size of the digitised 50

53 image. In the Image Adjustment tab indicated by the red arrow, the brightness and the contrast were set to +70 and -70 respectively. All other setups remained in the default values. The digitised image was saved as a JPEG file. This JPEG image was then imported into the RODOMS software, which gave the GSV pixel. In the RODOMS software, the GSV pixel value could be read at any position on the digitised image. The software automatically applied a smoothing function when taking the readings. Generally, when a point was chosen, its pixel value was read along with the values for the surrounding eight pixel points. Thus for each point nine GSV values were read. Then the software indexed these nine values from high to low, and calculate an average, disregarding the highest and the lowest readings. This average was used as the final reading for the inspected point. The aim of this automatic function was to reduce the high noise of the Gafchromic film, although to a certain extent it also reduced the spatial resolution of the digitised image due to the smoothing. This average value then underwent a simple internal correction function in the RODOMS software. First, although the image was scanned in as a 16-bit image, the pixel values were read based on 8-bit information. The advantage of this was that it greatly decreased the amount of processed information. Although the sensitivity was reduced from 16-bit to 8-bit, it was considered still adequate for clinical applications. Second, in an unprocessed image the raw pixel colour values corresponding to the 0 dose level would have the highest values. This was not ideal; it would be preferable to have the pixel value directly proportional to the dose level. So in RODOMS, for 8-bit information, the pixel values of which ranged from 0 to 255, the raw pixels would be processed such that the exported pixel value would be equal to (255 raw pixel value); thus the corresponding data would increase with the dose level. This also meant that 255 was the maximum pixel value that could be exported from RODOMS. This processed pixel colour value was used as the final reading for the Gafchromic EBT2 film. Finally, the readings from the Gafchromic film, the diode and the ionisation chambers were normalised to the same level and plotted on the same graph for comparison. Section 4.1 gives the results and discussions. 51

54 3.3.2 Gafchromic EBT2 Film Dose Response Calibration Currently the most commonly accepted method of processing Gafchromic EBT2 film is to scan the film as a 48-bit colour image (red-green-blue) (Devic, 2011), and extracting the red channel for characterisation and analysis, as it is believed to have the highest absorption as shown in measured absorption spectra (Devic et al., 2007). The pixel value (PV) of the red image is converted to the optical density (OD) after a calibration curve is established, and the OD is converted to dose for analysis. The reason why OD is used because it can provide a relatively linear response to the dose based on its definition (Devic 2011). However, the author used a simple and novel way to process the film in this research: scanning the film as a 16-bit grey scale image and using the GSV to directly convert to the dose. The author found that because this method avoids the step of conversion to OD, less uncertainty (due to calibration fitting) is introduced and efficiency is improved, while still providing superior linearity in the dose response than that achieved with conventional processing. To confirm the above statement, one piece of Gafchromic film was cut into several strips and irradiated: 0, 10, 20, 50, 100, 150, 200, 250, 300 and 500cGy. The film was then analysed in the red, green, blue, and GSV channels, with the dose-response curves plotted on the same graph and compared. Once the general linearity was confirmed using the GSV processing method and the assigned scanner settings, its dose response was analysed in more detail. Three ranges of doses were defined: low, middle and high, each with 12 equivalent energy gaps. The low range comprised doses of 0, 5, 10, 15, 20, 25, 30, 35, 40, 45, 50 and 55cGy; the middle range 0, 50, 100, 150, 200, 250, 300, 350, 400, 450, 500 and 550cGy; and the high range 0, 100, 200, 300, 400, 500, 600, 700, 800, 900, 1000 and 1100cGy. In the experiments, a total of four slices of Gafchromic film were used, each cut into 12 strips measuring 2cm x 20.2cm; each strip was used for one of the dose ranges. The strip was put between 10cm SW backscatter and 10cm SW build-up, with its geometric centre aligned at the centre of the optical crosshair at gantry 0, collimator 0 and couch 0. Finally the front pointer was used to set the top of the build-up phantom at 100cm SSD (Figure ). The setup was the same as that of the film part of Section After set up, the film strips were irradiated with the dose steps at 6MV with a constant field size of 10cm x 10cm. Then the exposed films were processed and analysed as described in section The film readings were plotted against the corresponding doses as the dose response curve of the Gafchromic film. 52

55 3.3.3 Output-Factor Measurements The next step was to measure the output factor Scp. The Scp was collected at field sizes of 1cm x 1cm, 2cm x 2cm, 3cm x 3cm, 4cm x 4cm, 6cm x 6cm, 8cm x 8cm, 10cm x 10cm, 14cm x 14cm, 20cm x 20cm and 30cm x 30cm, using CC01, CC04, CC13, silicon diode and Gafchromic film. The setup was at 10cm build-up depth and 100cm SSD. The readings were normalised to that taken at the 10cm x 10cm field. For the water-resistant detectors (the ionisation chambers and the diode), measurements were taken in liquid water. As a start, the Scanditronix Blue Phantom water tank was placed under the head of the linac. The central crosses of the water tank were aligned with the optical crosshairs of the linac at gantry 0 and collimator 0. The Blue Phantom was then filled with de-ionised water to the maximum height. With the water fully filled, the tank was leveled using the leveling system. Finally the water-resistant detector was inserted into the respective holder on the moving bar of the tank. In the next step, the water surface was aligned to 100cm SSD using the front pointer. Then the detector was aligned such that its effective measurement point was at the centre of the field crosshair in the X-Y direction and at the water surface in the Z direction. This point was defined as the iso-centre point in the OmniPro water-tank software. From this point, the detector descended vertically in the Z direction along the central axis (CAX) for 10cm in the tank. This placed the detector at 10cm depth along the central axis and at 100cm SSD to the surface of the water. Once set up, the detector was connected to an electrometer. A 300V bias was set for the ionisation chambers, but no high voltage (0V) was applied to the diode. Next, 1000MU at 06MV photon beam was delivered to the detector as a warm-up. Then the detector was irradiated with 200MU at the proposed field sizes from 30cm x 30cm down to as small as 1cm x 1cm. The reading from the electrometer was recorded. Once all the measurements were taken, the readings were normalised to that of the 10cm x 10cm field to obtain the output factors. As stated in section 2.4.3, Gafchromic film was expected to have a high noise level and high dependence on the external conditions such as temperature, pressure and production variation (Crijins et al., 2011). For these reasons, more readings were required to reduce the possible noise. Although Gafchromic film could be used for measurements in a water tank, a special jig would have been required. This was not available to the author at the time, so instead an RW3 solid water phantom was used. With the gantry and collimator at 0, the Gafchromic film strip was placed on top of the 10cm RW3 backscatter material, and its geometric centre was aligned 53

56 with the field light. Then another 10cm RW3 slab phantom was added on top of the film as the build-up. Finally the front pointer was used to set the top of the build-up phantom to 100cm SSD. This closely simulated the scanning environment for the ionisation chambers and the diode, except that the scanning material for Gafchromic film was not de-ionized liquid water but RW3 slab phantoms. However, as the RW3 slab phantom had very close physical properties (physical density) and chemical properties (electron density) to real water, the difference was considered negligible. After alignment, each film strip was irradiated with 600MU at 06MV photon beam with field sizes from 30cm x 30cm to 1cm x 1cm. Multiple film strips were used at the same field size to increase the sample number and thus reduce the noise. Readings were then acquired from the film using the method in section All the readings were normalised to the reading at 10cm x 10cm field to obtain the output factors. Once the output factors of all the tested detectors were collected, they were plotted against the corresponding field sizes in the same graph. The next section gives the results of the comparison and analysis. 54

57 3.3.4 Percentage Depth Dose Measurements Different detectors CC01, CC04, CC13 and silicon diode were used to measure the PDD curves of the linac at both 1cm x 1cm and 10cm x 10cm fields. Consideration was given to whether the Gafchromic film should be used for the PDD measurement here and the profile measurement later. Both measurements required continuous scanning at a constant movement speed to collect the complete curve. This continuous-scan mode was available in the Scanditronix Blue Phantom, where the ionisation chambers and the silicon diode could be used. On the other hand, the Gafchromic film was not suitable for measurements in the Blue Phantom; RW3 slab phantoms were used to simulate the scan environment. With RW3 slab phantoms, a few measurement points could be taken at discreet depths, but the collection of the full PDD/profile curve in a continuous mode was impossible. Extrapolation between the points to complete the curve was considered too inaccurate, and no accurate extrapolation function could be found. The Gafchromic film also had a high noise level, unreliable reproducibility and poor stability; these defects would greatly influence the quality and the accuracy of the collected curve. The air gap between the RW3 slab phantom and the piece of the film was also a concern, as it might well cause spikes in the curve. The measurement of the PDD with Gafhcromic film would require the film to be scanned parallel to the beam central axis, but Suchowerska et al. (2000) found that films could over-respond when exposed parallel to the central axis of the beam, as opposed to perpendicular exposure. While their paper explains many complexities of interaction effects on film, the existence of the air gap is a confounding factor. As a result, the Gafchromic film was not used for either the PDD collection described in this section or the profile collection in the next section. To set up, first the central cross marks of the Scanditronix Blue Phantom were aligned with the optical crosshairs of the linac at the maximum available field size at gantry 0 and collimator 0. The tank was then filled with de-ionised water. Then the dosimeter was fit into the chamber holder, which was attached to a frame that could move along all three axes. The limits of the movement were set with close observation to avoid any possible chamber collision during the scan. In the next step the reference point of the dosimeter was aligned to the iso-centre of the linac, with the optical crosshair and the height of the reference point exactly at the water surface. According to IAEA report TRS 398 section 4.2.5, for cylindrical chamber types the reference point referred to the centre of the chamber s cavity volume on the chamber axis. This position was saved as the iso-centre of the water-tank system, and its position reproducibility was consistently checked during the scan to ensure accuracy. With the field dosimeter fully in water, the water surface was set to 100cm SSD using the front pointer and checked by the laser. 55

58 Besides the field dosimeter, a reference chamber was also needed in the Blue Phantom system for PDD and profile measurements. The position of the reference chamber was usually closer to the head of the linac than that of the field detector. This chamber was used to detect the signal fluctuation in the primary beam and compensate for its influence on the field detector; thus the reference chamber was particularly useful in reducing the noise from the primary beam. The choice of the reference detector was very important. As data at very small fields (down to 1cm x 1cm) would be collected, this reference chamber needed to be small enough that its effective measurement region could be covered in the small fields, but it would not block the scan passage of the field dosimeter. On the other hand, it should not introduce much noise from the detector; otherwise, the application of a reference chamber to reduce noise would be meaningless. This was actually the aim of this thesis to find a detector with low noise but high spatial resolution but at this stage no conclusion had yet been drawn. From previous clinical experience, it was decided that CC04 ionisation chamber would be used as the reference chamber for all the PDD and profile measurements. It could also be used as the reference chamber for the diode because the Scanditronix Blue Phantom allowed the voltages for the reference detector and the field detector to be set separately. The CC04 chamber was placed into the reference chamber arm. Its position was carefully set so that it was fully covered in the field but not blocking the scan path of the field detector in any direction. Its height was set halfway between the water surface and the linac head. This was not critical, though, as the PDD and the profile measurements were both relative measurements; as long as the position of the reference detector did not change, its height would not influence the final result. Finally, the field detector was connected to the field-detector panel on the control unit of the Blue Phantom, and the reference detector to the reference-detector panel, via optical cables. This control unit, or CCU, functioned as an electrometer as well. The field size was at 1cm x 1cm. The energy was the 6MV photon beam. After the setup was complete, the OmniPro-Accept software, which had been packaged with the Blue Phantom, was used to perform the scan. Before measurement, 300V was applied to the ionisation chambers, including the reference CC04 chamber, but no bias was applied to the diode. The readings of the field detector at Dmax and the reference detector were then normalised to 100%, and a background reading was also taken. Then the PDD scan was performed. With beam s on, the software automatically controlled the motion of the dosimeter to ascend vertically along the CAX from a depth of 36cm in water to 0.5cm above the water surface and to collect the signals from the CCU. The scan was in continuous mode, at a constant scan speed of 5mm/s. The software 56

59 automatically plotted the PDD curve while the scan progressed. When all the scans at the 1cm x 1cm field were completed, the same procedure was repeated at the 10cm x 10cm field. The raw PDD curves were then slightly smoothed without changing the shape in the OmniPro. The software s smoothing function was set such that the smooth interpolation was linear with an internal step width of 0.5mm, and the smooth function was envelope with a mean-value region of 2mm. A normalisation to dmax (100% of the dose) was also applied to the curve. The processed curves were plotted on the same graph and analysed. 57

60 3.3.5 Off-Central-Axis Profile Measurements In this section, the ionisation chambers and the diode were used to measure the OCR profiles and the results were compared and analysed with care. The Gafchromic film was not tested, as discussed section The profile scan was also performed in the Blue Phantom, using the same process to set up the water tank, the field detector and the reference detector as that described in section The same reference detector (the CC04 ionisation chamber) was also used. During the alignment, the scan line was carefully positioned along the middle of an MLC leaf to avoid the MLC interleaf leakage, which is more significant for small-volume detectors. After both detectors were properly set up and the iso-centre of the measurement well defined, the field detector was then lowered vertically along the central axis to a depth of 10cm, where the OCR profiles would be measured. Both dosimeters were then connected to the CCU and the correct bias was applied. The field size was at 1cm x 1cm and the energy for the scan was 6MV photon. After the setup, a background and a normalisation were performed. Then the OmniPro-Accept was used to perform the OCR profile scan. When beam was on, the software automatically controlled the motion of the dosimeter to move horizontally from left to right in the cross-plane in a range of 7.15cm deviation from the isocentre in both directions. The scan was set in the continuous mode at a constant speed of 5mm/s. The signal was automatically recorded and plotted by the software. When the scans at the 1cm x 1cm field were completed, measurements were repeated at the 10cm x 10cm field with the same photon energy, except that the scan region was set to 11.17cm off the CAX in both directions since the field size was larger. The curves were then smoothed as described in section 3.3.4, normalised to the same level and plotted on the same graph for analysis. 58

61 3.3.6 Clinical IMRT Case Comparison In the last part of the experiment, a clinical IMRT plan generated by CMS XiO was delivered and quality was assured using different dosimeters. The IMRT plan chosen was a 6MV photon plan with four fields. The target of the plan was the right malignant neoplasm of a kidney gland. The major organs-at-risk were the spinal cord, the bladder and the rectum. The plan featured a uniform target dose and low doses at the surrounding OARs. Measurements from various dosimeters in this condition could be clearly observed, and any difference could easily be distinguished. For this part, the ionisation chambers (the CC01, CC04 and CC13) and the Gafchromic film were used. The silicon diode was not tested because of its directional dependence on different beam angles, which is discussed in detail in a number of publications. The directional dependence of the diode had not been a concern in the earlier parts of the experiment because the previous tests had been performed in a fixed gantry angle of 0, where the diode axis was kept parallel to the beam axis. In an IMRT treatment, on the other hand, the gantry could rotate to any angle in 360 during the treatment, in which case the directional dependence of the diode could not be neglected. The effect of the diode s directional dependence was very hard to quantify; thus a correction factor for this effect was not achievable. Consequently, the diode was not tested in this part of the experiment. Because of the difference in dosimetric properties and geometry, the tests for the ionisation chambers and the Gafchromic film were designed in two different ways based on the current popular IMRT QA techniques (discussed in section 3.1). The ionisation chambers were used to acquire a point dose, while the Gafchromic film was used to acquire a planar dose distribution. Different phantoms were designed for each purpose. In this experiment, the planning output was determined to be the comparison standard. For the ionisation chambers, this would be a point-dose comparison; consequently a certain point position needed to be specified for comparison. After a review of the nominated IMRT plan in the planning system, it was decided that this point of interest would be the iso-centre of the linear accelerator. Choosing this point had two advantages: a) it was easier to set up the phantom at the iso-centre than the other off-centre points; and b) for this chosen plan with four fields, the first three fields had a low-gradient, high-dose volume for the target cover surrounding the iso-centre, while the last field had a dose gradient sitting close to the iso-centre. Consequently, QA is accomplished by drawing conclusions from a comparison of the results for each of these four fields. For the Gafchromic film, a whole-plane dose distribution comparison was needed. This was chosen to be the coronal plane at a 100cm source axis 59

62 distance (SAD). There were three reasons to use this plane: a) this was the plane where the iso-centre lay, so that the setup of the film was kept as close to that of the ionisation chambers as possible; b) setup was easier for this plane; and c) after viewing the system, it was seen that the dose distribution in this plane was relatively uniform and the target surrounded the centre of the plane. To generate the planning-comparison standard, the two cylindrical phantoms that would be used in the measurements (which are described in detail later) were first scanned by the GE LightSpeed CT scanner and stored in the planning system. During CT scanning the phantom was set up the same way as the linac. After scanning, the body contour was drawn on the phantoms. Then the nominated IMRT plan was calculated on the two phantoms with the superposition algorithm and a calculation grid size of 2mm in XiO. The XiO superposition algorithm is an adaption of the collapsed cone dose-calculation method. In this algorithm, all calculations are done in beam coordinates, and the dose in the beam coordinates is interpolated to the user-specified calculation volume. The superposition algorithm directly emulates the kernel calculation process: that is, it calculates deposited energy by spreading energy released at the interaction points, to points in the volume of interest, according to the distribution implied by the kernel. The kernels can be modified to account for variations in the electron density (Muralidhar et al., 2009). For the ionisation chambers, the dose at the point of interest (iso-centre) was exported in a table. For the Gafchromic film, the dose distribution at the chosen plane was calculated, and then exported from the XiO planning software via the Dose Plane Output function; the exported dose map was in text format and was called a virtual film. It was then imported into the RODOMS, converted to the GSV map in colour and used as the standard. Following the planning output was the measurement of the ionisation chambers (the CC01, CC04 and CC13). First, a calibration factor was necessary to convert the electrometer reading into the absolute dose. To do this, 10cm RW3 slab phantoms were placed on the couch as the backscatter material at gantry 0 and collimator 0. Then different chamber-holder phantom slices were used for different ionisation chambers, and the central cross marks of the phantom were aligned to the light-field crosshairs of the linac. The intersection of the cross marks of the slice corresponded to the geometric centre of the chamber if properly inserted; thus the proper alignment of the holder phantom ensured the proper alignment of the ionisation chamber. The chamber holder slice had a build-up equivalent to 1cm water, so after the alignment, another 0.5cm RW3 slab phantom was added on top to make the accumulated build-up equivalent to 1.5cm of water, the nominated Dmax of the 6MV photon. Finally, the front pointer was used to set the top of the build-up phantom at 100cm SSD. According to the definition of the absolute-dose calibration for this linac, at this point 1MU exactly 60

63 equalled 1cGy, and consequently the chamber reading on the electrometer at this point could be directly converted to the absolute dose. When an ionisation chamber is used for absolute calibration, its position at the reference depth needs to be considered. A chamber positioned with its cavity centre at Z ref does not sample the electron fluence present at Z ref, due to the finite size of the secondary charged particles. This perturbation may be accounted for either by applying a displacement correction factor or by displacing the chamber by an amount that compensates for this effect (TRS-398, 2000). In this thesis, however, this displacement correction method was not applied. Instead, the final result the measured IMRT dose is actually a ratio of the two ionisation readings: one measured in solid water phantoms and one measured in the cylindrical phantom. This is a reasonable approximation because of the slow variation with depth of water/air stopping-power ratios and the constancy of perturbation factors beyond Z max for photon beams. As the final dose reading is obtained as a ratio of ionisations, it does not require the use of displacement correction factors at two depths. The ionisation chamber was then connected to the electrometer with 300V high bias applied. Following the connection, 100MU at 6MV was delivered to the ionisation chamber, which in this setup equalled 100cGy. The measurement was repeated three times and the average of the readings was taken as the calibration factor from the electrometer reading to the actual dose, according to the following equation: Dose delivered (Gy) = Electrometer reading (nc) / Conversion factor (nc/gy). When the conversion factor was ready, the IMRT plan was delivered on the linac, with the ionisation chambers sitting in the geometrical centre of the ionisation-chamber cylindrical phantom. The cylindrical phantom was placed firmly on the couch. Then the ionisation chambers were carefully placed into the corresponding chamber holder of the phantom. The geometric centre of the chamber was aligned with the lasers in the three dimensions to the iso-centre of the linac at 100cm SAD. After proper alignment, the chamber was connected to the electrometer with a 300V high bias applied. The chamber-electrometer system was warmed up before the measurements. Then the selected IMRT plan was delivered on the linac, with the chamber-electrometer system recording the reading at the iso-centre for each beam. The measurement was repeated and the average reading was converted into absolute dose by the calibration factor achieved. The film was measured with the film cylindrical phantom. The IMRT plan was first delivered on the linac with the Gafchromic film sitting in the film cylindrical 61

64 phantom. At gantry 0 and collimator 0, the lower half of the phantom was placed on the couch and its geometrical centre, marked with the crosses, was aligned with the optical crosshairs. Then the front pointer was used to set the surface of the lower half-phantom to 100cm SAD. After the alignment, the Gafchromic EBT2 film was placed onto the phantom, its landscape orientation corresponded to the gantry target (G-T) direction of the linac, the yellow marker dye of the film at the bottom right corner of the phantom. The film was also aligned to the centre of the phantom (and thus the centre of the crosshairs) by the film positioning marks on the phantom. Once properly placed, the film was then fixed to the phantom using micro pore. The irradiation orientation G-T and left-right (L-R) was marked on the film for later scanning-position registration together with the centre cross-positions of the film. Then the upper half of the phantom was carefully added to the top of the film. All films were set up in exactly the same way to keep consistency, as this was highly recommended by the manufacturer to reduce possible systematic error. Once properly set up, the nominated IMRT case was loaded to the linac and delivered to the Gafchromic film. In addition to relative comparisons, absolute dose measurements were also required from the film. As a result, as with the ionisation chambers, a calibration factor was needed for the film to convert the readings into absolute doses. This reference film was taken using exactly the same setup as that of the ionisation chambers; i.e., Dmax water-equivalent build-up at 100cm SSD with enough backscatter provided, where 100MU corresponded to 100cGy. The reference film together with the IMRT-exposed film were then scanned exactly as described in section to keep consistency, and were read into the RODOMS software for analysis. 62

65 3.4 Error-Bar Determination To quantify the noise and the precision of the measurements, error bars were added to all the plots in the results (Chapter 4). How these error bars were calculated is explained as follows. In the experiment, measurements were repeated to increase the sample number and, consequently, improve the precision. The average of the readings was taken as the final result, while the standard deviation was used to quantify the noise. The standard deviation was calculated in Excel using the following equation: 2 2 Σ x ( Σx) Stdev =, ( n 1) where x was the average of all the repeated readings, and n was the total number of the readings. The % stdev was then calculated by: % stdev = (stdev / avg) x 100%. This percentage standard deviation was taken as the noise of the readings and was added to all the plots in part 4. 63

66 3.5 Gamma-Function Assessment The gamma function is a quantitative method of comparing measured distributions to the calculated values using two parameters. It is a widely used tool for IMRT QA. The gamma function provides a quantitative comparison of a) point doses between the physical film and the calculated distribution (the virtual film), and b) distance to nearest agreement (DTA) for the point in question. It assumes both the dose and the distance to be of equal importance, and allows the user to specify a tolerance level for each parameter. The function then returns a numerical value for the quality of the comparison result. Users are free to determine the tolerance limits. A gamma value of zero indicates perfect agreement between the two points while the agreement degrades as the gamma index converges towards one. Values exceeding one mean a failure to meet the pre-set tolerance. The equation of the gamma function is displayed below: r d ( ) ( ) Dd DD ϒ= In the equation, r is the distance-to-agreement (DTA), Δd is the DTA tolerance in mm set by the user, δ is the dose difference at a certain point and ΔD is the dose difference tolerance in % set by the user (Low, 1998). In the thesis the assessment values are 3% and 3mm, which have been used clinically and been supported by different authors (Low, 2003). This tolerance was chosen as a practical level used for analysis. There have been recently articles (Nelms et al., 2011), discussing the merits and limitations of gamma analysis. A further examination of this method is outside the scope of this thesis. It suffices to suggest that at this time gamma analysis is by far the most accepted analysis method in IMRT QA. Its strength lies in the way it quickly identifies regions of failure during QA analysis. In conclusion, the gamma-value function provides a direct comparison map between the virtual film and the physical film with both parameters considered, and is used as a pass/fail criteria for the IMRT QA using Gafchromic EBT2 film. 64

67 Chapter 4: Results and Discussions All the results are displayed and discussed in this chapter. Section 4.1 examines noise measurements; section 4.2 describes the calibration of Gafchromic EBT2 film using a novel method; section 4.3, 4.4 and 4.5 give the measurements for S cp output factors, PDD and OCR profile, respectively; and section 4.6 discusses the clinical IMRT QA. 4.1 Noise Measurements Ten repeated readings were taken for the CC01, the CC04, the CC13, the silicon diode and the Gafchromic film. The % standard deviation was used to represent the signal-to-noise ratio of the detectors. Both the 1cm x 1cm and the 10cm x 10cm field sizes were tested. Table gives the results of the averaged readings, the standard deviation and the % standard deviation relative to the average. 1x1 (nc) CC-01 CC-04 CC-13 Diode Film 10x10 (nc) 1x1 (nc) 10x10 (nc) 1x1 (nc) 10x10 (nc) 1x1 (nc) 10x10 (nc) 1x1 (GSV) 10x10 (GSV) Avg Stdev % Stdev Table 4.1.1: Readings for 10 repeated measurements at both field sizes To provide a more direct overview, the results were normalised to the same level and plotted on the same graph by normalising the average values to the same number through a conversion factor, then applying this conversion factor to the other readings. These normalised readings were then plotted on the same graph to illustrate the signal-to-noise ratio (Figure and 4.1.2). 65

68 Signal Variance at 1x1 Field Size Normalised Signal Exp No. CC-01 CC04 CC13 Diode Film Figure 4.1.1: Noise illustration at 1cm x 1cm field Signal Variance at 10x10 Field Size Normalised Signal Exp No. CC-01 CC04 CC13 Diode Film Figure 4.1.2: Noise illustration at 10cm x 10cm field The condition of the measurement was strictly controlled: the electrometer-detector system was fully warmed up; zero measurement was performed for each dosimeter before measurement; and the temperature, pressure and humidity were monitored closely and kept constant over the whole measurement. As a result, the fluctuation of the signal was most likely to come from the detector itself. From the above figures and graphs, some points could be noted: 1) The ionisation chambers and the diode had an acceptable signal-to-noise ratio. The noise was less than 1% of the primary signal. 2) The Gafchromic EBT2 film, on the other hand, had a much larger noise: 66

69 close to 3%, even when great care was taken when setting up and processing the films. Its precision was also relatively low compared to the other dosimeters (readings up to only one decimal place due to the limitation of the analysis software). Based on these facts, it was recommended that the Gafchromic film not be used for absolute-dose determination and commissioning-data collection, where high precision and high accuracy are important. 3) Noise was more significant at the smaller field among the ionisation chambers and the diode. This was because a) the intensity of the primary beam was reduced for small fields, and consequently a small fluctuation from the detector would introduce a larger error; and b) the systematic of the setup error was also more significant at smaller fields due to obvious reasons. 4) At the 10cm x 10cm field, the CC13 ionisation chamber had the smallest noise. This made sense, as the CC13 had the largest dimension among all the detectors (5.8mm length and 3.0mm radius), so its active volume, or sampling region, was also the largest, giving rise to a smallest noise. 5) At the 1cm x 1cm field, the CC13 ionisation chamber no longer had the smallest noise; this might have been caused by the signal perturbation from the volume constraint of the detector, as the length of CC13 was 5.8mm, which was very close to the field edge (1cm, or 10mm); thus the CC13 chamber might well be unable to fully collect the secondary scatters, or even the primary beam signal if some setup error was involved. On the other hand, the CC04 chamber had the smallest noise. This was because its dimension was much smaller than CC13 and very close to the CC01 and diode (3.6mm length for CC04 and CC01, and 2.5mm for diode), but its sampling volume was much larger than CC01 s and diode s volumes. 6) Although the signal-to-noise ratio assessed here helped reveal which detector gave the best precision, it was not the most important factor to determine the best detector overall. Although precision is not to be neglected in small field dosimetry, accuracy is a far more crucial factor; a detector with 3% precision but 1% accuracy is far preferable to one with a % precision but with 10% accuracy. The accuracy was assessed as discussed below. The signal-to-noise ratio could be quantified not only by the point-dose measurements, but also the curve scans, such as an OCR profile scan. Figure shows the magnified inter-umbra region of the 10cm x 10cm OCR curve measured in the experiment, clearly indicating the influence of noise on the quality of the scan. 67

70 CC01 CC04 CC13 Diode Figure 4.1.3: Magnified inter-umbra region of the OCR curves (scanned in water) with different dosimeters under a 10cm x 10cm field without the application of a smoothing function In the above graph, it was noted that the red curve measured by the CC01 ionisation chamber had the largest amount of fluctuation, which came from the noise of the detector. The CC13 ionisation chamber had the smallest fluctuation, as expected. It was concluded that a poor signal-to-noise ratio would greatly affect the quality of the curve scan, and this effect could not be neglected in determining a suitable detector for data collection. One could always argue that in curve scans, a smoothing function could be applied to eliminate the effect of the noise. It must be pointed out that with the existence of a poor signal-to-noise ratio, the application of a smooth function would very likely change the overall shape of the measured curve. In this case, the noise would worsen not only the precision of the measurement, but the accuracy. This influence is explained in more detail in section

71 4.2 Calibration of Gafchromic EBT2 Film Using a Novel Method The first part of this section assessed the dose response of the Gafchromic film when analysed in the red, green, blue or grey-scale channel. Figures to show the digitised images. Figure 4.2.1: 16-bit grey-scale value image (brightness: -70; contrast: 70) 69

72 Figure 4.2.2: 48-bit colour image Figure 4.2.3: Red channel image (extracted from the colour image) 70

73 Figure 4.2.4: Green channel image (extracted from the colour image) Figure 4.2.5: Blue channel image (extracted from the colour image) 71

74 The pixel values were then read from the centre of each strip; Table gives the results. All readings are background corrected. Dose (cgy) Red Chn Green Chn Blue Chn Greyscale Table 4.2.1: Dose response of Gafchromic film at different channels The values were then plotted in Excel (Figure 4.2.6). Gafchromic(tm) Film Dose Response - Different Channels y = x R² = Pixel Value Red Chn Green Chn Blue Chn Greyscale Linear (Greyscale) Dose (cgy) Figure 4.2.6: Gafchromic film PV-dose response with different channels Figure shows that, once processed in RODOMS, the 16-bit grey-scale image provides a much better linearity (R 2 = ~ ) than the colour-channel images within a dose range of 5Gy. Of the colour channels, the red channel had the 72

75 best absorption but still a significant sub-linearity. The blue channel could barely detect any signal under 5Gy. The superior linearity provided by the GSV image means that when it is used for a relative dose measurement only, no correction function is needed to convert the pixel value into the dose; this greatly improves the accuracy and efficiency of the measurement. Even when the Gafchromic film is used for an absolute measurement where dose information is needed, the GSV image s linear PV-dose response means that conversion to the optical density used in conventional methods can be avoided. This helps improve the accuracy of the results because the calibration fitting error is reduced with the conversion to the OD saved. As is stated in section 3.3, an image adjustment that includes brightness and contrast will also change the calibration curve. To illustrate this, Table compares the calibration curves from GSV images with different contrast and brightness corrections. Dose (cgy) Brightness = -70; Contrast = 70 Brightness = -60; Contrast = 60 Brightness = 0; Contrast = Table 4.2.2: Dose response of Gafchromic film with different image adjustments The results were plotted in Excel (Figure 4.2.7). 73

76 Gafchromic(tm) Film Dose Response - Different Image Adjustment Pixel Value R² = R² = R² = Dose (cgy) B=-70/C=70 B=-60/C=60 B=0/C=0 Linear (B=-70/C=70) Linear (B=-60/C=60) Linear (B=0/C=0) Figure 4.2.7: Gafchromic film GSV PV-dose response with different image adjustments Figure shows that with a combination of different brightness and contrast corrections, the calibration curve could vary significantly in both linearity and sensitivity. The dosimeter s sensitivity is also important for IMRT QA to measure sharp dose gradients. It can be seen that with a brightness of -70 and a contrast of 70, the calibration curve displays the best sensitivity (from 0 to 95.5) and the best linearity (R 2 closest to 1). These optimum contrast and brightness settings were used during analysis in subsequent film experiments discussed in this thesis. Once the general linearity of the GSV processing method was confirmed, the dose response was studied more carefully within three dose ranges. Figures to show the digitised images of the irradiated Gafchromic films for the three dose ranges. The dose delivered to each film strip is marked at the bottom right. 74

77 Figure 4.2.8: Low-dose-range film from 0cGy to 55cGy in steps of 5cGy Figure 4.2.9: Middle-dose-range film from 0cGy to 550cGy in steps of 50cGy 75

78 Figure : High-dose-range film from 0cGy to 1100cGy in steps of 100cGy The GSV were read in RODOMS at the centre of each strip and recorded (Table 4.2.3). Low-dose-range film Middle-dose-range film High-dose-range film Dose (cgy) GSV Dose (cgy) GSV Dose (cgy) GSV Table 4.2.3: Dose response of Gafchromic EBT2 film with different ranges 76

79 The GSV was then plotted against the dose for the three dose ranges, using Excel (Figure to ). A linear trend line was added to the data set and the R 2 value was shown on the graph as a check of the linearity. Film Dose Response in Low Dose Range 40 R 2 = GSV cgy Figure : Film dose response in low-dose range Film Dose Response in Middle Dose Range R 2 = GSV cgy Figure : Film dose response in middle-dose range 77

80 Film Dose Response in High Dose Range R 2 = GSV cgy Figure : Film dose response in high-dose range A linear fit was achieved with the R 2 values shown on each graph in all three dose ranges. The best linearity was obtained in the low-dose range, which was very encouraging, as it meant that the Gafchromic film could measure all these small-segment doses of the IMRT without a correction. Even with the lowest R 2 value at the high-dose range, the linearity was still guaranteed within 1.5% (R 2 = = 98.65% > 98.5%). The dose ranges were then combined below to generate an overview of the dose response of the Gafchromic EBT2 film (Figure ). 78

81 GafchromicTM EBT2 Film PV-Dose Response y = -5E-05x x R² = GSV Pixel Value Dose (cgy) Figure : Overall dose response of the Gafchromic film, in a range from 0cGy to 1100cGy Figure shows that the overall GSV-to-dose response is not so linear in a dose range of 0 to 11Gy. The best fit of the data is achieved at a second-power polynomial (y = -5E-05x x with R 2 = ). The curve tends to go flat after 6Gy, and actually becomes saturated at 15Gy. However, within 0 to 7Gy the GSV-to-dose relationship is still rough linear; this means that the GSV pixel value is still suitable to be used for IMRT QA, for which the prescription dose per fraction is usually in a range of 1.5Gy to 5.0Gy. The linearity in the dose region below 1Gy is also of great importance because this is usually the dose received by the OAR in an IMRT plan. In conclusion, when scanned as a 16-bit grey-scale-channel image using the optimum contrast and brightness calibrated scan setup, the Gafchromic EBT2 film was considered suitable for use in IMRT QA. Because most IMRT tests range in dose between 0 and ~3Gy, a linear GSV-to-dose calibration curve could be implemented. 79

82 4.3 Output-Factor Measurements Table gives the output factors measured by the different ionisation chambers and the diode. MLC Field CC01 CC04 CC13 Diode Size (cm) Table 4.3.1: Output factors measured by different ionisation chambers and diode Measurements with the Gafchromic EBT2 film were repeated five times to improve precision. Figure shows one of the digitised films. Figure 4.3.1: JPEG image of film output factors, irradiated on 27/7/10 (field size marked on the bottom right of each strip) 80

83 The GSV values were obtained from the centre of each film stripe, with the background value corrected by subtraction in RODOMS. The maximum GSV value recorded was in Table As discussed in section 4.2, this GSV corresponded to a dose of about 600cGy, where the linearity of the film s dose response was still maintained. Thus the ratio of the GSV readings could be taken directly as the output factors without being converted to the dose. All five sets of the calculated output factors were put into the same table, with the average and the standard deviation calculated. The averaged output factor was taken as the final result, while the standard deviation as the error bar (Table 4.3.2). MLC Field Size (cm) OF1 OF2 OF3 OF4 OF5 Avg STDEV % STDEV Table 4.3.2: Average output factors and standard deviations measured with Gafchromic film All output factors were tabulated in Table and plotted by Excel in Figure FS (cm) CC01 CC04 CC13 Diode Film Table 4.3.3: Output factors measured with different detectors 81

84 Output Factor Scp Output Factor cc-01 cc-04 cc-13 Diode Film Field Size (cm x cm) Figure 4.3.2: Output-factor curves measured with different dosimeters; errors quoted in the previous tables are left off in this graph for clarity Some points were noted from the plot above: 1) The data were normalised so that all five curves intersected at 1.00 at 10cm x 10cm field size, which was the reference reading for the output-factor measurement. 2) The curve measured with the Gafchromic film had the largest discrepancy from the others. Firstly, the Gafchromic film had the highest spatial resolution. Rather than collecting the ion-pairs produced by ionisation, the Gafchromic film measured the dose based on its radiochromic layer, which produced a very fine-grained blue colour upon irradiation due to a solid-state polymerisation process. This polymerisation process could happen at any single pixel point of the layer based on the income radiation. As a result, the spatial resolution of the film was much higher than that of the ionisation chambers and the diode. Since in this experiment all the films were scanned at 100dpi, the size for each dot was only 0.01inch, or 0.254mm; consequently, the scanner resolution was 0.254mm, while the smallest dimension of the chambers was 1mm (CC01). Thus theoretically the Gafchromic film had the highest spatial resolution, and its curve was possibly closest to the true value for a zero-volume detector response. 82

85 On the other hand, the measurements with the Gafchromic film were very unreliable. The major concern was the poor signal-to-noise ratio of the film. The noise was so significant that despite its high spatial resolution, the Gafchromic EBT2 film was concluded to be unsuitable for commissioning measurements where both high accuracy and high precision were compulsory. 3) The curves of the ionisation chambers and the diode agreed quite well after 5cm x 5cm, but not in the smaller fields. To view the differences better, the output factors for small fields were plotted separately for the ionisation chambers and the diode (Figure 4.3.3). Output Factor for Different Detectors in Small Fields 1.05 Output Factor cc-01 cc-04 cc-13 diode Field Size (cm) Figure 4.3.3: Output-factor curves measured with different ionisation chambers and diode, normalised to the 4cm x 4cm field reading The largest discrepancy occurred at the 1cm x 1cm field. At this field size, the diode had the highest output factor, followed by the CC04, the CC01 and the CC13. The reason why CC13 had the lowest output factor at the smallest field was very obvious: the volume-averaging effects. The CC13 ionisation chamber had the largest volume, with a radius of 3.0mm and a length of 5.8mm, which made its measurement volume the largest. However, at very small fields such as 1cm x 1cm, this measurement volume could not be fully covered by the primary beam, leaving the CC13 to collect a lower signal than the true signal i.e., the volume-averaging effect which was most significant for the CC13 ionisation chamber. On the other hand, the diode had the least side length (2.5mm compared to 3.6mm for CC01 and CC04, and 5.8mm for CC13); its output factor at 1cm x 1cm was the highest because it displayed the least volume-averaging effect. 83

86 4) Although the diode was expected to have a higher output factor than the others because of the volume-averaging effect, the difference was not expected to be so high (nearly 5% higher than CC04 in the plot). This could possibly originate from the diode overestimating the dose from low energy photons due to its non-tissue equivalence and higher absorption coefficient than water (Bucciolini et al., 2003), and because at very small fields the beam includes scattered components whose energy is relatively low. The Gafchromic EBT2 film might have the truest curve, but its poor signal-to-noise ratio not only lowered the precision but also influenced the accuracy. The measurements from the ionisation chambers were all severely restricted by the volume-averaging effect, although among them the CC04 had the best curve, as the CC01 was also influenced by the stem effect. On the other hand, the diode s lack of lateral electron equilibrium due to a higher material density would reduce its accuracy in small-field measurement. It was noted, though, that because of the different performances of the dosimeters, it would be good if the quality assurance dosimeter was kept the same as that used for commissioning so that the variance in the dosimeters would not introduce unnecessary systematic error. 84

87 4.4 Percentage Depth Dose Measurements Figure shows the PDD curve of the 1cm x 1cm field, measured by different dosimeters and re-plotted in Excel. PDD at 1cm x 1cm Field for Different Detectors 100 Relative Dose (%) cc-01 cc-04 cc-13 Diode Depth (mm) Figure 4.4.1: PDD curves measured with different dosimeters under the 1cm x 1cm field To see the difference better, the whole curve was split into two parts: the build-up region and the fall-off region, with the intersect point being the dmax depth. Figure shows the magnified build-up region of the PDD curves. 85

88 Build-up Region of the PDD Curve for Different Detectors Relative Dose (%) Depth (mm) cc-01 cc-04 cc-13 Diode Figure 4.4.2: Build-up region of the PDD curves measured with different dosimeters under the 1cm x 1cm field For high-energy photon beams, the point of maximum dose lies deeper into the tissue or phantom, giving rise to the dose build-up region. As the high-energy photon beam enters the patient or the phantom, high-speed electrons are ejected from the surface and the subsequent layers. These electrons deposit their energy a significant distance away from their site of origin. Hence the electron fluence and the absorbed dose increase with depth until they reach a maximum, then start decreasing with depth (Khan, 1992). Historically it has been difficult to measure the dose at the build-up region accurately, and different detectors usually show large discrepancies (Alves et al., 2001). In the build-up region, little difference could be seen among the curves of the CC04, the CC13 and the diode. On the other hand, the CC01 showed a lower response than the other detectors in this region. When scanning in the depth direction, the dimension that would affect the measurement resolution was mainly the radius of the dosimeters. As a result, the CC01 was expected to have the highest spatial resolution (1mm radius), the CC04 and the diode to have a similar spatial resolution (both 2mm radius), and the CC13 to have the lowest spatial resolution (3mm radius). The reason why the curve measured with the CC01 was steeper than the others might be one of the following: 1) All the curves were normalised to the same level (100%) by the maximum dose value, so if the peak value measured by one dosimeter was higher than 86

89 the others, after normalisation the total curve would be shallower. This was highly possible for the CC01 ionisation chamber, as its volume-averaging effect was the smallest, rendering the highest peak value and consequently the shallowest build-up region. For 6MV photon peaks the effect should not be significant, as the peak was not very sharp; however this could be a big concern to small-field profiles where the peak was very narrow. 2) It could also come from the electron fluence perturbation, which was caused by the electron fluence from the air cavity in the ionisation chamber. The effect of the electron fluence perturbation depended on the size of the chamber, but for high-energy photons this effect was usually very small. 3) Another possible reason for this dose variation at shallow depth might be errors in the setup. If the effective point of the chamber was not aligned perfectly with the surface of water, an extra build-up of the chamber might be introduced. Thus at the 0 depth indicated on the graph, the detector might not have measured the surface dose but a dose at deeper depth. This effect depended totally on the setup of each chamber, consequently causing the difference in the surface dose. Water-surface perturbation might bring in the same effect. 4) Although parallel plate chambers and solid water phantom are more ideal for exact surface-dose measurements (Javeden, 2010), near the surface the visualisation with isodose curves in such high-dose gradients is difficult. Hence, while it is well known that cylindrical data overestimates surface dose, it is generally more practical for deployment into the RTPS, as differences cannot be easily visualised. While combining data with parallel-plate ionisation chamber data may be more accurate, this is not a common practice. The curve measured by the CC01 was assumed to be the most accurate curve because of its highest spatial resolution. Other detectors could not provide as high spatial resolution due to the volume-averaging effect. The diode also had silicon perturbation as described in the previous section, but it is assumed this was eliminated by normalisation. For the current planning protocols adapted in the author s hospital, the use of the build-up region in the treatment of a patient was avoided, especially in the IMRT treatments. Treatments close to the skin would use lower X-ray energies or electrons with or without a bolus. The concern was exactly what was encountered in this study: that the build-up region of a beam could not be measured accurately enough for clinical assessment. The part of the PDD curve after the Dmax point, or the fall-off region, where electron contamination was negligible, is of more clinical importance for treatment planning. Figure shows the PDD fall-off region of the 1cm x 1cm field. 87

90 Fall-off Region of PDD for Different Detectors Relative Dose (%) Depth (mm) cc-01 cc-04 cc-13 Diode Figure 4.4.3: Dose fall-off region of the PDD curves measured with different dosimeters under the 1cm x 1cm field It was very clear from the above plot that no difference could ever be observed in the photon PDD fall-off region. At least for the 1cm x 1cm field, it appeared that the PDD measurement difference only existed in the build-up region. To see how the detectors acted in measuring the PDD curve with a larger field, the 10cm x 10cm field was used (Figure 4.4.4). PDD at 10cm x 10cm Field for Different Detectors 100 Relative Dose (%) cc-01 cc-04 cc-13 Diode Depth (mm) Figure 4.4.4: PDD curves measured with different dosimeters under the 10cm x 10cm field. 88

91 Again, for better detail, the curve was divided: the build-up region (Figure 4.4.5) and the fall-off region (Figure 4.4.6). Build-up Region of the PDD Curve at 10cm Field for Different Detectors Relative Dose (%) Depth (mm) cc-01 cc-04 cc-13 Diode Figure 4.4.5: Build-up region of the PDD curves measured with different dosimeters under the 10cm x 10cm field A similar difference was seen between the CC01 and all the other detectors in the build-up region of the 10cm x 10cm PDD; this might again be due to the reasons described above. No difference was seen among the CC04, the CC13 and the diode. 89

92 Fall-off Region of the PDD Curve at 10cm Field for Different Detectors Relative Dose (%) Depth (mm) cc-01 cc-04 cc-13 Diode Figure 4.4.6: Dose fall-off region of the PDD curves measured with different dosimeters under the 10cm x 10cm field Again, no difference could be seen for all the detectors in the fall-off region. In conclusion, in the PDD measurement, all the tested dosimeters collected quite similar results. The only difference was that the CC01 ionisation chamber had a lower response than the other detectors in the build-up region of the PDD curve, possibly because of the volume-averaging effect of the peak, or the electron fluence perturbation. Besides this, all the detectors agreed well in the fall-off region of the PDD curve. Since the treatment using the build-up region of the photon beam was avoided in the author s hospital, the CC01 s difference in the build-up region was assumed negligible. All the detectors that were tested the CC01, the CC04, the CC13 and the diode were concluded suitable for the measurement of the small-field PDD curve, as they all provided adequate accuracy in this part. 90

93 4.5 Off-Central-Axis Profile Measurements Figure shows the overall OCR curves measured by different dosimeters under the 1cm x 1 cm field at 10cm depth in water. OCR at 1cm x 1cm for Different Detectors Relative Dose (%) Off-axis Distance (mm) cc-01 cc-04 cc-13 Diode Figure 4.5.1: OCR curves measured with different dosimeters under the 1cm x 1cm field For better detail, the left-hand penumbra region was magnified for comparison (Figure 4.5.2). Penumbra Region of OCR for Different Detectors 100 Relative Dose (%) Off-Central Distance (mm) cc-01 cc-13 Diode cc-04 Figure 4.5.2: Penumbra region of the OCR curves measured with different dosimeters under the 1cm x 1cm field 91

94 Once the penumbra region was magnified, it was seen that no curves were similar, which indicated that the four detectors measured the profiles differently due to their characteristics. Differences were mostly in the curve shape of the penumbra and the level of the penumbra tail. To give a better illustration, Table shows the curves values for d 20 (depth of the 20% dose) and d 80 (depth of the 80% dose). Dosimeter Type d 80 (mm) d 20 (mm) d 80 d 20 (mm) cc cc cc diode Table 4.5.1: d 80 and d 20 values of the 1cm x 1cm OCR curves collected by different dosimeters First, the curve shape of the penumbra, or the gradient, was different. From Figure and Table 4.5.1, it can be seen that the diode s curve had the largest drop-off in the penumbra region, the CC01s the second, the CC04s the third, while the CC13s showed the flattest curve, as the smaller the d 80 -d 20 value is, the sharper the curve is. This was caused by the volume-averaging effect of the dosimeters. For a profile measurement, besides the detector s radius, the side length of the detector would also affect the spatial resolution of the scan. Consequently the diode was expected to have the highest resolution because its side length (2.5mm) was the smallest among all the detectors, followed by the CC01 (3.6mm and a radius of 1.0mm), the CC04 (3.6mm and a radius of 2.0mm), and the CC13 (5.8mm and a radius of 3.0mm).For the profile of a very small field, such as 1cm x 1cm, the OCR curve is mainly composed of the penumbra, where the most contribution comes from scatters of the primary beam (including scattered component of primary photons from the linac head and backscattered from the patient, and photon-electron interactions) and the dose changes rapidly over a short distance. Thus, the spatial resolution of the detector is very critical to the accuracy of the measurement. As a result, in this study the curve from the diode had the largest drop-off gradient and was concluded to be the most accurate, followed by the CC01, the CC04 and finally the CC13, which gave a much flatter curve than the others because of its much larger dimension. Second, the level of the penumbra tail was also different among the four detectors. As clearly demonstrated by the d 20 values in Table 4.5.1, the CC01 and the diode had a lower-dose penumbra tail, while the CC04 and the CC13 had a higher and flatter tail. This was also caused by the volume-averaging effect of the dosimeters. In the profile measurement, the curve was always normalised to the peak/maximum dose value; thus the measurement accuracy of this normalisation point was of particular importance. For the 1cm x 1cm OCR curve, there was nearly no central region, which is defined as the portion from the beam central axis to within 1 to 1.5cm from the geometric field edges of the beam (Podgorsak, 2005); the largest contribution was from the penumbra proportion, as stated above. As a result, measuring the dose 92

95 maximum demanded a very high resolution from the detector, as the dose maximum, or the central region, was very small, and if the detector was not small enough, the uniformity over the sensitive volume could not be ensured. Because of their small sizes, the diode and the CC01 measured this dose maximum more accurately. On the other hand, the dose maximum measured by the CC04 and the CC13 was lower because their large volumes meant that their sampling region also included part of the penumbra, thus decreasing the measured Dmax. All the curves were normalised to 100% at the measured dose maximum, so the CC04 and the CC13 had a higher tail than the diode and the CC01; this was actually because their measured Dmax was lower or less accurate. Figure shows the plots for the OCR profiles of the 10cm x 10cm field measured at 10cm depth in water. OCR at 10cm x 10cm Field For Different Detectors 100 Relative Dose (%) cc-01 cc-04 cc-13 Diode Off-Central Distance (mm) Figure 4.5.3: OCR curves measured with different dosimeters under the 10cm x 10cm field The left side of the OCR curve was magnified and plotted in Figure to better show the differences. 93

96 Penumbra Region of OCR at 10cm Field for Different Detectors FWHM 100 Relative Dose (%) Off-Central Distance (mm) cc-01 cc-04 cc-13 Diode Figure 4.5.4: Penumbra region of the OCR curves measured with different dosimeters under the 10cm x 10cm field Table tabulates the values of d 20 and d 80 for analysis. Dosimeter Type d 80 (mm) d 20 (mm) d 80 d 20 (mm) CC CC CC Diode Table 4.5.2: d 80 and d 20 values of the 10cm x 10cm OCR curves collected by different dosimeters Figure and Table showed that the gradient of the shoulders of the OCR curves demonstrated a similar difference to that seen in the small field, again because of the volume averaging of the dosimeters. The diode had the smallest penumbra, while the CC13 had the largest. The d 80 -d 20 difference among the diode, CC01 and CC04 was not significant. On the other hand, the difference in d 20 values in the table was very small, demonstrating a similar dose percentage in the tail region below the 20% level: because for a 10cm x 10cm profile curve, the central region was large (about 8cm wide) and flat, so the accurate measurement of the dose maximum for normalisation did not demand high resolution, and even a CC13 ionisation chamber was adequate in this region. It is also possible to derive virtual-zero detector volume profiles based on extrapolation of data from multiple detector diameters (Yuen, 2009, Laub, 2003, Metcalfe, 1993). In this study, the d 80 -d 20 measurements were extrapolated to obtain 94

97 the zero-detector value. According to Yuen, extrapolation of penumbra with detector size was found to have an invalid linearity in the region of low detector size, so the extrapolation performed here excluded the data measured by the diode. Another reason for not including the diode was that, unlike the ionisation chambers, it did not have a cylindrical geometry. The extrapolations for the 1cm x 1cm and the 10cm x 10cm are shown in Figures and cm x 1cm Penumbra Extrapolation 5 y = 0.375x d80-d20 (mm) Detector Diameter (mm) Figure 4.5.5: Virtual zero-volume detector profile extrapolation for 1cm x 1cm field 95

98 10cm x 10cm Penumbra Extrapolation 7 6 y = 0.350x d80-d20 (mm) Detector Diameter (mm) Figure 4.5.6: Virtual zero-volume detector profile extrapolation for 10cm x 10cm field The d 80 -d 20 value measured by a virtual zero-volume detector, or the true d 80 -d 20 value, can be determined by the intercept of the linear extrapolation. Figures and show that this intercept was 2.63mm for the 1cm x 1cm profile, and 4.40mm for the 10cm x 10cm profile. The closest measurement was given by the diode, then the CC01, CC04, and CC13. These results were also very close to the values reported in the literature. In conclusion, for the profile measurement, the choice of the detector would greatly influence the accuracy of the measurement. High-resolution detectors were required to ensure accuracy, as discussed above. From the comparison, it was concluded that the diode was the most suitable detector for the OCR profile measurement, because it provided a very high spatial resolution, and was not affected by electron fluence perturbation. Among the ionisation chambers, the CC01 could also be used, but it was still influenced by volume averaging as well as electron fluence perturbation. The CC04 and the CC13 were not good choices because they could not provide high spatial resolution due to their larger sizes; this compromised the accuracy. However, it was possible to obtain a profile measured by a virtual zero-volume detector using multiple measurements and extrapolation method (Yuen, 2009). 96

99 4.6 Clinical IMRT QA Comparison In the last part of the experiment, the IMRT plan was used for comparison. The ionisation chambers were used to measure a single-point dose, while the Gafchromic film was used to obtain the dose distribution of a plane. The results were then compared to the planning-system output at the same position, which was taken as the standard. Section describes how the planning standard was generated Planning Standard Output The dose at the selected interest point was calculated in XiO and exported (Table ). Beam No. Gantry Angle ( ) Dose (Gy) Total Table : Doses of interest point exported from XiO 97

100 The dose distribution of the selected interest plane was calculated and exported from the XiO. It was then imported into RODOMS and converted to a GSV map. Figure shows a screen capture of this virtual film. Figure : Virtual film from XiO 98

101 4.6.2 Ionisation Chamber Measurement This section lists the point dose measurements from the ionisation chambers, describes the process of converting the readings to the dose, compares the results to the planning standard and provides an analysis. (a) CC13 First, the reference readings for the CC13 ionisation chamber were taken. For 100cGy, the results were: 3.651nC, 3.650nC, 3.651nC, 3.651nC and 3.652nC. The average was 3.651nC. The conversion factor for the CC13 was then calculated as: kc = 3.651nC / 100cGy = 3.651nC/Gy. The IMRT plan was delivered to the CC13, and the readings for each beam were recorded and converted to doses using kc. The doses were then compared to the planning standard (Table ). Beam No. Gantry Angle ( ) Reading (nc) Measured Dose (cgy) Planning Dose (cgy) Difference (%) Total Table : IMRT point-dose comparison CC13 99

102 (b) CC04 Next, the reference readings for the CC04 ionisation chamber were taken. For 100cGy, the results were: 1.020nC, 1.020nC, 1.020nC, 1.020nC and 1.020nC. The average was 1.020nC. The conversion factor for the CC04 was then calculated as: kc = 1.020nC / 100cGy = 1.020nC/Gy. The IMRT plan was delivered to the CC04, and the readings for each beam were recorded and converted to doses using kc. The doses were then compared to the planning standard (Table ). Beam No. Gantry Angle ( ) Reading (nc) Measured Dose (cgy) Planning Dose (cgy) Difference (%) Total Table : IMRT point-dose comparison CC04 100

103 (c) CC01 Finally, the reference readings for the CC01 ionisation chamber were taken. For 100cGy, the results were: 0.316nC, 0.315nC, 0.315nC, 0.314nC and 0.315nC. The average was 0.315nC. The conversion factor for the CC01 was then calculated as: kc = 0.315nC / 100cGy = 0.315nC/Gy. The IMRT plan was delivered to the CC01, and the readings for each beam were recorded and converted to doses using kc. The doses were then compared to the planning standard (Table ). Beam No. Gantry Angle ( ) Reading (nc) Measured Dose (cgy) Planning Dose (cgy) Difference (%) Total Table : IMRT point-dose comparison CC01 101

104 For better comparison, the minimum discrepancy, the maximum discrepancy and the total difference of the three ionisation chambers were tabulated in Table Ionisation CC01 CC04 CC13 Chamber Max Diff (%) Min Diff (%) Total Diff (%) Table : IMRT point-dose comparison summary Tables to showed that the chambers measurements agreed well with the planning system from beam 1 to beam 3, but did not agree as well at beam 4. After looking into the plan, it was found that although the chosen point of interest sat in a relatively uniform dose region with the sum of all four beams, it was on the dose gradient of beam 4 individually. Thus, due to their volume limitations, the ionisation chambers measured a reading with a much higher discrepancy at beam 4. However, for all the other beams and the total dose, the measurements agreed quite well with the planning outcome, with a difference of less than 3.0%. When beam 4 was included in the analysis, the CC13 ionisation chamber measured the largest maximum difference and the largest total difference. This was caused mostly by the volume-averaging effect of the ionisation chamber. In the RTPS, the number of calculation points is determined by the calculation grid size chosen by the user. In this thesis, for the best accuracy, the author chose the grid size to be the smallest available size, which was 0.01mm x 0.01mm x 0.01mm. This grid size is much smaller than the active volume of the detectors. On the other hand, in an ionisation chamber, although the dose at a point (effective point of measurement) is measured, it is calculated by measuring the charge produced in a volume of air surrounding that point. The resulting volume-averaging effect occurs, because the energy deposition in the medium varies spatially. Because the CC13 ionisation chamber had the largest volume, its volume-averaging effect was the most significant, contributing to the largest maximum dose difference and total dose difference. The result could be even worse if the interest point was not chosen to be at a uniform dose region but at a sharp dose gradient, as was seen from the results of beam 4. Based on the data in Table , it was concluded that with its medium volume, the CC04 achieved a good balance between accuracy and reliability, as it had the lowest maximum dose difference, minimum dose difference and total difference. Table shows the statistics when beam 4 was excluded from the analysis. 102

105 Ionisation CC01 CC04 CC13 Chamber Max Diff (%) Min Diff (%) Total Diff (%) Table : IMRT point-dose comparison beam 4 excluded Considering the volume-averaging effect, the CC01 was expected to have smaller maximum, minimum and total differences to the CC04 and the CC13, as it had a smaller volume, and thus higher resolution and, in turn, greater accuracy. However, the results showed the contrary: the CC01 had the largest discrepancy in all three parameters, while the CC04 and CC13 showed a similar difference. The author concluded that this would most likely come from the systematic error of the setup. When setting up the cylindrical phantom, a misalignment of the chamber could happen due to operator error. In this case the extremely small volume of the CC01 was a disadvantage rather than an advantage, as it made the setup much harder. What s more, the same degree of misalignment would cause a larger error with the CC01 than with the CC04 and the CC13. For example, a 0.1mm misalignment would make the CC01 totally offset from the target interest point, while with the same amount of the misalignment the CC04 and the CC13 would still cover the target interest point. Although the CC01 ionisation chamber could possibly provide the highest accuracy because of its smallest volume, it was not always reliable due to the setup difficulty. On the other hand, with the largest volume, the CC13 most likely offset the small systematic error, leading to the smallest discrepancy among the three. In conclusion, limited by volume averaging, the ionisation chambers could not provide very high accuracy in the point-dose comparison for IMRT QA. A single-point-dose comparison was also not adequate for the IMRT QA; a dose-distribution comparison would be more helpful, but it could not be achieved using the ionisation chambers. However, the ionisation chambers could still be used as a secondary IMRT QA tool if combined with other dosimeters such as the Gafchromic film. Their advantage is that they can accurately convert the electrometer reading into the absolute dose, thus supplying an absolute-dose comparison, which is useful for the QA. The best ionisation chamber used in these experiments for IMRT QA was concluded to be the CC04 ionisation chamber. Its volume was small enough to provide a reasonably high spatial resolution, and thus accuracy, but not too small to be severely influenced by possible systematic errors that could significantly reduce the accuracy. 103

106 4.6.3 Gafchromic EBT2 Film Measurements The Gafchromic EBT2 film was used to perform the dose-distribution comparison. Figure and show the digitised IMRT measurement film and the reference film, respectively. Figure : IMRT Gafchromic film 104

107 Figure : Reference Gafchromic film 105

108 The IMRT dose-distribution comparison using Gafchromic film could be performed in either the relative or the absolute mode. In the relative mode, the GSV value of the IMRT film are directly normalised (or scaled) to that of the virtual film based on a normalisation point specified before comparison. The normalisation point is usually the point at which the ionisation chamber measurement is performed, and should be within the low-gradient, high-uniform dose volume to reduce the potential uncertainty. However, re-normalisation to a relative dose point is not the best solution, as errors from this normalisation point will be inherited by all the other measurement points, and the choice of this normalisation point cannot be reproduced easily, introducing significant uncertainty. Thus, in this study, an absolute-dose measurement was compulsory as well. In the absolute mode, the reference Gafchromic film was used. A conversion factor from the GSV to the dose was achieved from the reference film whose irradiated dose was known. This conversion factor was then applied to the IMRT film to convert its GSV to absolute dose. The IMRT film s absolute dose was then compared to the planning results point by point. The QA analyses in both modes were performed in RODOMS. First, the relative mode comparison was performed. Before analysis, the digitised IMRT film was re-scaled and co-registered to the virtual film. Then a background point was chosen where the film was not irradiated, and the value of this point subtracted as the background. Finally, the normalisation point was chosen (in this case the iso-centre), which brought the IMRT film to the same level as the virtual film. Then different comparison or analysis could be performed in RODOMS. Figure shows a screen capture. 106

109 Figure : Screenshot of the relative-dose analysis mode of RODOMS. (1): Virtual film with iso-dose lines; (2) Digitised IMRT film with iso-dose lines; (3) Central axis line profile (horizontal profile, left; vertical profile, right); (4) Point-dose comparison in relative mode. Three comparisons were made: a) the iso-dose line comparison; b) the dose-profile comparison; and c) the point-dose comparison in relative mode. The isodose line comparison gave an overview of the dose-distribution comparison by connecting the dots that had the same GSV value. Different colours indicated different isodose levels (90%, 80%, 50% and 20%). Figure shows that the isodose line distribution was similar in the IMRT film to that in the virtual film. In the dose-profile comparison, the dose profiles along both the horizontal and the vertical lines of a cross were plotted from both films and compared. The centre of the cross was selected by the user and could be moved to any position in the plane; thus the dose profile could be compared at any point. In item 3 in Figure , the horizontal dose-profile comparison is on the left, and the vertical dose-profile comparison is on the right. The red line is the profile from the virtual 107

110 film, while the blue line is the profile from the actual film. No obvious difference could be observed, which indicated good agreement between the measurement and the planning goal. The point-dose comparison in relative mode also provided a point-to-point comparison. There was a major advantage in using the Gafchromic film for the point-dose comparison: while a single ionisation-chamber measurement could only supply a single-point comparison, a single film could supply multiple-point comparison, and the point could be chosen at any position in the plane. Table gives the results. Point Coordinate X Coordinate Y Planning %Dose Film %Dose % Dev Average 2.07 Table : Point-dose comparison in relative mode From Table , it was noted that the total dose difference at the iso-centre (point 8, position (0, 0)) measured by the film was 1.48%, which was close to the results from the ionisation chambers (1.20% for CC04). Generally, the points at different positions agreed quite well, with a maximum difference of 5.00% (point 6) and an average difference of 2.07%. However, it must be pointed out that in the relative mode, the point comparison was not based on the absolute dose, but on the GSV values from the virtual film and the physical film at a selected normalisation point. The error in this normalisation point was then inherited by all the other points, accounting for the general QA result. Although carefully choosing this normalisation point to be in a uniform and high-intensity dose region could potentially reduce this error, it can never be fully eliminated. The purpose of the relative point-dose comparison is only to provide a quick overview of the QA. Consequently, using a calibration film to do an absolute point-dose comparison is necessary. In absolute mode, the reference film was used and the GSV from the measurement was converted to dose before comparison. It showed a very close result, with 1.60% difference at the iso-centre and an average difference of 3.47%. In general, if the systematic error is reduced as much as possible in the 108

111 setup and the processing of both the reference film and the IMRT film, the point-dose comparison results should not show any disagreement in either relative or absolute mode. In the absolute mode only the point-dose comparison could be conducted, but not other functions available in the relative mode, so the result was not listed here. Finally, a gamma-function assessment was conducted in RODOMS in the relative mode, as explained in section 3.5. The tolerance for the gamma function was set at 3% and 3mm. A new green-red colour image was generated based on the gamma results (Figure ). In this image, green indicates a pass result and red indicates a fail Figure : Screenshot of the gamma-value function result from RODOMS. (1): Virtual film; (2) Physical film; (3) Gamma-value tolerance limits set by the user (3mm and 3%); (4) Green-red colour image based on the results of the gamma-value function. As seen from the gamma-function image result, the measurement from the Gafchromic film had a very good agreement with the planning output. Only at the edge of the PTV was there a small region of red, or failure to meet the 3%-3mm criteria. To demonstrate how accurate and sensitive the gamma-function assessment was, the measurement film was shifted to the left by 109

112 0.5mm in the software; then the gamma function was re-assessed (Figure ). Figure : Gamma-function result with the measurement film shifted to the left by 0.5mm The influence of the movement was is plain: the right side of the image, which had shown good agreement before, totally failed the tolerance after the shift. Since the movement made only 0.5mm, the accuracy and the sensitivity of the RODOMS gamma function was well assessed. When the image was shifted in the same direction by 1.0mm, the effect was even more obvious (Figure ). 110